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Текст
Advanced Dental Biomaterials
WOODHEAD PUBLISHING SERIES IN BIOMATERIALS Advanced Dental Biomaterials Edited by Zohaib Khurshid Shariq Najeeb Muhammad Sohail Zafar Farshid Sefat
Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom Copyright © 2019 Elsevier Ltd. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-08-102476-8 (print) ISBN: 978-0-08-102477-5 (online) For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals Publisher: Matthew Deans Acquisition Editor: Sabrina Webber Editorial Project Manager: Joshua Mearns Production Project Manager: Debasish Ghosh Cover Designer: Greg Harris. Typeset by MPS Limited, Chennai, India
List of contributors Mohamed-Nur Abdallah Faculty Toronto, ON, Canada of Dentistry, University of Toronto, Azeem Ajaz Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Kingdom of Saudi Arabia Mai Saleh Ali Faculty of Dentistry, University of Toronto, Toronto, ON, Canada; Private Practice, Amman, Jordan Saqib Ali Department of Biomedical Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia Sanam Almassi Almassi Specialist Clinic, Tehran, Iran Ahmad A. Alnazzawi Department of Substitutive Dental Sciences, College of Dentistry, Taibah University, Medina, Saudi Arabia Mothanna Alrahabi Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia Abdullah Alwadaani Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia Marina Amaral Department of Dentistry, University of Taubaté, Taubaté, Brazil Faiza Amin Department of Science of Dental Materials, Dow Dental College, Dow University of Health Sciences, Karachi, Pakistan Paul Anderson Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom Sukumaran Anil Department of Dentistry, Hamad Medical Corporation, Doha, Qatar Anderson Catelan Faculty of Health Sciences, University of Western São Paulo, Presidente Prudente, Brazil
xxii List of contributors Elna Paul Chalisserry Interdisciplinary Program of Marine-Biomedical, Electrical and Mechanical Engineering, Center for Marine-Integrated Biomedical Technology (BK21 Plus), Pukyong National University, Busan, South Korea Amritpaul Singh Dhillon Dental Institute, King’s College London, London, United Kingdom Sergey V. Dorozhkin Moscow, Russia Tomas Duminis Centre for Oral Bioengineering, Institute of Dentistry, Barts and the London School of Medicine and Dentistry, London, United Kingdom Daghigh Ahmadi Ehsaneh Centre for Nanohealth, College of Engineering, Swansea University, Swansea, United Kingdom Ahmed El-Banna Dental Biomaterials Department, Faculty of Dentistry, AinShams University, Cairo, Egypt Muhammad A. Fareed Adult Restorative Dentistry, Dental Biomaterials and Prosthodontics Oman Dental College, Muscat, Sultanate of Oman Azita Farhadi Shamsabadi Centre for English Language Education, Nottingham University, Nottingham, United Kingdom Imran Farooq Department of Biomedical Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia Amr S. Fawzy UWA Dental School, University of Western Australia, Nedlands, WA, Australia Pegah Firouzmanesh Babol University of Medical Sciences, Babol, Iran Hani Ghabbani Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia Shadi Ghalami Department of Anatomy and Pathology, University of Siena, Siena, Italy Michael Glogauer Faculty of Dentistry, University of Toronto, Toronto, ON, Canada; Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto, ON, Canada
List of contributors xxiii Marc Grynpas Lunenfeld-Tanenbaum Research Institute, Mount Sinai Hospital, Toronto, ON, Canada; Department of Laboratory Medicine and Pathobiology, University of Toronto, Toronto, ON, Canada; Princess Margaret Cancer Centre, Department of Dental Oncology and Maxillofacial Prosthetics, Toronto, ON, Canada Nader Hamdan Department of Dental Clinical Sciences, Faculty of Dentistry, Dalhousie University, Halifax, NS, Canada Zoe Hancox Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom Nasira Haque Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom Robert G. Hill Dental Physical Sciences, Institute of Dentistry, Barts and The London School of Medicine and Dentistry, Queen Mary University of London, London, United Kingdom Wei-Te Huang Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom Shehriar Husain Department of Dental Materials Science, Jinnah Sindh Medical University, Karachi, Pakistan Seyed Hassan Jafari School of Chemical Engineering, College of Engineering, University of Tehran, Tehran, Iran Bassel Kano Division of Endodontics, Faculty of Dentistry, McGill University, Montreal, QC, Canada Abdul Samad Khan Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia Erum Khan Bhitai Dental and Medical College, Liaquat University of Medical and Health Sciences, Jamshoro, Pakistan; Faculty of Dentistry, King Abdulaziz University, Jeddah, Saudi Arabia Zohaib Khurshid Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia Tiantong Lou Faculty of Dentistry, University of Toronto, Toronto, ON, United States Maria Mali Department of Orthodontics, Islamic International Dental College & Hospital, Riphah International University, Islamabad, Pakistan
xxiv List of contributors Jukka P. Matinlinna Dental Materials Science, Applied Oral Sciences, Faculty of Dentistry, The University of Hong Kong, Prince Philip Dental Hospital, Sai Ying Pun, Hong Kong SAR, P.R. China Kyung-san Min School of Dentistry, Chonbuk National University, Jeonju, South Korea Masoud Mozafari Bioengineering Research Group, Nanotechnology and Advanced Materials Department, Materials and Energy Research Centre (MERC), Tehran, Iran; Department of Tissue Engineering & Regenerative Medicine, Faculty of Advanced Technologies in Medicine, Iran University of Medical Sciences (IUMS), Tehran, Iran; Cellular and Molecular Research Center, Iran University of Medical Sciences, Tehran, Iran Shariq Najeeb Independent Researcher and Private Practitioner, Alberta, Canada; National Center for Proteomics, University of Karachi, Pakistan Seung Yun Nam Interdisciplinary Program of Marine-Biomedical, Electrical and Mechanical Engineering, Center for Marine-Integrated Biomedical Technology (BK21 Plus), Pukyong National University, Busan, South Korea; Department of Biomedical Engineering, Pukyong National University, Busan, South Korea Hafiz Muhammad Owais Nasim Department of Dental Materials, Sharif Medical and Dental College, Lahore, Pakistan Touraj Nejatian Eastman Dental Institute, University College of London, London, United Kingdom; Nottingham Dental Clinic, Nottingham, United Kingdom; Royal College of Surgeons of England, London, United Kingdom Rafael Rocha Pacheco School of Dentistry, University of Detroit Mercy, Detroit, MI, United States Brouki Milan Peiman Cellular and Molecular Research Center, Iran University of Medical Sciences, Tehran, Iran; Department of Tissue Engineering & Regenerative Medicine, Faculty of Advanced Technologies in Medicine, Iran University of Medical Sciences, Tehran, Iran Sajjad Pezeshki Babol University of Medical Sciences, Babol, Iran Zeeshan Qamar Department of Oral and Maxillofacial Surgery, Riyadh Elm University, Riyadh, Saudi Arabia Jean-Marc Retrouvey Division of Orthodontics, Faculty of Dentistry, McGill University, Montreal, QC, Canada
List of contributors xxv Sahba Rezaei School of Chemical Engineering, College of Engineering, University of Tehran, Tehran, Iran Francisco Javier Rodrı́guez-Lozano School of Dentistry, University of Murcia, Murcia, Spain Vinicius Rosa Faculty of Dentistry, National University of Singapore, Singapore, Singapore; Centre for Advanced 2D Materials and Graphene Research Centre, National University of Singapore, Singapore, Singapore Mahsa Roshandel Department of Anatomy and Pathology, University of Siena, Siena, Italy Mohammad Reza Saeb Color and Polymer Research Center (CPRC), Amirkabir University of Technology, Tehran, Iran; Advanced Materials Group, Iranian Color Society (ICS), Tehran, Iran; Department of Resin and Additive, Institute for Color Science and Technology, Tehran, Iran Farshid Sefat Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom; Interdisciplinary Research Center in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom Saroash Shahid Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom Zeeshan Sheikh Faculty of Dentistry, University of Toronto, Toronto, ON, Canada; Lunenfeld-Tanenbaum Research Institute, Mount Sinai Hospital, Toronto, ON, Canada; Department of Laboratory Medicine and Pathobiology, University of Toronto, Toronto, ON, Canada Dalia Sherief Dental Biomaterials Department, Faculty of Dentistry, Ain-Shams University, Cairo, Egypt Sunjay Suri Division of Orthodontics, Faculty of Dentistry, University of Toronto, Toronto, ON, United States Ahmed Talal Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia Waqas Tanveer Maxillofacial Prosthetic Service, Department of Prosthodontics, Faculty of Dentistry, Mahidol University, Bangkok, Thailand
xxvi List of contributors James K.H. Tsoi Dental Materials Science, Discipline of Applied Oral Sciences, Faculty of Dentistry, The University of Hong Kong, Pokfulam, Hong Kong SAR, P.R. China Rizwan Ullah Department of Oral Biology, Sindh Institute of Oral Health Sciences, Jinnah Sindh Medical University, Karachi, Pakistan Daniel Varley Medical Engineering Department, Faculty of Engineering and Informatics, University of Bradford, Bradford, United Kingdom Gaurav Vasudeva School of Dentistry, James Cook University, Townsville, QLD, Australia; Oral Health Services, Hobart, TAS, Australia Jayachandran Venkatesan Yenepoya Research Centre, Yenepoya University, Mangalore, India Rafael Pino Vitti Department of Dentistry, University of Taubaté, Taubaté, Brazil; Department of Dentistry, Araras Dental School (FHO|UNIARARAS), Araras, Brazil Syed Azeem Ul Yaqin Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia Safiyya Yousaf Medical Engineering Department, Faculty of Engineering and Informatics, University of Bradford, Bradford, United Kingdom Mansour Youseffi Medical Engineering Department, Faculty of Engineering and Informatics, University of Bradford, Bradford, United Kingdom Muhammad S. Zafar Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan; Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia Muhammad Sohail Zafar Department of Restorative Dentistry, College of Dentistry, Taibah University, Almadinah Almunawwarah, Saudi Arabia Payam Zarrintaj Polymer Engineering Department, Faculty of Engineering, Urmia University, Urmia, Iran; Color and Polymer Research Center (CPRC), Amirkabir University of Technology, Tehran, Iran; Advanced Materials Group, Iranian Color Society (ICS), Tehran, Iran
Introduction to dental biomaterials and their advances 1 Zohaib Khurshid1, Muhammad S. Zafar2,3, Shariq Najeeb4, Touraj Nejatian5,6 and Farshid Sefat7 1 Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia, 2Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia, 3Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan, 4 National Center for Proteomics, University of Karachi, Pakistan, 5Royal College of Surgeons of England, London, United Kingdom, 6Nottingham Dental Clinic, Nottingham, United Kingdom, 7Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom Chapter Outline References 3 Further reading 5 Dental biomaterials and tissue engineering are rapidly developing approaches being used for the production of new organs and body tissues, particularly for bone implants or dental tissue replacements. Yet, over the past few decades, there has been a wide range of research conducted on the provision of tissue engineered dental grafts that has led to a significant improvement in the production of scaffolds with similar characteristics to a natural tooth (Zafar and Ahmed, 2015). Tissue engineering in the 21st century has become a cutting-edge science in the field of medicine and it is expected in the near future to replace traditional therapies which cause enormous side effects. In the tissue engineering principles, one of the main elements after cells, environmental factors, and signaling molecules is the biomaterial, which plays an important role in successful functional tissue engineered products. In recent years significant improvement and progress have been reported in the reconstruction of various human tissue replacements and prostheses, including bone (Sefat et al., 2010, 2014), cartilage (Daghigh Ahmadi et al., 2018; Raja et al., 2018), skin (Mahjour et al., 2015; Bye et al., 2014), oral tissues (Nejatian et al., 2017; Zafar and Ahmed, 2015; Qasim et al., 2018; Najeeb et al., 2017), cornea (Deshpande et al., 2013; Ortega et al., 2014), nerve (Mohamadi et al., 2017; Mohammadi et al., 2018), and adipose tissue (Amini et al., 2018). The most common methods used in the fabrication of tissue engineering scaffolds, particularly as dental biomaterials, consist of hydrogels, molecular self-assembly, thermally induced phase separation, solvent casting, particulate-leaching techniques, and Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00001-3 Copyright © 2019 Elsevier Ltd. All rights reserved.
2 Advanced Dental Biomaterials the electrospinning process (Gentile et al., 2017; Mahjour et al., 2016). In particular, for dental tissue engineering much research has been carried out using electrospinning (Zafar et al., 2016; Qasim et al., 2018), and this is the most widely used method in the production of tissue engineered teeth. Dentistry, similarly to medicine and other related subjects such as biomedical engineering, biomedical science, pharmacy, and pharmacology, has been revolutionized by the introduction of new technologies, for example, dentists use digital technology to carry out dental treatments. For this reason, we have included an interesting chapter on digital dentistry in this book. This book covers both the basic and clinical sciences of dental biomaterials with the view to meeting the needs of researchers and practitioners. In this book, the authors mainly look at various dental biomaterials’ properties and characterization techniques, such as metal alloys, polymers, composites, and ceramics. A chapter will focus on dental gypsum and investment materials covering composition, types, manufacturing, and applications. Gypsum products can be found as a powder which forms a fluid mass once combined with water and, therefore, has the capability to be shaped, before finally hardening into a rigid, stable mass. The main application of gypsum products is to replicate the shape of oral structures, which will be discussed in this chapter. Acrylic resin is a biomaterial that has the required mechanical and physical properties (Zafar and Ahmed, 2014), and it has been used as a denture material both in research and the clinic. Much research has been done on the biocompatibility and biodegradability of this material (Nejatian et al., 2015). Acrylic resin will be discussed in a separate chapter in this book. Another chapter is mainly focused on dental amalgam, which has been used successfully in the dental industry for decades. Dental amalgam is composed of a mixture of metal alloy and liquid mercury. Amalgam is mainly used for dental restorations. The current trend in the United Kingdom is to phase-down amalgam application due to the increasing concern over the safety of mercury, however, it is still widely used in many other countries, including the United States. Amalgam stays soft for a short period of time after it is mixed, which allows enough time for it to be condensed and shaped onto the prepared tooth (Gay et al., 1979; Bates, 2006). Dental resin composites are important biomaterials that have been increasingly used as the main restorative materials (Nejatian et al., 2017; Khurshid et al., 2015). In this book the specific types of composites as well as the potential biological issues of dental composites are discussed in detail. In addition, resin-based dental composites for tooth filling are addressed in a separate chapter due to their importance. Cements are another important group of biomaterials that have been explored intensively by researchers and clinicians. A separate chapter is allocated to dental cements, including base, liner, luting, and temporary cements, as well as pulp capping materials. Due to the high clinical demand for dental cements, the chemistry and applications of glass ionomer cement (GIC) and nano-GIC have been discussed broadly here in this book.
Introduction to dental biomaterials and their advances 3 Dental impression materials have been utilized as negative replicas to fabricate dental models. Various natural and synthetic materials have been utilized as dental impression materials, which are frequently classified on the basis of their properties into two groups, that is, elastic and nonelastic families. In this book, impression materials and their properties are discussed with the aim of providing a practical guide for dentists and prosthodontists. Fiber reinforced composites (FRCs) are combinations of conventional dental resin composites and glass fibers. This combination provides a high level of strength in the same way as that found in boats, light aircraft, and F1 racing cars. For many decades, engineers have been using fibers as fillers to construct materials/devices with high levels of strength and fracture toughness. That is why these materials attracted attention as desirable restorative materials in the field of dentistry (He et al., 2017). FRCs have been the center of attention in dentistry in the 21st century because of their excellent adhesion and appearance. Nowadays these materials have many clinical applications, such as fixed prosthodontics, restorative dentistry, periodontology, orthodontics, and repair of prosthetic devices (Nagata et al., 2016). FRCs are discussed in detail in a chapter of this book. Additionally, a number of emerging materials that have been extensively explored for various dental applications, such as nanomaterials (Najeeb et al., 2015, 2016a,b,c,d; Zafar et al., 2017), polyether ether ketones (Najeeb et al., 2016a,b), natural silk (Zafar and Ahmed, 2014), and antimicrobial peptides (Khurshid et al., 2016a,b, 2017, 2018), will be discussed in parts of various chapters. The use of GICs as direct restorative dental materials, which have been extensively studied and used in dentistry (Zafar and Ahmed, 2015; Najeeb et al., 2016a, b,c,d), is also discussed in detail in this book. Other advanced dental biomaterials are also covered in great detail including endodontic materials, advanced ceramics, bone cements, calcium phosphate, bioactive glasses, graphene, and silicon, as well as orthodontic materials including wires, orthodontic brackets, elastomeric ligatures, and chains. The final chapter discusses biomaterials used for maxillofacial prosthetic rehabilitation/reconstruction, as well as biomaterials for craniofacial tissue engineering and regenerative dentistry which are hot topics in the field of bioengineering and regenerative medicine. References Amini, N., Vousooghi, N., Alizade, A., Ramezani, S., Joghataei, M.T., Brouki Milan, P., et al., 2019. Transplantation of adipose tissue-derived stem cells into brain through cerebrospinal fluid in rat models: protocol development and initial outcome data. Curr. Stem Cell Res. Ther. 14 (2), 191 195. Bates, M.N., 2006. Mercury amalgam dental fillings: an epidemiologic assessment. Int. J. Hyg. Environ. Health 209, 309 316. Bye, F.J., Bullock, A.J., Singh, R., Sefat, F., Roman, S., Macneil, S., 2014. Development of a basement membrane substitute incorporated into an electrospun scaffold for 3D skin tissue engineering. J. Biomater. Tissue Eng. 4, 1 7.
4 Advanced Dental Biomaterials Daghigh Ahmadi, E., Raja, T.I., Khaghani, S.A., Soon, C.F., Mozafari, M., Youseffi, M., et al., 2018. The role of photonics and natural curing agents of TGF-β1 in treatment of osteoarthritis. Mater. Today Procedia 5, 15540 15549. Deshpande, P., Sefat, F., Ramchadaran, C., Mariappan, I., Johnson, C., Mckean, R., et al., 2013. Simplifying corneal surface regeneration using a biodegradable synthetic membrane and limbal tissue explants. Biomaterials 34 (21), 5088 5106. Gay, D.D., Cox, R.D., Reinhardt, J.W., 1979. Chewing releases mercury from fillings. Lancet 1 (8123), 985 986. Gentile, P., Mccolgan-Bannon, K., Ceretto, N., Sefat, F., Dalgarno, K., Ferreira, A.M., 2017. Biosynthetic PCL-graft-collagen bulk material for tissue engineering applications. Materials 10, 693. He, J., Vallittu, P., Lassila, L.V., 2017. Preparation and characterization of high radio-opaque E-glass fibre-reinforced composite with iodine containing methacrylate monomer. Dent. Mater. 33 (2), 218 225. Khurshid, Z., Zafar, M., Qasim, S., Shahab, S., Naseem, M., AbuReqaiba, A., 2015. Advances in nanotechnology for restorative dentistry. Materials 8 (2), 717 731. Khurshid, Z., Naseem, M., Sheikh, Z., Najeeb, S., Shahab, S., Zafar, M.S., 2016a. Oral antimicrobial peptides: types and role in the oral cavity. Saudi Pharm. J 24 (5), 515 524. Khurshid, Z., Zohaib, S., Najeeb, S., Zafar, M.S., Rehman, R., Rehman, I.U., 2016b. Advances of proteomic sciences in dentistry. Int. J. Mol. Sci. 17 (5), 728. Khurshid, Z., Najeeb, S., Mali, M., Moin, S.F., Raza, S.Q., Zohaib, S., et al., 2017. Histatin peptides: pharmacological functions and their applications in dentistry. Saudi Pharm. J. 25 (1), 25 31. Khurshid, Z., Zafar, M.S., Naseem, M., Khan, R.S., Najeeb, S., 2018. Human Oral Defensins antimicrobial peptides: a future promising antimicrobial drug. Curr. Pharm. Des. 24 (10), 1130 1137. Mahjour, S.B., Fu, X., Yang, X., Fong, J., Sefat, F., Wang, H., 2015. Rapid creation of skin substitutes from human skin cells and biomimetic nanofibers for acute full-thickness wound repair. Burns 41 (8), 1764 1774. Mahjour, S.B., Sefat, F., Polunin, Y., Wang, L., Wang, H., 2016. Improved cell infiltration of electrospun nanofiber mats for layered tissue constructs. J. Biomed. Mater. Res. Part A 104 (6), 1479 1488. Mohamadi, F., Ebrahimi, S., Nourani, M.R., Mansoori, K., Alizadeh, A.A., Tavangar, S.M., et al., 2017. Enhanced sciatic nerve regeneration by human endometrial stem cells in an electrospun poly (ε-caprolactone)/collagen/NBG nerve conduit in rat. Artif. Cells Nanomed. Biotechnol. 46 (8), 1731 1743. Mohammadi, A., Maleki-Jamshid, A., Sanooghi, D., Brouki Milan, P., Rahmani, A., Sefat, F., et al., 2018. Transplantation of human chorion-derived cholinergic progenitor cells: a novel treatment for neurological disorders. Mol. Neurobiol. 56 (1), 307 318. Nagata, K., Garoushi, S.K., Vallittu, P.K., Wakabayashi, N., Takahashi, H., Lassila, L.V.J., 2016. Fracture behaviour of single-structure fibre-reinforced composite restorations. Acta Biomater. Odontol. Scand. 2 (1), 118 124. Najeeb, S., Khurshid, Z., Matinlinna, J.P., Siddiqui, F., Nassani, M.Z., Baroudi, K., 2015. Nanomodified peek dental implants: bioactive composites and surface modification—a review. Int. J. Dent. 2015, 381759. Najeeb, S., Bds, Z.K., Bds, S.Z., Bds, M.S., 2016a. Bioactivity and osseointegration of PEEK are inferior to those of titanium: a systematic review. J. Oral Implantol. 42 (6), 512 516. Najeeb, S., Zafar, M.S., Khurshid, Z., Siddiqui, F., 2016b. Applications of polyetheretherketone (PEEK) in oral implantology and prosthodontics. J. Prosthodont. Res. 60 (1), 12 19.
Introduction to dental biomaterials and their advances 5 Najeeb, S., Khurshid, Z., Zafar, M.S., Khan, A.S., Zohaib, S., Martı́, J.M., et al., 2016c. Modifications in glass ionomer cements: nano-sized fillers and bioactive nanoceramics. Int. J. Mol. Sci. 17 (7), 1134. Najeeb, S., Khurshid, Z., Agwan, A.S., Zafar, M.S., Alrahabi, M., Qasim, S.B., et al., 2016d. Dental applications of nanodiamonds. Sci. Adv. Mater. 8 (11), 2064 2070. Najeeb, S., Khurshid, Z., Agwan, M.A., Ansari, S.A., Zafar, M.S., Matinlinna, J.P., 2017. Regenerative potential of platelet rich fibrin (PRF) for curing intrabony periodontal defects: a systematic review of clinical studies. Tissue Eng. Regener. Med 1, 1 8. Nejatian, T., Sefat, F., Johnson, T., 2015. Impact of packing and processing technique on mechanical properties of acrylic denture base materials. Materials 8 (5), 2093 2109. Nejatian, T., Khurshid, Z., Zafar, M.S., Najeeb, S., Zohaib, S., Mozafari, M., et al., 2017. Dental biocomposites (Chapter 5). Biomaterials for Oral and Dental Tissue Engineering. Elsevier, pp. 65 83. Ortega, I., Sefat, F., Paterson, T., Deshpande, P., Ramchadaran, C., Claeyssens, F., et al., 2014. Combination of microstereolithography and electrospinning to produce membranes equipped with niches for corneal regeneration. J. Vis. Exp 91, e51826. Qasim, S.,B., Zafar, M.S., Najeeb, S., Khurshid, Z., Shah, A.H., Husain, S., et al., 2018. Electrospinning of chitosan-based solutions for tissue engineering and regenerative medicine. Int. J. Mol. Sci. 19 (2), 407. Raja, T.I., Khaghani, S.A., Zafar, M.S., Khurshid, Z., Mozafari, M., Youseffi, M., et al., 2018. Effect of TGF-β1 on water retention properties of healthy and osteoarthritic chondrocytes. Mater. Today Proc. 5 (7), 15717 15725. Sefat, F., Youseffi, M., Denyer, M.C.T., 2010. Imaging via widefield surface plasmon resonance microscope for studying bone cell interactions with micro-patterned ECM proteins. J. Microsc. 241 (3), 282 290. Sefat, F., Denyer, M.C.T., Youseffi, M., 2014. Effects of different transforming growth factor beta (TGF-β) isomers on wound closure of bone cell monolayers. Cytokines 64, 75 86. Zafar, M.S., Ahmed, N., 2014. Nanoindentation and surface roughness profilometry of poly methyl methacrylate denture base materials. Technol. Health Care 22 (4), 573 581. Zafar, M.S., Ahmed, N., 2015. Therapeutic roles of fluoride released from restorative dental materials. Fluoride 48, 184 194. Zafar, M.S., Najeeb, S., Khurshid, Z., Vazirzadeh, M., Zohaib, S., Najeeb, B., et al., 2016. Potential of electrospun nanofibers for biomedical and dental applications. Materials 9 (2), 73. Zafar, M.S., Khurshid, Z., Najeeb, S., Zohaib, S., Rehman, I.U., 2017. Therapeutic applications of nanotechnology in dentistry. Nanostructures for Oral Medicine. Elsevier, pp. 833 862. Further reading Zafar, M.S., Al-Samadani, K.H., 2014. Potential use of natural silk for bio-dental applications. J. Taibah Univ. Med. Sci. 9 (3), 171 177. Zafar, M.S., Khurshid, Z., Almas, K., 2015. Oral tissue engineering progress and challenges. Tissue Eng. Regener. Med. 12 (6), 387 397.
Properties of dental biomaterials 2 Muhammad S. Zafar1,2, Rizwan Ullah3, Zeeshan Qamar4, Muhammad A. Fareed5, Faiza Amin6, Zohaib Khurshid7 and Farshid Sefat8,9 1 Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia, 2Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan, 3Department of Oral Biology, Sindh Institute of Oral Health Sciences, Jinnah Sindh Medical University, Karachi, Pakistan, 4 Department of Oral and Maxillofacial Surgery, Riyadh Elm University, Riyadh, Saudi Arabia, 5Adult Restorative Dentistry, Dental Biomaterials and Prosthodontics Oman Dental College, Muscat, Sultanate of Oman, 6Department of Science of Dental Materials, Dow Dental College, Dow University of Health Sciences, Karachi, Pakistan, 7Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia, 8Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom, 9Interdisciplinary Research Centre in Polymer Science and Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom Chapter Outline 2.1 Introduction 8 2.2 Optical properties (color) 2.3 Thermal properties 9 2.3.1 2.3.2 2.3.3 2.3.4 2.3.5 2.3.6 2.3.7 2.4 2.5 2.6 2.7 2.8 8 Temperature 9 Transition temperatures 9 Heat of fusion (L) 11 Thermal conductivity (K) 12 Specific heat (Cp) 13 Thermal diffusivity (Δ) 13 Coefficient of thermal expansion (α) 14 Viscosity 14 Electrical conductivity and resistivity 15 Mechanical properties and characterization methods Limitation of mechanical testing methods 22 Biological properties 22 2.8.1 2.8.2 2.8.3 2.8.4 Biocompatibility 22 In vitro testing 23 In vivo testing 24 Usage tests 24 2.9 Toxicity and cytotoxicity 24 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00002-5 Copyright © 2019 Elsevier Ltd. All rights reserved. 16
8 Advanced Dental Biomaterials 2.10 Cytotoxicity tests 26 2.11 Fluoride and caries 26 2.11.1 Fluoride toxicity 27 2.12 Carcinogenicity 27 2.13 Biodegradation 28 2.14 Bioactivity 28 2.15 Osseointegration 29 2.16 Osteoinduction 29 2.17 Foreign body reaction 29 2.18 Conclusive remarks 30 References 30 2.1 Introduction Understanding the properties of dental biomaterials is important in order to compare with the properties of oral hard and soft tissues prior to any clinical application. Various dental restorations tend to fail due to a number of reasons such as distortion or mechanical failure. The dental restorations are sometimes incompatible with the supporting oral tissues due to the interface or substrate failure. Although no dental biomaterial has been reported as having ideal properties to date, but it is worth mentioning that the performance and clinical success of dental biomaterials is strongly associated with their properties. In this chapter, various physical, mechanical, and biological properties of dental biomaterials and related interaction are discussed. 2.2 Optical properties (color) Color is perceived as an end result of the sensory response to light (Costa, 2016). The physiological stimuli or sensory response is experienced by an individual, whereas the light beam is considered as the physical stimuli which produce the sensory response. Color is perceived due to reflection or transmission (partial or complete) of white light. According to Grassmann’s law, the normal eye can differentiate in three color parameters, which are dominant wavelengths, luminous reflectance, and the excitation purity (Mausfeld, 1998; Grassmann, 1853). The dominant wavelength (λ) is the wavelength of a monochromatic light which on combination with achromatic color (gray) matches the perceived color (Klein and Meyrath, 2010). Light with short (400 nm) and long (700 nm) wavelengths are violet and red in color, respectively (Klein and Meyrath, 2010). The wavelength range of visible light (400700 nm) corresponds to specific colors (e.g., blue, green, yellow, and orange). This property of color that is distinguished by light’s wavelength is called “hue.” Among all the colors and shades, there are only three primary colors i.e., red, green, and blue. These three colors when used in appropriate
Properties of dental biomaterials 9 proportions can produce different colors, for example, yellow color can be produced by an appropriate combination of green and red. The second parameter is luminous reflectance of color (value), which classifies objects equivalent to a series of achromatic scale (López Camelo and Gómez, 2004). For the light diffusing objects, it ranges from black to white. While in the case of transmitting objects, it ranges from black to clear or colorless. The black is specified as a standard with luminous reflectance of 0 and the white assigned as 100. The third component (saturation of color, also called “Chroma”) defines the degree of difference from achromatic color toward the color spectrum; it numerically ranges from 0 to 1 (Sakaguchi and Powers, 2012). 2.3 Thermal properties Thermal properties are important in regulating the performance of dental materials. At the molecular level, the structural arrangement and movements of atoms are affected as a function of temperature variations. Therefore it is necessary to have an understanding of different thermal techniques used to identify thermal properties of dental materials. 2.3.1 Temperature A thermometer or a thermocouple can be used to measure the temperature changes of various substances (Peyton, 1952). For instance in the dental clinics, the temperature variations are observed as a result of heat generated during cavity preparation or while curing of resin composite materials. Factors such as headpiece rotational speed and use of coolants are likely to influence the temperature changes during tooth cutting (Fig. 2.1). Thermocouples are used for the measurement of temperature by inserting it near the dentino-enamel junction. 2.3.2 Transition temperatures A number of dental materials, such as dental waxes, are highly sensitive to minute thermals changes. The differential thermal analysis is a technique used to identify different constituents (such as paraffin and carnauba) of dental waxes (Brauer et al., 1970; Craig et al., 1965). A thermogram was developed on observing temperature differences between both waxes under the standard conditions using thermocouples. Temperature differences were recorded as a function of temperature surroundings. Decrease in temperature difference (ΔT) indicated an endothermic reaction in the sample (Brauer et al., 1970; Craig et al., 1967). The solidsolid endotherms at 31.5 C and 35 C were observed as a result of changes in the crystal structure for the paraffin wax. The solidliquid transition endotherms were observed at 52 C for paraffin wax; on the other hand the endotherms at 68.7 C and 80.2 C result from melting of carnauba wax (Craig et al., 1965; Craig et al., 1967). The transition of
Advanced Dental Biomaterials Temperature rise (ºF) 10 37 Steel bur 180 170 130 120 110 100 90 80 70 60 50 40 30 20 10 0 Water spray Air No coolant 10,000 20,000 30,000 Speed of rotation (RPM) Figure 2.1 The temperature fluctuations in the tooth tissue during cavity preparation at variable speeds (Peyton, 1958). 0 0 80.2 25 68.7 ΔT (°C) 1.0 50 1.5 75 2.0 Penetration (%) 31.5 0.5 35 DTA 2.5 TMA 0.013 MN/m2 100 2 TMA 0.26 MN/m 52 20 40 60 80 ºC 100 120 Figure 2.2 Thermogram of a paraffin (75%) and carnauba (25%) wax mixture comparing their behavior in response to lower (0.013 MPa) and higher (0.26 MPa) stress (Brauer et al., 1970; Powers and Craig, 1974). heat for two solidsolid transitions is approximately 8 cal/g, the melting transition for the paraffin was 39 cal/g and for carnauba wax 11 cal/g (Craig et al., 1965). The penetration of the testing indenter was recorded using two quantitative stresses 0.013 and 0.26 MPa (Fig. 2.2). At lower stress (0.013 MPa), the penetration was regulated by the melting transition of the carnauba wax. On the other hand, at the
Properties of dental biomaterials 11 higher stress (0.26 MPa), the penetration was facilitated mainly by solidsolid and solidliquid transitions of paraffin wax. Nearly 44% of the wax penetration occurred before the melting of paraffin wax, which is corresponding to its flowing potential (Powers and Craig, 1974). Another important property correlated with the thermograms is the coefficient of thermal expansion. This coefficient is raised approximately from 300 3 1026/ C to 1400 3 1026/ C prior to solidsolid transition. The property of flow is also observed to increase in this temperature range. The dynamic mechanical analysis is another technique in order to analyze the thermal properties of a material. For this purpose, a thin film of di-methacrylate copolymer is subjected to tensile strain at a specific frequency of 11 Hz (Wilson and Turner, 1987). The elastic modulus and loss of tangent are obtained by raising the temperature. The glass transition temperature (Tg) is measured from the decrease of elastic modulus with temperature changes. The Tg determines the temperature at which polymer is transformed to a softer, rubbery state on heating (Wilson and Turner, 1987). The lower value of glass transition temperature can be affected by the lower degree of alteration of double bonds, less cross-linking, and better flexibility of networks. The coefficient of thermal expansion of a polymer is altered at the glass transition temperature. 2.3.3 Heat of fusion (L) The heat in calories (Cal) or Joules (J) which is required for conversion of 1 g of material from the solid phase to the liquid phase (melting temperature) is referred to as heat of fusion. The heat of fusion can be calculated by L5 Q m where L is the heat of fusion, Q refers to the total amount of heat absorbed, and m refers to the mass of the substance being melted. Thus, practically, the mass of material is directly proportional to the heat required in changing the total mass to liquid. The heat of fusion of a material is correlated with melting or freezing. For the change in the state from a solid mass to liquefaction it is important to add heat, as long the heat of fusion is retained by the liquid, the mass remains molten. As soon as the heat is liberated from the liquid state the material solidifies. The difference in the energy content is of key importance in order to maintain the kinetic molecular motion, an important property of a material in a liquid state. The heat of fusion values for commonly used materials are listed in the Table 2.1 (Craig and Powers, 2002). It can be observed from Table 2.1 that the values for dental alloys, particularly silver, gold, and copper, are lower than that of many other alloys and metals.
12 Advanced Dental Biomaterials Table 2.1 Heat of fusion of commonly used materials (Craig and Powers, 2002). Materials Temperature ( C) Heat of fusion (cal/g) [J/g] Metals Mercury Gold Silver Platinum Copper Cobalt Chromium Aluminum 239 1063 960 1773 1083 1495 1890 660 3 [12] 16 [67] 26 [109] 27 [113] 49 [205] 58 [242] 75 [314] 94 [393] 2114 52 62 18 0 25 [104] 35 [146] 42 [176] 47 [196] 80 [334] Compounds Alcohol Paraffin Beeswax Glycerin Ice 2.3.4 Thermal conductivity (K) The heat (in calories or joules per second) that is passing through the 1 cm thick body with cross section of 1 cm2 when the temperature difference is 1 C is known as the thermal conductivity. The unit for the thermal conductivity is cal/s/cm2 ( C/cm). Changes in thermal conductivity are observed on variation in the surrounding temperature, but these changes are negligible in comparison to that which exists between different materials. There is an important role of thermal conductivity in the dental materials. A classic representation for the thermal conductivity can be shown by an example of a tooth restored with dental amalgam filling or dental crown made of gold alloy in close proximity to the dental pulp. Such restoration possibly will lead to discomfort on use of hot or cold food products which can produce changes in the temperature. The effect of thermal conductivity can be alleviated if adequate dental tissue is intact or cavity liners are placed between the tooth structure and the dental restorative material, which has a capability of providing insulation since it is a poor thermal inductor. Dental materials that provide adequate insulation are listed in Table 2.2 (Brown et al., 1970; Lisanti and Zander, 1950). Nonmetallic materials have less thermal conductivity than metals making them desirable and good insulators. Nonmetallic restorative dental materials usually have similar thermal conductivity to those of the enamel and dentin.
Properties of dental biomaterials 13 Table 2.2 Thermal conductivity of various dental restorative materials (Brown et al., 1970; Lisanti and Zander, 1950). Materials Thermal conductivity Cal/s/cm2/( C/cm) J/s/cm2/( C/cm) 1.006 0.918 0.710 0.055 4.21 3.84 2.97 0.23 0.0031 0.0026 0.0022 0.0015 0.013 0.011 0.0092 0.0063 Metals Silver Copper Gold Dental amalgam Nonmetals Gypsum Composite resin Enamel Dentin 2.3.5 Specific heat (Cp) The specific heat is the amount of heat energy required to raise 1 g of a substance’s temperature by 1 C. Water is usually used as a standard for comparison. The total energy required in order to raise the temperature by 1 C is dependent on the total mass of the material. For example, in order to increase the temperature by 1 C, 100 g of water requires more calories than would be required for 50 g of water. Similarly due to the variance in specific heat of water and alcohol, 100 g of water needs a higher degree of heat in comparison with alcohol of the same amount. Considered on a broad spectrum, the specific heat of liquids is higher than solids. For the melting and casting process, a metal or alloy’s specific heat is of key importance for the heat required to raise the temperature of metallic mass until it reaches the melting point. The specific heat of gold and gold alloys is lower than majority of restorative materials and dental tissues; therefore heating for a longer duration is not required. 2.3.6 Thermal diffusivity (Δ) The measurement of transient heat flow is referred to as thermal diffusivity. The thermal diffusivity is measured as: Δ5 K Cp 3 ρ where, K, Cp, and ρ represent the thermal conductivity, specific heat, and the density of the material, respectively (Brady et al., 1974; Carslaw and Jaeger, 1959; Tay
14 Advanced Dental Biomaterials and Braden, 1987). The unit for the thermal diffusivity is mm2/s. The gold crown or amalgam restorations have high thermal conductivity and low specific heat and are likely to cause potent thermal shock in comparison with the normal tooth structure. Similar to thermal conductivity, the material thickness is important in regulating the thermal diffusivity (Brady et al., 1974). 2.3.7 Coefficient of thermal expansion (α) The coefficient of thermal expansion is referred to as dimensional changes in a material per unit for the change by 1 C in temperature (Fairhurst et al., 1980). The coefficient of thermal expansion can be calculated using the following equation: α5 ðLfinal 2 Loriginal Þ ½Loriginal 3 Cfinal 2 Coriginal The units for thermal expansion are expressed as the notation / C as the values are small and they are therefore generally used in exponential form, for example, 12 3 1026/ C (Fairhurst et al., 1980). 2.4 Viscosity The fluid viscosity is observed to have a direct relation with the shear rate; it increases with increasing shear rate. The proportionality of the viscosity varies for different fluids. The viscosity of the fluids may vary according to the shear rate; therefore fluid can further be classified as Newtonian, pseudoplastic, or dilatant. The Newtonian fluids are reported to have constant viscosity which is independent of its shear rate. Classic examples for Newtonian fluids are dental cements and impression materials. For the pseudoplastic fluid the viscosity decreases with an increasing shear rate. The best example for the pseudoplastics is the monophase elastomeric materials (Combe and Moser, 1978). These materials tend to have high viscosity therefore on mixing or while placing the impression material containing tray into the mouth, it remains in its place without flowing (Herfort et al., 1977). Elastomeric materials can be used in syringes as they encounter higher shear rates while passing through the syringe tip. On exiting the syringe the viscosity decreases by 10-fold (Herfort et al., 1977). This property of material is referred to as thixotropy, although the term “thixotropy” particularly defines the alterations in the viscosity of a material with time. A material whose viscosity increases with increasing shear rate is referred to as a dilatant fluid. In dentistry, a classic example for dilatant fluids is the denture base resins (Vermilyea et al., 1978).
Properties of dental biomaterials 2.5 15 Electrical conductivity and resistivity The conductance or conductivity can be described as the potential of a material to conduct electrical current, or contrariwise, as the specific resistance or resistivity. At a constant temperature, the homogeneous conductor’s resistance is directly proportional to the length and inversely proportional to the cross section of the specimen and can be calculated using the following equation: R5 ρ3L A The resistance is presented by R, resistivity by ρ, length by L, and section area by A. Resistivity (Ω cm) varies with the type of the material. For example, in the case of a 1 cm3 cube, the length and section area are equal to each other and can be expressed as R 5 ρ (Mumford, 1967). The variation in ohms can be used to understand the changes occurring in the structure of different alloys when subjected to heat. On investigation of electrical conductivity for the goldcopper alloy system, there are changes in the internal crystal structure in conjunction with its conductivity. The resistivity values of dental hard tissues are shown in Table 2.3. Resistivity of the structure is of key importance to determine the perception for pain threshold which results from the application of electrical stimuli and the displacement of fluid in tooth structure due to the ionic movements (Mumford, 1967). The electrical resistance of a healthy tooth differs from that of the infected structure, as less resistance is offered by the carious tissues. An intact enamel structure is relatively a poor electric conductor compared to dentin (Table 2.3) (Mumford, 1967). The electrical conductivity of restorative dental materials is of key concern. Various researchers have measured the resistivity of dental restorative materials (Table 2.3). The dental cement zinc oxideeugenol is found to have the highest resistivity values followed by the zinc phosphate cements. The glass ionomer cements’ conductivity matches closely to that of dentin. Table 2.3 The resistivity (Ω cm) values of dental hard tissues and dental restorative materials (Mumford, 1967; Tay and Braden, 1981; Braden and Clarke, 1974; Phillips et al., 1955). Material Ω cm Human enamel 2.93.6 3 106 2.62.9 3 106 0.76.0 3 104 1.15.2 3 104 0.82.5 3 104 1091010 2 3 105 Human dentin Glass ionomer Zinc oxideeugenol Zinc phosphate
16 2.6 Advanced Dental Biomaterials Mechanical properties and characterization methods The vibrant human oral atmosphere is able to influence the dental biomaterials used for tooth restoration. One of the key requirements of dental materials is to match the properties of the tooth structure which are required in a particular application (Tillberg et al., 2008). In order to identify the appropriate mechanical properties relevant to a particular use various test methods are employed. Therefore, an understanding of mechanical properties allows to distinguish the potential causes of clinical deficiencies related to the mechanical failure of dental materials under oral load. The basic parameter to determine mechanical behavior is to understand the stressstrain relationship for deformation (Vaderhobli, 2011). Mechanical properties determine the resistance fracture or deformation under an applied force or pressure. Analysis of mechanical properties will help to understand the failure and longevity of dental materials in the dynamic oral environment under various types of stresses and forces (De Jager et al., 2006). The mechanical behavior of a material is mainly related to the response to a force or load (Zafar and Ahmed, 2014b,c) and this behavior ultimately decides the usefulness of a material in a specific application. The mechanical testing and characterization are usually carried out according to existing standards set by various international organizations such as american dental association (ADA), the British Standards, or the Federation Dentaire Internationale (McCabe and Walls, 2008). One of the main purposes of these standards is to provide the technical information regarding an unbiased and dependable selection process for the materials used by health care providers. Certain claims of the materials’ quality arise from manufacturers as a result of standard test methodologies for mechanical, physical, chemical, and biological tests. The mechanical testing performed by the manufacturers to measure mechanical properties such as compression strength (CS), tensile strength, flexural strength (FS), wear, modulus, and biocompatibility must be performed under the specific testing conditions to identify the safest and the most efficient materials for specific clinical applications (Basu et al., 2010). The mechanical properties demonstrate how a dental material and the tooth structure react to the applied forces. Therefore it is important to understand the application of force or stress and pressure on a material; such forces will change the shape or structure of the materials if not resisted. 1. Force or pressure: Force is an applied energy to move or to deform a material. During the normal function and biting, various types of forces are applied on the tooth structure, alveolar socket, and periodontal complex (Van Noort and Barbour, 2013). For example, axial, vertical forces, horizontal forces, torsion and bending forces, or any combination of these forces (Fig. 2.3) are applied to the tooth or materials by the muscles of mastication while biting and through parafunctional habits. External forces due to dental trauma, orthodontics, and tooth movement are also present. Three characteristics of force determine the outcome of the applied force, that is, the magnitude, the direction, and the point of application of force and these result in various deformations of the materials (Fig. 2.3).
Properties of dental biomaterials 17 Figure 2.3 Various types of forces and possible deformation according to force directions. Several test methods are designed to measure the mechanical properties of dental material by destructive testing methodologies, that is, compressive, tensile, impact tests, hardness, brittleness, fatigue, and wear resistance. The unit of force is newton or pound. 2. Stress and strain: When a force is applied on a material, stress is equal to the force applied divided by per unit area, that is, forces/unit area, and the unit is Newton/meter square. Stress or a force that is applied to a material produces a strain which is opposite to the external force but equal in magnitude. Strain is the change in the size (length) of the material divided by the original size (length). The unit of stress is pascal (Pa) or megapascal (MPa). Compressive stresses are produced when a material is subjected to forces in a straight line directed in the same direction toward the center. Compressive stress produces comprehensive strain (Sakaguchi and Powers, 2012), as shown in Fig. 2.4. Tensile stress results in a material when two set of forces are applied in a straight line but in opposite directions, that is, the material is pulled apart (Fig. 2.4). Tensile stress produces tensile strain. 3. Elastic and plastic regions: In the stressstrain graph, two types of stresses are common, elastic and plastic, which produce the corresponding elastic strain and plastic strain. Elastic stresses do not produce any deformation; therefore after the removal of elastic stress, a material will come back to its original dimension and it remains unchanged due to elastic strain, whereas plastic stresses produce permanent deformation and the material will not come back to its original size or shape (Sakaguchi and Powers, 2012). Several mechanical behaviors are determined from elastic strain, for example, elastic modulus, flexibility, resilience, and Poisson’s ratio. On the other hand, the initial plastic deformation region and/or the end of the elastic region of a stressstrain graph help to determine the yield point, proportional limit, elastic limit, and yield strength. It is important to understand stress and strain because they are independent of the size of the material used. For example, for two cylindrical shape specimens of composite
18 Advanced Dental Biomaterials Figure 2.4 Various types of strain and deformations based on force direction. having 6 mm length but different cross-sectional area (2 and 4 mm), it will require double the force to break the larger cylinder; however, the stressstrain relationship will remain the similar. Therefore the stressstrain curve relationship is considered as the most predictable method to measure the mechanical properties. Various types of forces are applied to determine the corresponding stressstrain values before or after deformation or fracture. Some of these properties are discussed here: a. Proportional limit: The highest amount of stresses that a material can withstand without deviation from the proportionality of stressstrain. In this region both stress and strain are directly proportional to each other and it is present in the elastic region, therefore no permanent deformation occurs. The stressstrain cure region ahead of the proportional limit is known as the plastic region; stresses higher than the proportional limit lead to plastic deformation. A clinical application of the proportional limit is the connectors of cast partial denture as it withstand high stresses with plastic deformation. b. Elastic limit: The maximum amount of stress that a material can withstand without deformation or fracture. Elastic limits deal with the elasticity of a material as it is the point beyond which strain is not recoverable. Theoretically, the values of elastic limit and proportional limit remain the same. c. Yield strength: These are the stresses at which materials show a specific limiting diversion from the proportionality of stress to strain and a material begins to function in a plastic manner without failure or fracture. For example, while shaping an orthodontic wire or appliance and during clasp adjustment of cast removable, partial denture stresses are induced in excess of yield strength to achieve the desired outcomes. d. Elastic modulus (stiffness or rigidity): Elastic modulus is the measure of relative rigidity or stiffness of a material within the elastic region. This is an inherited measurement of a material’s mechanical behavior and is measured by the ratio of stress and strain in the elastic range (Modulus of elasticity 5 stress/strain). It is the measure of intrinsic properties of a material: the stronger the interatomic forces
Properties of dental biomaterials 4. 5. 6. 7. 19 (basic interaction forces), the greater the value of elastic modulus, and therefore, the materials would be more rigid and stiff (resistance to elastic deformation). Generally, the modulus of elasticity of dental materials is directly related to hardness (Zafar and Ahmed, 2014a,b; Zafar, 2014). e. Resilience and toughness: The amount of recoverable energy within the elastic region absorbed by a material when it is stressed to its proportional limit is known as resilience. Whereas the energy absorbed by a material while under plastic deformation until it fractures is called toughness. Resilience is defined as the unit area under the elastic region and toughness and is measured by the area under the plastic region till the point of fracture. Denture soft relining materials and orthodontic wires are example of resilient materials (Fig. 2.5). Ductility and malleability: Ductility is a materials’ ability to sustain huge amounts of permanent deformation under tensile loading without fracture or failure. Metals are ductile materials which can be drawn into wire with the help of tensile forces. On the other hand, malleability is the ability of a material to resist permanent deformation under compressive load without failure or fracture. Metals are malleable as it can be deformed in thin sheets under compressive forces, that is, gold foils, and wires or cords under tensile stress such as orthodontic wires and stainless steel crowns. Brittleness: This is the ability of a material to undergo failure or fracture without any or very little deformation. Therefore, it is the opposite to ductility. Most brittle materials fracture at the yield point or at the proportional limit. The examples of brittle materials are gypsum, dental cements, resin composite, dental ceramics, and dental amalgam (Fig. 2.6). Fracture stress: The amount stress that breaks a material break is known as fracture stresses or breaking stresses. It is not always necessary that a material fractures at the point of maximum stress. Strength: Defined as the maximum amount of stresses a material can withstand without failure of fracture. It is not an intrinsic property of the materials as it is recorded after an external force produces internal stresses leading to failure or fracture. The unit of strength is MPa or GPa. Various types of testing methodologies are employed to determine the various types of strengths, which depend upon the type of forces, direction, and magnitude of forces (International Standards Organization, 2000). a. Compressive strength (CS): The maximum amount of stress a material can withstand before fracture under compressive load or forces (ASTM, 2010). In compressive strength measurement, a cylinder-shaped specimen of the brittle material is vertically Figure 2.5 Stressstrain curves showing mechanical behavior of materials: (A) resilience (shaded area) and (B) toughness (shaded area).
20 Advanced Dental Biomaterials Figure 2.6 Stressstrain curves showing mechanical behavior for brittle, ductile, stiff, and elastic materials. compressed until it fractures. The CS was calculated in MPa from the following equation: CS 5 4P/πd2, where P is the maximum force applied at fracture and d is the diameter. b. Tensile strength: The maximum amount of stress a material can withstand before fracture under tensile load or forces. Tensile strength is sometimes known as ultimate tensile strength. To ensure good gripping, test specimens for tensile strength are often enlarged compared to the central portion of specimen (ASTM Standard, 2014). c. Diametral tensile strength or dismetral compression test is an alternative way to measure the tensile strength; therefore it is also known as an indirect tensile test. In this test method a disk-shaped specimen (4 3 2 mm2) of the brittle material is diametrically compressed until it fractures. While testing, compressive forces applied on the specimen induce tensile stresses perpendicular to the direction of applied force (tensile stresses are directly proportional to the applied force) and are calculated as follows: Tensile strength 5 2F/πDT, where F is the force, D is the diameter, and t is the thickness of the specimen. d. FS or transverse strength or modulus of rupture is measured by employing the threeor four-point bend test method (Fig. 2.7).
Properties of dental biomaterials 21 Figure 2.7 Schematic presentation of force/stress application used for various methods of flexural strength testing. e. In the three-point bend test method an external force is applied in the middle of a rectangular-shaped bar (25 3 2 3 2 mm3) specimen until it fractures, and the FS is calculated as: FS 5 3FL/2bd2 where, F is the applied force and L, b, and d represent the distance between two supports, width, and depth of the specimen, respectively (ASTM Standard, 2010). For the disk-shaped specimens the calculated stress value is known as biaxial FS (ASTM, 1996). It is believed that the FS measurement values are higher in three-point than in four-point bending tests (Chitchumnong et al., 1989). f. Fatigue strength is measured when the repetitive application of a small load to a material results in fatigue failure. Fatigue strength is measured by bending or twisting a test specimen and counting the number of repeated stress cycles a material can withstand before it fractures. The fatigue strength gives an indication of ultimate failure when subjected to repeated load over a prolonged period of time. g. Shear strength: The maximum stress that a material can bear prior to failure fracture under shear loading. Punch-test method is a common test to determine the shear strength of dental material. Shear strength 5 F/πdh, where F is the force, d is the diameter of punch, and h is the thickness of the specimen. h. Impact strength is the ability of materials to withstand fracture under a sudden impact force. Impact strength is tested by Charpy impact tester (specimen is held horizontally at the end) or by Izod impact tester (specimen is held vertically at one end) and force is applied by the swinging pendulum. 8. Fracture toughness: The amount of energy required to resist crack propagation in a brittle material under an applied force. In a single-edge-notched test method, a notch of very well-defined dimensions (notch depth and notch tip radius) is created in the specimen to measure fracture toughness. Fracture toughness (K1c) is determined by the applied force, the specimen dimensions, and the size and shape of the notch (ASTM Standard, 2007). 9. Hardness: The ability of a material to resist or withstand indentation or penetration. To measure the hardness of a material, an indenter of known dimension is pressed with a known load for measured length of time and hardness is measured from the area or depth of the indentation. The smaller the indenter area/depth, the higher the hardness. There are various types of indenters used to measure hardness, for example, Birnell, Rockwell, Vickers, and Knoop hardness testers (ISO, 2002). 10. Wear (two body, three body, and wear simulators): In the oral cavity, materials may be subjected to a variety of wear mechanisms including three-body abrasion, attrition, and erosion (ISO, 2001). Wear deals with the relationship of two or more contacting surfaces in relative motion. A number of studies have compared the wear behavior of various dental materials (Zafar and Ahmed, 2014c; Nayyer et al., 2018; Zafar, 2018). The wear
22 Advanced Dental Biomaterials behavior of dental materials is not only dependent on the material properties but also on the contact conditions including the material itself, the surface roughness, motion pattern, the rate of loading, the shape and contour of the antagonist material, the local environment, and lubrication. The traditional methods of classifications of wear are based on the type of motion and the mechanism of wear (adhesion, abrasion, attrition, surface fatigue). In vitro wear test methods attempt to mimic the masticatory processes. The wear assessment of restorative materials has been conducted using a variety of machines which simulate diverse wear mechanisms. Most commonly used wear simulators include the IVOCLAR wear simulator, the Zurich wear simulator, the MTS wear simulator, the Oregon Health & Science University wear simulator, the Dento-munch-robo-simulator, and the University of Alabama wear simulator (Heintze, 2006). However, due to complex oral biology and a limited number of internationally recognized standards for the in vitro wear testing, simulating and interpreting complexity of the wear behavior mechanisms is not straightforward (Heintze et al., 2012). 11. Standards of dental materials testing: Various studies reporting mechanical characterization can only be compared if following certain standards of testing parameters, specimen preparation, and data interpretation. In addition, following standards facilitates the reproducibility of results. Unfortunately, there are only a few standards available to describe the mechanical testing parameters for dental restorative materials, specifically for glass ionomer cements (GICs) and resin-based composites. For example, the specification standard for acidbase cements (ISO 9917-1:2007) describes the CS, working time, and setting time specifications; the standard for light activated system (ISO 9917-2:2000) describes FS, depth of cure, shade, and color stability; the ISO 10477-2004 provides standard specification for bond strength, FS, and water sorption; and the ISO 4049:2000 describes the specification for FS, film thickness, depth of cure, working, and setting time. 2.7 Limitation of mechanical testing methods It is very difficult to duplicate the human oral conditions in the laboratory and this makes it difficult to properly test materials in a similar environment into which they will be placed. In addition, the quality of the test specimens prepared for mechanical testing may also influence the outcome of test results. All mechanical test methods are limited as they address the individual properties without measuring the interaction between various properties. Therefore the phenomenon of aging in the oral cavity and the determination of the potential service of life of dental materials are challenging but an important factor to determine the mechanical properties. 2.8 Biological properties 2.8.1 Biocompatibility Biocompatibility is an ability of a material to perform its desired function without causing any local or systemic adverse response in the recipient of the material (Perrotti et al., 2017; Schmalz, 2014). Biocompatibility is a dynamic process because there is a change in properties of material and host response over the period
Properties of dental biomaterials 23 Table 2.4 Various types of tests used to assess the biocompatibility of biomaterials (Perrotti et al., 2017; Moharamzadeh et al., 2009). Tests Testing methods 1 In vitro 2 In vivo 3 Usage Cytotoxicity test Membrane permeability test Agar overlay method Dentin barrier tests Mucous membrane irritation test Skin sensitization Implantation tests Dental pulp irritation tests Periapical tissue damage and Endodontic usage Dental implantation in bone of time, for instance due to corrosion, disease, and aging (Wataha, 2001). Biocompatibility testing is not easy as new technologies are arriving. The simplest way of classifying these tests is in vitro testing, in vivo testing on animal models, and usage tests (Table 2.4). 2.8.2 In vitro testing In vitro testing is performed outside a living organism in a laboratory setting using cells of human or animal origin (Perrotti et al., 2017; Cintra et al., 2017). The objective of in vitro testing is to predict the response and further improve the material before in vivo testing (de Souza Costa et al., 2014). The in vitro testing includes the following: G G Indirect contact tests, in which the viability of cell is assessed through indirect contact of material, for instance, contact of elute or extract of material with the cells, mostly fibroblasts or keratinocytes (Perrotti et al., 2017; Cintra et al., 2017; Moharamzadeh et al., 2009). Among the other methods of indirect contact are agar overlay method and dentin slice technique (Moharamzadeh et al., 2009). Direct contact tests, in which the material is in direct contact with the cell. The morphological features of cells are observed under a scanning electron microscope (Moharamzadeh et al., 2009; Chan et al., 2017). The cell viability is also assessed by methyl thiazol tetrazolium (MTT) assay in which the viable cells convert the MTT 3-[4,5dimethylthiazol-2-yl]-2,5-diphenyltetrazolium bromide into formazan dye and the dye is quantified using a reader at a specific wavelength (Chan et al., 2017). The in vitro testing has numerous advantages, such as being faster and more cost-effective compared to in vivo testing, with greater reproducibility of the results, and no ethical and legal issues (Sakaguchi and Powers, 2012; Perrotti et al., 2017; Wataha et al., 1999). The disadvantage of in vitro testing is that the exact tissue response of the material cannot be predicted (Sakaguchi and Powers, 2012; Perrotti et al., 2017).
24 Advanced Dental Biomaterials 2.8.3 In vivo testing In vivo testing is done in animal models (de Souza Costa et al., 2014). The in vivo testing includes: G G testing by implantation of the material subcutaneously in animals, followed by monitoring the inflammatory response of the adjacent tissues by excising and examining the tissue under the microscope (Cintra et al., 2017; Garcia Lda et al., 2010; Lacerda-Santos et al., 2015) and implantation test of the material into the experimental animal bone followed by histological evaluation (Chan et al., 2017). The main advantages of in vivo testing is that the biological response is comprehensive and has lower costs compared to clinical studies (Perrotti et al., 2017; de Souza Costa et al., 2014). The disadvantages that are associated with in vivo tests are high costs as compared to in vitro testing, long duration, and ethical and legal issues (Sakaguchi and Powers, 2012; Perrotti et al., 2017). 2.8.4 Usage tests The usage tests are considered as gold standard and are performed on human volunteers. The material or device is placed in a situation that is similar to its intended use (Perrotti et al., 2017; de Souza Costa et al., 2014; Moharamzadeh et al., 2009). The advantage of the usage tests is that the results are clinically relevant and comprehensive (Perrotti et al., 2017; Moharamzadeh et al., 2009). The major disadvantages that are associated with usage tests are their high costs, prolonged duration, and legal and ethical issues (Perrotti et al., 2017; de Souza Costa et al., 2014). It is vital to understand that just one type of testing cannot entirely predict the biocompatibility of a material and a combination of tests is usually carried out to evaluate the overall biocompatibility of the material. For any specific biological application, a systematic protocol (Fig. 2.8) is used to evaluate the biocompatibility of materials or devices. 2.9 Toxicity and cytotoxicity Toxicity is the ability of a material to damage a biological system by chemical means (Schmalz, 2009a). In vitro cytotoxicity testing is assessed during the initial screening of the material. In vitro studies have shown that when the setting reaction in the restorative materials (such as resin modified glass ionomer cements and resin composites) is incomplete, the cytotoxic effects are due to the release of ingredients from the incompletely set material that alters the metabolic processes of the cell (Goldberg, 2008). The toxicity is classified as local or the systemic toxicity. Local toxicity: The damage occurs in the form of inflammation or necrosis at the site of application of a biomaterial (Schmalz, 2009a). Systemic toxicity: The damage occurs in an area that is distant from the site of application of material (Schmalz, 2009a).
Properties of dental biomaterials 25 First step New biomaterial or device In vitro testing of biomaterial or device Second step Promising findings In vivo testing of biomaterial or device Third step Promising findings Usage tests and clinical studies Biocompatible material or device Postmarket evaluation or information system Figure 2.8 Schematic presentation used to evaluate the biocompatibility of materials or devices for biological applications. Toxic and allergic reactions in the oral cavity are rare, due to the following possible reasons (Setcos et al., 2006): G G G Saliva constantly neutralizes and washes away potential allergic and toxic substances. The oral mucosa’s rich vascular supply allows the rapid dispersion of the allergic substances. The stratum corneum layer of the oral mucosa acts as a barrier and prevents the potential allergic and toxic reaction.
26 2.10 Advanced Dental Biomaterials Cytotoxicity tests The cytotoxicity tests assess the cellular reaction by a material or its eluates and in vitro testing assesses the cell viability, cell growth or function, enzyme activity, membrane integrity, and formation of reactive oxygen species before and after exposure of cells to the material (de Souza Costa et al., 2014; Moharamzadeh et al., 2009; Goldberg, 2008). The outcomes of these tests cannot exactly predict the behavior of the material in vivo. The outcomes of cytotoxic tests depend on numerous factors, such as storage time of material before testing, material manipulation/mixing ratio, and polymerization method (Jorge et al., 2003). The test should be selected according to the desired use of the material, for instance the in vitro cytotoxicity of denture adhesives should be evaluated by exposure of extracts of denture adhesives with the gingival fibroblasts, followed by evaluation of the cell density and metabolic activity of the viable cells (Soares et al., 2018). The in vivo local cytotoxicity of oral hygiene products can be assessed by the collection of oral mucosa cells of the individual, followed by scanning of nuclear anomalies under high magnification (Tadin et al., 2018). 2.11 Fluoride and caries Dental caries is one of the most prevalent and multifactorial community problems in the developed as well as developing societies. The relationship of fluoride administration and decreased incidence of dental caries is well documented and fluoride has made an impact on the dental health of the population (Levine, 2010; Tenuta and Cury, 2010). The anticaries effect of fluoride is dependent on the type of fluoride compound, concentration and contact time of fluoride with oral tissues, and the method of fluoride delivery (Goldberg, 2016). Fluoride is an ionic form of a trace element that belongs to the halogen group of elements (McGrady et al., 2010). It is one of the most abundant and highly reactive elements that are present naturally in water, soil, plants, and animals (Harrison, 2005; Ullah et al., 2017). Fluoride is recognized by the US Food and Drug Agency for the prevention of dental caries (Carey, 2014). Fluoride is naturally present in water and seafood (Martı́nez-Mier, 2012). It has been delivered into the oral cavity by water fluoridation, food products (milk, salts, and juices), devices (fluoride nanoparticles), and oral hygiene products (toothpastes, mouth rinses, gels, and varnishes containing fluoride) (McGrady et al., 2010; Ullah and Zafar, 2015). In addition, fluoride has been incorporated into various dental restorative materials such as conventional and resin-modified glass ionomers and modified resin-based composites (Zafar and Ahmed, 2015a; Zafar, 2013) that facilitate caries prevention by releasing fluoride into the oral cavity. The major effect of fluoride on caries prevention is primarily due to the topical effect of fluoride after eruption. The caries preventing effect of fluoride is mainly
Properties of dental biomaterials 27 due to the inhibition of bacterial plaque metabolism and acid production by inhibiting bacterial enzyme enolase (Levine, 2010; Ullah and Zafar, 2015) and through formation of a more stable mineral phase with less solubility. There is also incorporation of fluoride into the enamel prior to the eruption of teeth which also increases the resistance of enamel against dental caries (Martı́nez-Mier, 2012). 2.11.1 Fluoride toxicity The total fluoride dose of 0.050.07 mg/kg body weight is considered as safe (Buzalaf and Levy, 2011). The fluoride containing products for the control of dental caries should be prescribed carefully. During tooth development consumption of fluoride above the safe dose is associated with numerous toxic effects (Tinanoff and O’Sullivan, 1997). However, the primary focus here is dental fluorosis. Dental fluorosis is a developmental disturbance in enamel formation, which occurs because of exposure of fluoride above the safe dose. The enamel is hypomineralized with increased porosity (Carey, 2014; Atia and May, 2013). The dental fluorosis risk is related to the fluoride dose, duration of consumption, nutritional status, renal and gastrointestinal function, genetics, and timing of intake (Tinanoff and O’Sullivan, 1997; Denbesten and Li, 2011). The clinical presentation of dental fluorosis may range from chalky white areas of hypomineralization to dark stained surface with pitting and brittle enamel (Carey, 2014; Tinanoff and O’Sullivan, 1997; Atia and May, 2013). The severity of dental fluorosis depends on when and how long the excess exposure to fluoride occurs (Atia and May, 2013). 2.12 Carcinogenicity Carcinogenesis is the progression of a normal healthy cell to a potentially malignant cell (Scully, 2011). Carcinogenicity is an ability of a material to cause cancer or neoplasm in vivo and the carcinogens are the substances that are known to cause cancer (Craig and Powers, 2002; Anusavice and Phillips, 2003). There are various risk factors that are associated with carcinogenesis; the most commonly associated factors are tobacco, alcohol consumption, Candida albicans, viruses such as human papilloma, herpes simplex, and ionizing radiations. These factors either act alone or in combination (Scully, 2011). Formaldehyde is an antimicrobial compound. It is used in the manufacturing industry (plywood, glue), agriculture (seed treatment), and cosmetics. Beside this it is also present naturally in fruits and fishes. It is also a constituent of dental materials such as composite resins, epoxy resin used in endodontic treatment, and formocresol an agent for pulpotomy (Athanassiadis et al., 2015). Formaldehyde is considered as a carcinogenic material but the amount of release in vivo is very negligible to cause any significant risk (Athanassiadis et al., 2015). Nickel-containing alloys have various applications in dentistry, such as crowns, bridges, partial dentures, metallic brackets, wires, springs, and endodontic instruments (Setcos et al., 2006).
28 Advanced Dental Biomaterials The nickel-containing alloys within restorations and appliances release a very negligible amount of nickel as compared to exposure through other routes, for instance drinking water, occupational exposure, and intravenous exposure during drug administration (Setcos et al., 2006). Although there are concerns of carcinogenicity associated with nickel among laboratory personnel involved in the casting of alloys containing nickel, which can be minimized if manufacturer’s instructions are followed (Setcos et al., 2006). Dental adhesives containing glutaraldehyde, TEGMA, and bis-GMA products have been found to be mutagenic in vitro. However, the concentrations are very negligible in subjects with these restorations compared to those concentrations necessary to cause these effects (Craig and Powers, 2002; Schmalz, 2009b). 2.13 Biodegradation The degradation of material that is not harmful for the body is known as biodegradation. The products that are produced as a result are usually not toxic to the body fluids. However, a high rate of biodegradation can have a few side effects (Goharian, 2017b). Examples of the biodegradable materials include natural polymer-based materials such as silk (Zafar and Al-Samadani, 2014; Cao and Wang, 2009; Vasconcelos et al., 2008), chitosan (Husain et al., 2017; Qasim et al., 2018), and synthetic polymers polylactic coglycolic acid, poly(urethanes), and polylactic acid. The advantages of the biodegradable materials are that their subsequent removal from the patient’s body is not required and it is easier to deliver therapeutic agents locally (Goharian, 2017a). 2.14 Bioactivity Bioactivity is the ability of a material to elicit a specific biological effect and form a bond with the tissue. The term bioactivity is also defined as the cellular effects induced because of release of bioactive substances and ions from the material (Vallittu et al., 2018; Gerhardt and Boccaccini, 2010; Hoppe et al., 2011). The incorporation of therapeutic substances favorably affects the behavior of the cells and enhances bone and new blood vessels formation (Hoppe et al., 2011). Examples of bioactive materials are bioactive glasses for bone regeneration and tissue engineering (Gorustovich et al., 2009), composite materials with bioactive inorganic components (Gerhardt and Boccaccini, 2010), glass ionomers (Najeeb et al., 2016a), hydroxyapatite (HA) nanoparticles modified titanium implants (Vanzillotta et al., 2006; Javed et al., 2014), and titanium dioxide nanoparticle-modified adhesives or coatings (Welch et al., 2010; Najeeb et al., 2016b). Before in vivo testing of bioactivity, it is standard to conduct in vitro testing of the bioactivity of the material by immersion of the material into simulated body fluid and assessing the formation of HA on its surface using different surface
Properties of dental biomaterials 29 analysis characterization techniques such as X-ray diffraction and energy dispersive spectroscopy (Loof et al., 2008). The bioactivity of a material depends on numerous factors, among them the key factors are the chemical composition (presence of bioactive ions for instance Sr, Cu, Zn, F, and Mg), surface roughness or topography of the material, and wettability of the material (Chan et al., 2017; Hoppe et al., 2011). 2.15 Osseointegration It is defined as a direct anchorage or contact between the implant surface and bone without any fibrous tissue in between that can be observed histologically (Albrektsson and Johansson, 2001; De Meurechy et al., 2017). Osseointegration depends on a number of factors, for instance (De Meurechy et al., 2017; Goharian and Abdullah, 2017): G G G material related factors, such as biocompatibility, surface chemistry, roughness, topography, and laser ablation of material; tissue related factors, such as bone quality and quantity; and factors related to surgical procedures, including surgical trauma, sterilization, and primary stability (De Meurechy et al., 2017). Early in vitro assessment of the osteogenic differentiation was evaluated by measuring the in vitro activity of alkaline phosphatase (Chan et al., 2017). Osseointegration is assessed in vivo by implantation of the material within the bone of an animal model followed by micro-CT measurement to evaluate the bone volume density (Chan et al., 2017). Further assessments are done by evaluation of bone implant histologically by assessment of new bone formation and the presence of inflammatory cells (Chan et al., 2017). 2.16 Osteoinduction The process by which the undifferentiated bone cells are stimulated to differentiate into osteoblast cells or the process by which the bone formation is induced (Albrektsson and Johansson, 2001). The materials reported to have osseoinductive effects are poly-hydroxyethylmethacrylate, titanium, HA, biphasic calcium phosphate, carbonated apatite, and octacalcium phosphate (Barradas et al., 2010). Implants with moderate surface roughness 12 μm are considered as optimal for cell attachment. The bioactive glass fibers on the implant surface also improve osseointegration (Chan et al., 2017). 2.17 Foreign body reaction This is an inflammatory process and is marked by the deposition of an avascular dense collagen capsule with inflammatory cells around the material or device that is implanted (Trindade et al., 2016; Soto et al., 2018). These reactions are usually
30 Advanced Dental Biomaterials reported with implant material or devices and less commonly it is associated with local dressing materials (Tasoulas et al., 2017). The inflammatory cells that are mainly present at the site of a foreign body reaction are monocytes, macrophages, Langerhans, and giant cells (Vasiljević et al., 2018; Jordan et al., 2018). These reactions start as a result of an interaction between tissue fluid and blood with the material followed by an inflammatory reaction and granulated tissue formation (De Meurechy et al., 2017). These reactions also lead to the degradation of the materials as a result of which the material becomes brittle and more prone to physical damage (De Meurechy et al., 2017). The foreign body reaction is dependent on the properties of the material, surface form, and topography of the biomaterial (Vasiljević et al., 2018). The strategies that are employed to reduce the foreign body reactions are to decrease surface porosity, incorporation of antiinflammatory agents, and the release and delivery of growth factors (Soto et al., 2018). 2.18 Conclusive remarks The properties of dental materials may vary vastly from one type to another. Understanding the properties of dental biomaterials is important in order for comparison with the properties of the oral hard and soft tissues to be replaced. Ideally, the biomaterials’ properties should closely match the properties of the lost tissues to be replaced. For clinical performance and success, a biomaterial must pass the required criteria for various physical, mechanical, and biological properties during the characterization and testing phase. References Albrektsson, T., Johansson, C., 2001. Osteoinduction, osteoconduction and osseointegration. Eur. Spine J. 10, S96S101. Anusavice, K., Phillips, R., 2003. Phillips’ Science of Dental Materials. Saunders, St. Louis, MO; Great Britain. ASTM, 1996. 394-78, Standard test method for biaxial flexure strength (modulus of rupture) of ceramic substrates. American Society for Testing and Materials, West Conshohocken, PA (reapproved in 1996 and withdrawn in 2001). ASTM, 2010. ASTM 695 Standard Test Method for Compressive Properties of Rigid Plastics. ASTM Materials Standards. ASTM Standard, 2007. Standard Test Methods for Determination of Fracture Toughness of Advanced Ceramics at Ambient Temperature. ASTM Standard, 2010. ASTMD790-10. Standard Test Methods for Flexural Properties of Unreinforced and Reinforced Plastics and Electrical Insulating Materials. ASTM International, West Conshohocken, PA. ASTM Standard, 2014. ASTM D638-14 Standard Test Method for Tensile Properties of Plastics; ASTM International. American Society for Testing and Materials, West Conshohocken, PA.
Properties of dental biomaterials 31 Athanassiadis, B., George, G., Abbott, P., Wash, L., 2015. A review of the effects of formaldehyde release from endodontic materials. Int. Endod. J. 48, 829838. Atia, G., May, J., 2013. Dental fluorosis in the paediatric patient. Dent. Update 40, 836839. Barradas, A.M., Yuan, H., van Blitterswijk, C., Habibovic, P., 2010. Osteoinductive biomaterials: current knowledge of properties, experimental models and biological mechanisms. Eur. Cell. Mater. 21, 407429. Basu, B., Katti, D.S., Kumar, A., 2010. Advanced Biomaterials: Fundamentals, Processing, and Applications. John Wiley & Sons. Braden, M., Clarke, R., 1974. Dielectric properties of zinc oxide-eugenol type cements. J. Dent. Res. 53, 12631267. Brady, A., Lee, H., Orlowski, J.A., 1974. Thermal conductivity studies of composite dental restorative materials. J. Biomed. Mater. Res. 8, 471485. Brauer, G., Termini, D., Burns, C., 1970. Characterization of components of dental materials and components of tooth structure by differential thermal analysis. J. Dent. Res. 49, 100110. Brown, W., Dewey, W., Jacobs, H., 1970. Thermal properties of teeth. J. Dent. Res. 49, 752755. Buzalaf, M.A., Levy, S.M., 2011. Fluoride intake of children: considerations for dental caries and dental fluorosis. Monogr. Oral Sci. 22, 119. Cao, Y., Wang, B., 2009. Biodegradation of silk biomaterials. Int. J. Mol. Sci. 10, 15141524. Carey, C.M., 2014. Focus on fluorides: update on the use of fluoride for the prevention of dental caries. J. Evid. Based Dent. Pract. 14, 95102. Carslaw, H.S., Jaeger, J.C., 1959. Conduction of Heat in Solids, second ed. Clarendon Press, Oxford. Chan, Y., Lew, W., Lu, E., Loretz, T., Lu, L., Lin, C., et al., 2017. An evaluation of the biocompatibility and osseointegration of novel glass fiber reinforced composite implants: in vitro and in vivo studies. Dent. Mater. 34, 470485. Chitchumnong, P., Brooks, S.C., Stafford, G.D., 1989. Comparison of three- and four-point flexural strength testing of denture-base polymers. Dent. Mater. 5, 25. Cintra, L.T.A., Benetti, F., de Azevedo Queiroz, I.O., de Araujo Lopes, J.M., Penha de Oliveira, S.H., Sivieri Araujo, G., et al., 2017. Cytotoxicity, biocompatibility, and biomineralization of the new high-plasticity MTA material. J. Endod. 43, 774778. Combe, E., Moser, J., 1978. The rheological characteristics of elastomeric impression materials. J. Dent. Res. 57, 221226. Costa, M.F., 2016. Color vision sensation and perception. Front Psychol. 7, 1084. Craig, R., Powers, J., 2002. Restorative Dental Materials. Mosby, St. Louis, MO; London. Craig, R., Eick, J., Peyton, F., 1965. Properties of natural waxes used in dentistry. J. Dent. Res. 44, 13081316. Craig, R., Powers, J., Peyton, F., 1967. Differential thermal analysis of commercial and dental waxes. J. Dent. Res. 46, 10901097. De Jager, N., de Kler, M., van der Zel, J.M., 2006. The influence of different core material on the FEA-determined stress distribution in dental crowns. Dent. Mater. 22, 234242. De Meurechy, N., Braem, A., Mommaerts, M., 2017. Biomaterials in temporomandibular joint replacement: current status and future perspectives—a narrative review. Int. J. Oral Maxillofac. Surg. 47, 518533. Denbesten, P., Li, W., 2011. Chronic fluoride toxicity: dental fluorosis. Monogr. Oral Sci. 22, 8196. de Souza Costa, C.A., Hebling, J., Scheffel, D.L., Soares, D.G., Basso, F.G., Ribeiro, A.P., 2014. Methods to evaluate and strategies to improve the biocompatibility of dental materials and operative techniques. Dent. Mater. 30, 769784.
32 Advanced Dental Biomaterials Fairhurst, C.W., Anusavice, K.J., Hashinger, D.T., Ringle, R.D., Warren Twiggs, S., 1980. Thermal expansion of dental alloys and porcelains. J. Biomed. Mater. Res. 14, 435446. Garcia Lda, F., Marques, A.A., Roselino Lde, M., Pires-de-Souza Fde, C., Consani, S., 2010. Biocompatibility evaluation of Epiphany/Resilon root canal filling system in subcutaneous tissue of rats. J. Endod. 36, 110114. Gerhardt, L., Boccaccini, A.R., 2010. Bioactive glass and glass-ceramic scaffolds for bone tissue engineering. Materials 3, 38673910. Goharian, A., 2017a. Further development of trauma plating fixation. In: Goharian, A. (Ed.), Trauma Plating Systems. Elsevier. Goharian, A., 2017b. General concepts. In: Goharian, A. (Ed.), Trauma Plating Systems. Elsevier. Goharian, A., Abdullah, M.R., 2017. 7—Bioinert Metals (Stainless Steel, Titanium, Cobalt Chromium). Trauma Plating System. Elsevier, Amsterdam, pp. 115142. Goldberg, M., 2008. In vitro and in vivo studies on the toxicity of dental resin components: a review. Clin. Oral Investig. 12, 18. Goldberg, M., 2016. Understanding Dental Caries: From Pathogenesis to Prevention and Therapy. Springer. Gorustovich, A.A., Roether, J.A., Boccaccini, A.R., 2009. Effect of bioactive glasses on angiogenesis: a review of in vitro and in vivo evidences. Tissue Eng., B: Rev 16, 199207. Grassmann, H., 1853. Zur theorie der farbenmischung. Annalen der Physik. 165, 6984. Harrison, P.T., 2005. Fluoride in water: a UK perspective. J. Fluorine Chem. 126, 14481456. Heintze, S.D., 2006. How to qualify and validate wear simulation devices and methods. Dent. Mater. 22, 712734. Heintze, S.D., Faouzi, M., Rousson, V., Ozcan, M., 2012. Correlation of wear in vivo and six laboratory wear methods. Dent. Mater. 28, 961973. Herfort, T.W., Gerberich, W.W., Macosko, C.W., Goodkind, R.J., 1977. Viscosity of elastomeric impression materials. J. Prosthet. Dent. 38, 396404. Hoppe, A., Güldal, N.S., Boccaccini, A.R., 2011. A review of the biological response to ionic dissolution products from bioactive glasses and glass-ceramics. Biomaterials 32, 27572774. Husain, S., Al-Samadani, K.H., Najeeb, S., Zafar, M.S., Khurshid, Z., Zohaib, S., et al., 2017. Chitosan biomaterials for current and potential dental applications. Materials 10, 602. International Standards Organization, 2000. ISO 4049 Dentistry-Polymer-Based Filling, Restorative and Luting Materials. International Organization for Standardization, Geneva, Switzerland. ISO, 2001. Dental materials—guidance on testing of wear—Part 2: Wear by two-and/or three body contact. In: Technical Specification; No. 14569-14572. ISO, 2002. ISO14577-1: metallic materials—instrumented indentation test for hardness and materials. In: ISO/TC 164/SC 3. Javed, F., Vohra, F., Zafar, S., Almas, K., 2014. Significance of osteogenic surface coatings on implants to enhance osseointegration under osteoporotic-like conditions. Implant. Dent. 23, 679686. Jordan, S.W., Fligor, J.E., Janes, L.E., Dumanian, G.A., 2018. Implant porosity and the foreign body response. Plast. Reconstr. Surg. 141, 103e112e.
Properties of dental biomaterials 33 Jorge, J.H., Giampaolo, E.T., Machado, A.L., Vergani, C.E., 2003. Cytotoxicity of denture base acrylic resins: a literature review. J. Prosthet. Dent. 90, 190193. Klein, G.A., Meyrath, T., 2010. Industrial Color Physics. Springer. Lacerda-Santos, R., de Meneses, I.H., de Morais Sampaio, G.A., Pithon, M.M., Alves, P.M., 2015. Effect of degree of conversion on in vivo biocompatibility of flowable resin used for bioprotection of mini-implants. Angle Orthod. 86, 157163. Levine, M., 2010. Topics in Dental Biochemistry. Springer Science & Business Media. Lisanti, V., Zander, H., 1950. Thermal conductivity of dentin. J. Dent. Res. 29, 493497. Loof, J., Svahn, F., Jarmar, T., Engqvist, H., Pameijer, C.H., 2008. A comparative study of the bioactivity of three materials for dental applications. Dent. Mater. 24, 653659. López Camelo, A.F., Gómez, P.A., 2004. Comparison of color indexes for tomato ripening. Hortic. Bras. 22, 534537. Martı́nez-Mier, E.A., 2012. Fluoride its metabolism, toxicity, and role in dental health. J. Evid. Based Complement. Alternat. Med. 17, 2832. Mausfeld, R., 1998. 12. Color perception: from Grassmann codes to a dual code for object and illumination colors. In: Color Vision: Perspectives From Different Disciplines. Walter de Gruyter Berlin, New York, p. 231. McCabe, J.F., Walls, A.W.G., 2008. Applied Dental Materials. Blackwell, Oxford. McGrady, M.G., Ellwood, R.P., Pretty, I.A., 2010. Why fluoride? Dent. Update 37, 595598. 601602. Moharamzadeh, K., Brook, I.M., Van Noort, R., 2009. Biocompatibility of resin-based dental materials. Materials 2, 514548. Mumford, J.M., 1967. Resistivity of human enamel and dentine. Arch. Oral Biol. 12, 925927. Najeeb, S., Khurshid, Z., Zafar, M.S., Khan, A.S., Zohaib, S., Martı́, J.M.N., et al., 2016a. Modifications in glass ionomer cements: nano-sized fillers and bioactive nanoceramics. Int. J. Mol. Sci. 17, E1134. Najeeb, S., Khurshid, Z., Zohaib, S., Zafar, M.S., 2016b. Bioactivity and osseointegration of PEEK are inferior to those of titanium—a systematic review. J. Oral Implantol. 42, 512516. Nayyer, M., Zahid, S., Hassan, S.H., Mian, S.A., Mehmood, S., Khan, H.A., et al., 2018. Comparative abrasive wear resistance and surface analysis of dental resin-based materials. Eur. J. Dent. 12, 5766. Perrotti, V., Piattelli, A., Quaranta, A., Gómez-Moreno, G., Iezzi, G., 2017. Biocompatibility of dental biomaterials. In: Shelton, R. (Ed.), Biocompatibility of Dental Biomaterials. Woodhead Publishing. Peyton, F., 1952. Temperature rise and cutting efficiency of rotating instruments. New York DJ. 18, 439. Peyton, F., 1958. Effectiveness of water coolants with rotary cutting instruments. J. Am. Dent. Assoc. 56, 664675. Phillips, L.J., Phillips, R.W., Schnell, R.J., 1955. Measurement of the electric conductivity of dental cement: IV. Extracted human teeth; in vivo tests; summary. J. Dent. Res. 34, 839848. Powers, J., Craig, R., 1974. Penetration of commercial and dental waxes. J. Dent. Res. 53, 402409. Qasim, S.B., Zafar, M.S., Najeeb, S., Khurshid, Z., Shah, A.H., Husain, S., et al., 2018. Electrospinning of chitosan-based solutions for tissue engineering and regenerative medicine. Int. J. Mol. Sci. 19, 407.
34 Advanced Dental Biomaterials Sakaguchi, R.L., Powers, J.M., 2012. Craig’s Restorative Dental Materials. Elsevier/Mosby, Philadelphia, PA. Schmalz, G., 2009a. Biocompatibility of Dental Materials. Springer, Berlin; London. Schmalz, G., 2009b. Resin-based composites. In: Schmalz, G., Arenholt-Bindslev, D. (Eds.), Biocompatibility of Dental Materials. Springer Berlin Heidelberg, Berlin, Heidelberg. Schmalz, G., 2014. Strategies to improve biocompatibility of dental materials. Curr. Oral Health Rep. 1, 222231. Scully, C., 2011. Oral cancer aetiopathogenesis; past, present and future aspects. Med. Oral Patol. Oral Cir. Bucal 16, e306e311. Setcos, J.C., Babaei-Mahani, A., Silvio, L.D., Mjor, I.A., Wilson, N.H., 2006. The safety of nickel containing dental alloys. Dent. Mater. 22, 11631168. Soares, A.S., Scelza, M.Z., Spoladore, J., Gallito, M.A., Oliveira, F., Moraes, R.C.M., et al., 2018. Comparison of primary human gingival fibroblasts from an older and a young donor on the evaluation of cytotoxicity of denture adhesives. J. Appl. Oral Sci. 26, e20160594. Soto, R.J., Merricks, E.P., Bellinger, D.A., Nichols, T.C., Schoenfisch, M.H., 2018. Influence of diabetes on the foreign body response to nitric oxide-releasing implants. Biomaterials 157, 7685. Tadin, A., Gavic, L., Zeravica, A., Ugrin, K., Galic, N., Zeljezic, D., 2018. Assessment of cytotoxic and genotoxic effects of conventional and whitening kinds of toothpaste on oral mucosa cells. Acta Odontol. Scand. 76, 6470. Tasoulas, J., Daskalopoulos, A., Droukas, C., Nonni, A., Nikitakis, N.G., 2017. An unusual microscopic pattern of foreign body reaction as a complication of dry socket management. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. 125, e118e123. Tay, W., Braden, M., 1981. Materials science: dielectric properties of glass ionomer cements. J. Dent. Res. 60, 13111314. Tay, W., Braden, M., 1987. Thermal diffusivity of glass-ionomer cements. J. Dent. Res. 66, 10401043. Tenuta, L.M.A., Cury, J.A., 2010. Fluoride: its role in dentistry. Braz. Oral Res. 24, 917. Tillberg, A., Jarvholm, B., Berglund, A., 2008. Risks with dental materials. Dent. Mater. 24, 940943. Tinanoff, N., O’Sullivan, D., 1997. Early childhood caries: overview and recent findings. Pediatr. Dent. 19, 1216. Trindade, R., Albrektsson, T., Tengvall, P., Wennerberg, A., 2016. Foreign body reaction to biomaterials: on mechanisms for buildup and breakdown of osseointegration. Clin. Implant. Dent. Relat. Res. 18, 192203. Ullah, R., Zafar, M.S., 2015. Oral and dental delivery of fluoride: a review. Fluoride 48, 195204. Ullah, R., Sohail Zafar, M., Shahani, N., 2017. Potential fluoride toxicity from oral medicaments: a review. Iran. J. Basic Med. Sci. 20, 841848. Vaderhobli, R.M., 2011. Advances in dental materials. Dent. Clin. North Am. 55, 619625. x. Vallittu, P.K., Boccaccini, A.R., Hupa, L., Watts, D.C., 2018. Bioactive dental materials—do they exist and what does bioactivity mean? Dent. Mater. 34, 693694. Van Noort, R., Barbour, M.E., 2013. Introduction to Dental Materials 4. Elsevier Health Sciences. Vanzillotta, P.S., Sader, M.S., Bastos, I.N., Soares Gde, A., 2006. Improvement of in vitro titanium bioactivity by three different surface treatments. Dent. Mater. 22, 275282.
Properties of dental biomaterials 35 Vasconcelos, A., Freddi, G., Cavaco-Paulo, A., 2008. Biodegradable materials based on silk fibroin and keratin. Biomacromolecules 9, 12991305. Vasiljević, P.J., Živković, J., Vukelić-Nikolić, M., Najman, S., 2018. Determining the biological properties of biomaterials in vivo. Anonymous Biomaterials in Clinical Practice. Springer. Vermilyea, S., Powers, J., Koran, A., 1978. The rheological properties of fluid denture-base resins. J. Dent. Res. 57, 227231. Wataha, J.C., 2001. Principles of biocompatibility for dental practitioners. J. Prosthet. Dent. 86, 203209. Wataha, J., Lockwood, P., Nelson, S., Rakich, D., 1999. In vitro cytotoxicity of dental casting alloys over 8 months. J. Oral Rehabil. 26, 379387. Welch, K., Cai, Y., Engqvist, H., Stromme, M., 2010. Dental adhesives with bioactive and on-demand bactericidal properties. Dent. Mater. 26, 491499. Wilson, T., Turner, D., 1987. Characterization of polydimethacrylates and their composites by dynamic mechanical analysis. J. Dent. Res. 66, 10321035. Zafar, M.S., 2013. Effects of surface pre-reacted glass particles on fluoride release of dental restorative materials. World Appl. Sci. J. 28, 457462. Zafar, M.S., 2014. A comparison of dental restorative materials and mineralized dental tissues for surface nanomechanical properties. Life Sci. J. 11, 1924. Zafar, M.S., 2018. Wear behavior of various dental restorative materials. Mater. Technol. 33, 2531. Zafar, M.S., Ahmed, N., 2014a. Nanoindentation and surface roughness profilometry of poly methyl methacrylate denture base materials. Technol. Health Care 22, 573581. Zafar, M.S., Ahmed, N., 2014b. Nanomechanical characterization of exfoliated and retained deciduous incisors. Technol. Health Care 22, 785793. Zafar, M.S., Ahmed, N., 2014c. Effects of wear on hardness and stiffness of restorative dental materials. Life Sci. J. 11, 1118. Zafar, M.S., Al-Samadani, K.H., 2014. Potential use of natural silk for bio-dental applications. J. Taibah Univ. Med. Sci. 9, 171177. Zafar, M.S., Ahmed, N., 2015a. Therapeutic roles of fluoride released from restorative dental materials. Fluoride 48, 184194. Zafar, M.S., Ahmed, N., 2015b. The effects of acid etching time on surface mechanical properties of dental hard tissues. Dent. Mater. J. 34, 315320.
3 Dental gypsum and investments Touraj Nejatian1,2, Pegah Firouzmanesh3 and Azeem Ajaz4 1 Eastman Dental Institute, University College of London, London, United Kingdom, 2 Nottingham Dental Clinic, Nottingham, United Kingdom, 3Babol University of Medical Sciences, Babol, Iran, 4Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Kingdom of Saudi Arabia Chapter Outline 3.1 3.2 3.3 3.4 3.5 Introduction 37 Desirable properties of gypsum products 38 Production of calcium sulfate hemihydrate 38 Types of gypsum products 40 The setting and manipulation characteristics of gypsum products 3.5.1 3.5.2 3.5.3 3.5.4 42 Mixing technique 42 Pouring the impression 43 The setting processes 44 Setting time 46 3.6 Setting expansion hygroscopic setting expansion 47 3.6.1 3.6.2 3.6.3 3.6.4 3.6.5 Reproduction of detail 49 Compressive strength 49 Tensile strength 50 Surface hardness and abrasion resistance 50 Dimensional stability 51 3.7 Dies and models produced from digital data 3.8 Conclusion 52 References 52 3.1 51 Introduction Replicas of oral tissues are used in dentistry for the purpose of assessment, treatment planning, and manufacturing indirect restorations. They may also be used as part of medicolegal records. These replicas are called Study Models when used for assessment and treatment planning, whereas the ones used for manufacturing removable appliances are classified as Casts. Dies are the third type and are more accurate replicas which are used for making fixed restorations. The accuracy of the replicas depends on the accuracy of both the impression and the replica materials. Various materials are used to pour the impressions including gypsum products and epoxy materials. Recently, polymers are being used by novel three-dimensional Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00003-7 Copyright © 2019 Elsevier Ltd. All rights reserved.
38 Advanced Dental Biomaterials (3D) printing systems to make replicas. Gypsum products are the most commonly used material to pour the impressions due to their versatile nature. There are five types of gypsum products with the same chemical composition but differing physical forms with different properties for different applications. 3.2 Desirable properties of gypsum products A true replica of oral tissue is required to be reasonably accurate and dimensionally stable to enable making an accurate restoration; however, the degree of accuracy required varies depending on the application. For example, indirect fixed restoration should be made with high accuracy and fine details, whereas a moderate level of accuracy is acceptable for most of the study models. Strength, hardness, and resistance against scratching are other important properties of gypsum materials, especially for die materials as they are prone to damage during handling in a dental lab. Compatibility with impression material is another requirement of the replica materials as poor compatibility may result in imperfection in the replicas. Finally, like many other materials in the clinical environment, gypsums should have a long shelf life, short setting time, be easy and safe to use, inexpensive, and be easy to dispose of or recycle. Gypsum products of five different types have a range of properties to meet the various clinical needs, which have made it the most popular replica material in dentistry. For gypsum investment materials, the desirable properties should be same as above along with some more detailed features (Luk and Darvell, 1991; Anusavice, 2013; Asaoka et al., 2012; Luk and Darvell, 2003): G G G G G G G G G G Sufficient strength at room temperature and high temperatures to withstand the impact of molten metal. Easy to manipulate and become hard in a reasonably short period of time. Stable and resistant to higher temperatures to prevent damage to the surface of the alloy by releasing gas. Sufficient expansion to compensate the shrinkage of the wax pattern and metal alloys. Beneficial thermal expansion over a range of other casting temperatures. Enough porosity to let off the gases produced during casting procedures. Preserve all the fine details and margins. Easily separated from the surface of the metal. No chemical reaction with casting metals. Inexpensive. 3.3 Production of calcium sulfate hemihydrate The main constituent of gypsum-based products, including plasters and dental stones, is calcium sulfate hemihydrate (CaSO4 1/2H2O) or [2(CaSO4) H2O]. As gypsum is in the form of dihydrate calcium sulfate (CaSO4 2H2O), in order to
Dental gypsum and investments 39 produce calcium sulfate hemihydrate, first gypsum is crushed and milled and then exposed to heat at 110 C 120 C (230 F 250 F) in order to lose three-fourths of its water. Losing this amount of water is essential for the conversion of CaSO4 2H2O to CaSO4 1/2H2O (calcium sulfate hemihydrate). This process is the first stage of the calcination reaction (Table 3.1). Based on the calcination method, different types of hemihydrate may be produced. These types are known as α-hemihydrate and β-hemihydrate. Application of the α and β prefixes does not represent different substance phases, rather these prefixes are used only due to tradition and their easy use. Indeed, the α and β types are not different in terms of crystal shape, rather their difference lies in the crystal size, surface area, and the extent of network perfectness. If gypsum is exposed up to 120 C under dry calcination in an open space, a hemihydrate crystalline species is produced called β-hemihydrate. The β form, also known as dental plaster, consists of large crystals and irregular orthorhombic particles with capillary pores. If gypsum is heated under pressure in the presence of water vapor up to around 125 C, α-hemihydrate is formed. The α form, also known as dental stone, is composed of crystal particles with a smaller size and more irregular prismatic shape. The chemical reaction during the manufacturing and setting reaction of all types of gypsums is exactly the same (Fig. 3.1). When α-hemihydrate is mixed with water, reaction 1 (Fig. 3.1) is reversed. The product is far stronger and harder than the one obtained from mixing β-hemihydrate with water. The main reason for this difference is that β-hemihydrate particles absorb more water, as their crystals have a more irregular shape and are more porous. Although the particle size and sum of the surface area are the main factors in measuring the water content required for mixing, particles with different sizes should be scattered uniformly in the mixture. Pulverizing the particles after hemihydrate preparation removes needle-like crystals, causing the crystals to stand together more densely, thereby decreasing the level of water required for mixing. Attachment and adhesion among hemihydrate particles are also important factors in determining the water required for pourable products. Adding trace amounts of some materials with an active surface such as Arabic Gum along with calcium carbonate can considerably decrease the extent of water required in gypsum products. Therefore different gypsum products need different amounts of water and the extent of this need can be estimated based on the shape and density of crystals. It is worth Table 3.1 Calcination process for gypsum. CaSO4.2H2 O gypsum (calcium sulfate dihydrate) 110120°c 130200°c CaSO4.1/2H2O Plaster or stone (calcium sulfate hemihydrate) 2001000° c CaSO4 Hexagonal anhydrate CaSO4 (orthorhombic anhydrite)
40 Advanced Dental Biomaterials Figure 3.1 Chemical reaction of gypsum during (1) manufacturing (gray box) and (2) setting process (reverse reaction). noting that these factors are adjusted and calculated by the manufacturer company. Modified α-hemihydrate is obtained through heating gypsum in an aqueous solution of calcium chloride 30%. These particles are denser than the two previous types and need less water for mixing. The obtained powder is used in synthesized dyes. Modified α-hemihydrate is also known as die stone or modified stone (Worner, 1944). 3.4 Types of gypsum products Selecting a particular type of gypsum product depends on its use and the required physical properties for that purpose. Gypsum products are categorized into following five types (Fig. 3.2): 1. Type I: Impression plaster Today, impression plasters are rarely used. Their uses include initial molding from toothless ridges to make denture (mucostatic impression). 2. Type II: Model plaster Model plaster or Type II is a type of β-hemihydrate, which is used to prepare initial diagnostic and after-treatment models, fill in the flasks when making dentures, and mount the casts in the articulator. With low compressive strength and tensile strength, model plaster is considered a relatively weak substance. 3. Type III: Dental stone In 1930 with the discovery and introduction of α-gypsum into dentistry, a new area began. The fabricated gypsum, made under vapor pressure in a closed-lid container, results in a slower outflow of water off the crystals and the development of powdered particles with a uniform shape and less porosity. This substance is harder than the β type.
Dental gypsum and investments 41 Figure 3.2 Different types of gypsum products. Thus in the vicinity of hydrocolloid molding materials, it loses its superficial strength to a lesser extent and allows for fabricating indirect pattern. Dental stone is suitable for synthesizing acrylic dentures and making diagnostic casts. 4. Type IV: Dental stone, high strength Stone Type IV is a modified α-hemihydrate which is obtained by calcination of gypsum in calcium chloride solution and has very dense and cubic-shaped particles. They have a great strength, hardness, and resistance to abrasion, while their setting expansion is low. As a result, they are suitable for preparing die. 5. Type V: Dental stone, high strength, high expansion Dental stone Type V has the maximum compressive strength. In this type, expansion occurs during setting leading to a larger mold, which compensates for the solidification shrinkage of base-metal alloys during the later casting (Mccabe and Walls, 2013) (Fig. 3.3).
42 Advanced Dental Biomaterials Figure 3.3 Setting expansion of Type V gypsum utilized to compensate for solidification shrinkage of base-metal alloys during casting. 3.5 The setting and manipulation characteristics of gypsum products The technique to use gypsum products is relatively simple. The required equipment includes mixing bowl, mixing spatula, room-temperature water, and correct powder/water ratio. To achieve the optimum properties powder and water should be mixed with each other at the correct ratio. It is important that the mixing bowl and other mixing equipment used for dental gypsum be used only for mixing dental gypsum. Use of this equipment for mixing phosphate investment or molding substances causes contamination that may alter the dental stone properties later. It is also important that all mixing equipment be cleaned after the job and no dry gypsum remains on the bowl or spatula. 3.5.1 Mixing technique 3.5.1.1 Measuring the water The best outcome is obtained by distilled water, though mixing with normal tap water also leads to favorable results. The water should be at ambient temperature. Water is usually measured by a graduated cylinder, as 1 g water has almost a volume of 1 mL. 3.5.1.2 Measuring the powder The powder weight is measured by an accurate scale. The powder volume should not be measured by a cup, as different powders become dense at different levels.
Dental gypsum and investments 43 3.5.1.3 Adding powder and water The preferred method for mixing involves adding the measured level of water into the container. Then, the preweighed powder should be gradually added to it. Approximation-based addition of water and gypsum for reaching the desired consolidation should be avoided, as it can result in a low strength and inconsistent expansion. Gypsum should be added to water slowly to avoid entrapment of large amounts of air. The powder should be left to soak in water for at least 30 seconds while moving the bowl around to wet the powder before mixing. Using hand mixing, back and forth strokes are used. Typically, the bowl is rotated so that all nonmixed and dry substances enter the mixture. This continues until a creamy homogenous mixture is obtained. Mechanical mixing accelerates the setting reaction and reduces work and setting time of gypsums. The strength and hardness obtained under conditions where mixing is done mechanically and in a vacuum are usually greater than when mixing has been done by hand. As mechanical mixers are faster, the powder is added to water and it is stirred by hand until complete soaking of powder. It is then placed in a mechanical mixer and is mixed for a shorter time of 10 30 seconds (Mccabe and Walls, 2013). 3.5.2 Pouring the impression It begins by holding the impression tray so that the handle rests against the vibrator. Start at one end of the arch and place a small amount of gypsum mixture into the impression letting it flow slowly enough to fill each tooth imprint one after another. After the impression is filled, the remaining bubbles should be popped with a small instrument. After covering all the critical surfaces of the impression, you may safely add progressively the larger amount of the mix. Do not flow stone over the outside of the tray, because it must be removed before the impression can be separated from the cast. Now let this first pour set for about 45 minutes or at least until the stone loses its glazed appearance before making the cast base. Continue filling the impression stone to the level slightly above the height of the impression walls (about 2 mm thick). Lastly, add retention nodules to this first pour to retain the base mechanically. Place the handle of the tray in a holding device; do not lay the impression on the counter or the cast will be distorted. Add stone around the retention nodules and invert the first pour into the patty. Use your spatula to shape the sides of the cast base. Take care not to bring stone up onto the tray embedding the tray into the base. With a mandibular cast, you must take time to smooth and contour the tongue space while the second mix is still soft. There is a rate of vibration that is best for each mix’s ability to flow. The vibration intensity should be set high enough to make the material move across the surface of the impression. If the vibrator is set too high the impression “jumps” in your hand and gypsum moves so fast that it skips over surface detail, or vibration wave patterns develop on the surface of the mix (O’Brien, 2008).
44 Advanced Dental Biomaterials 3.5.3 The setting processes The precursor substance for production of casts and gypsum models and some casting investments and molding gypsum is calcium sulfate hemihydrate, which is converted to calcium sulfate dehydrate through an exothermic reaction when setting (Fig. 3.4). Three theories of gypsum setting exist: 1. crystalline 5 dissolution 2 precipitation; 2. gel; and 3. hydration. Nowadays, the most accepted theory is crystalline theory, which is based on the dissolution of hemihydrate powder and instant recrystallization of gypsum followed by the interlocking of the crystal to form the set solid. Calcium sulfate dehydrate has a relatively low solubility compared to calcium sulfate hemihydrate, which causes the setting of these substances. When hemihydrate powder is mixed with a proper ratio of water, it changes into a thick slurry. Dissolution of hemihydrate continues until it turns into a saturated solution. Thereafter, the saturated solution of hemihydrate mixes with water and converts to supersaturated dehydrate mixture and precipitates as needle-like crystals of nucleation type in the solution. At room temperature (20 C), the solubility of hemihydrate is four times greater than that of dehydrate forms. With the precipitation of calcium sulfate dehydrate, the solution is no longer saturated with hemihydrate. Thus hemihydrate continues to dissolve. Next, while new crystals are being formed or previous crystals are growing further, the dissolution of hemihydrate and precipitation of the dehydrate form continues. The links of the needle-like crystals which have been distorted cause the final consistent structure (Docking, 1965; Winkler et al., 1998). 3.5.3.1 Stages of setting The setting process is continuous, from the beginning of mixing until the setting reaction is complete, by which time the material has reached its full wet strength. However, important physical changes can be recognized during this process. The stages in setting may be described as fluid, plastic, friable, and carvable. 3.5.3.2 The rate of setting reaction G G G G G Water/powder ratio (W/P ratio) Spatulation Temperature Colloidal system and pH Additives: (1) accelerates and (2) retarders Figure 3.4 Setting reaction of calcium sulfate hemihydrate.
Dental gypsum and investments 45 3.5.3.3 Water/powder ratio In general, the effects of using less water (less W/P ratio) can be summarized as follows: G G G G G thickening of the mixture and making it harder to work; retaining more air bubbles by reducing the amount of water to fill the space between particles; strengthening of the mixture; enhancing the dimensional changes; and decreasing the number of nuclei per unit resulting in a longer setting time. 3.5.3.4 Spatulation The setting time will be shorter by prolonging and increasing the rate of gypsum mixing. This is probably due to the fact that when the stone or plasters are in contact with water, some gypsum crystals will form immediately. The formation of these crystals increases with the onset of mixing. At the same time, the crystals are broken down through spatula stirrer and dispersed within the mixture, causing an increase in the number of crystalline nuclei and thus shortening the setting time (Mccabe and Walls, 2013). 3.5.3.5 Temperature Increasing temperature has complex effects on the setting time of gypsum products, including a decrease in the relative solubility of calcium sulfate hemihydrate and calcium sulfate dihydrate. Reducing the lower solubility ratio leads to a slower reaction and longer setting times. On the other hand, increasing the temperature will accelerate the reaction rate and shorten the setting time. In practice the effects of these two phenomena accumulate together, to reveal the overall effect. Increasing the temperature to about 40 C enhances the solubility of hemihydrate and thus accelerates the setting reaction. Increasing the temperature over 40 C leads to a decrease in the solubility of hemihydrate and slows down the setting reaction until the solubility of hemihydrate and dihydrate is equal at 100 C, where no reaction occurs and the plasters will not set (Singh and Garg, 1997). 3.5.3.6 Modifying agents One of the methods of controlling the setting time is adding certain modifiers to the dental stone and plaster mix. If the added chemical reduces the setting time, it will be known as an accelerator, and if it increases the setting time, it will be considered as a retarder. The retarders usually reduce the solubility of a layer absorbed on the surface of the hemihydrate and prevent further growth of the gypsum crystals by coating their surface. The organic materials such as glues, gelatin, and some gums act in this manner. Other types of retarders consist of salts forming a layer of calcium salt on the plaster surface, which is less soluble in comparison to sulfate, including borax, potassium citrate, and calcium chloride. It should be noted that many alkaline mineral salts, such as sodium chloride, act as
46 Advanced Dental Biomaterials accelerators in low concentrations but as retarders in high concentrations (more than 20%). Less than 2% potassium sulfate or sodium chloride can be added to the gypsum mixture to accelerate the setting reaction rate and reduce the setting time. The gypsum is usually appended by adding a small amount of slurry to the water, but not a lot of powder to accelerate the reaction because it will be very difficult to produce a mixture with the desired viscosity. It is better to choose a product that essentially has longer setting time when a prolonged setting time is required. It should be kept in mind that the use of colder water has no major effect on the setting time. To ensure achieving the longest working time, it is better, therefore, to use a product with a regular setting time or to use the shortest time and rate of mixing. 3.5.3.7 Fineness Minimizing the size of hemihydrate particles leads to an increase in the hemihydrate dissolution rate, the number of gypsum nuclei and setting rate. 3.5.3.8 Effect of pH Fluids with a low pH (such as saliva) delay the setting reaction (Earnshaw and Marks, 1964). If the mold taken from the patient’s mouth is not washed properly and is still contaminated with saliva, it will slow down the setting reaction. 3.5.4 Setting time 3.5.4.1 Initial setting time Initial setting time or working time is the interval between the onset of mixing and the setting mass reaching the semihard stage. Initial setting time is clinically the period when the plaster loses its gloss. 3.5.4.2 Final setting time This is the interval between the start of mixing the powder and water to the completion of the substance reaction and setting, when the material can be removed from the mold without deformation or break. The final setting time is chemically the period when all water required for hydration is consumed by hemihydrate. The following tests are used to measure the setting time: G G Loss of gloss As the reaction progresses, part of the excess water is absorbed and the dihydrate is produced; as a result, the mixture loses its gloss. This happens usually in the ninth minute, but the plaster setting is still incomplete and cannot be removed from the mold. Initial Gilmore test In the initial Gilmore test the mixture is spread over the target area and the needle is placed on the surface. The initial setting time will be when the needle no longer makes any track on the material surface.
Dental gypsum and investments G 47 Vicat test for setting time A needle with a diameter of 1 mm and a pressure of 300 g is used in this test. The setting time is obtained when the needle no longer penetrates to a depth of 2 mm in the material. This test can be also employed to control the quality of the mixture. The measurement of Vicat and initial Gilmore occur simultaneously in some cases, while there is a little difference at other times. 3.5.4.3 Gilmore test for final setting time This is similar to the initial Gilmore test, but a heavier Gilmore needle is used. The final setting time will be when the needle leaves an only minor barely visible trace on the surface (Dental Science, 2017; Sanad et al., 1982; Williams and Williams, 1984). 3.6 Setting expansion hygroscopic setting expansion Gypsum products, when mixed with water and allowed to set, exhibit measurable linear expansion along their extremities, which is known as “normal setting expansion.” When during setting additional water is brought into contact, an increased expansion is observed due to the outward thrust of the growing crystals. This phenomenon of expansion is known as “hygroscopic setting expansion” (Mahler and Bruce Ady, 1960). All gypsum products show some linear expansion, the percentage of this expansion varies from one type of gypsum product to another (Table 3.2). Under normal conditions the round particles of hemihydrate are surrounded by water and the crystals of dihydrate begin to form on the nuclei (Mahler and Bruce Ady, 1960; Lyon et al., 1955). Hydration helps reduce water around the particles while the surface tension of water draws the particles closer. The crystals of dihydrate begin to grow and contact each other, further decreasing the water around the particles. Furthermore, water is also consumed by dihydrate growth which should bring the crystals closer as before, but the outward thrust of the growing crystals repels this contraction and ultimately the crystals are entangled and intermeshed (Mahler and Bruce Ady, 1960; Lyon et al., 1955). Consequently, gypsum products when setting under water have been observed with more setting expansion than gypsum products which set under air (Fig. 3.5). Table 3.2 Setting expansion by volumetric percentage of different gypsum products when mixed with water. Gypsum products Setting expansion (%) Plaster Low moderate strength dental stone High-strength dental stone High-strength, high-expansion dental stone 0.2 0.3 0.15 0.25 0.08 0.10 0.10 0.20
48 Advanced Dental Biomaterials I Water 1 2 II 3 Water Water Water 1 2 3 Figure 3.5 Illustration of the setting expansion of gypsum products. (I) The gypsum product is mixed with water and left to set. The crystal growth is restricted by the surface tension of water surrounding the growing crystals. (II) Hygroscopic expansion occurred when the mixture is exposed to water during the setting reaction which provides room for longer crystal growth (Mahler and Bruce Ady, 1960). The hygroscopic setting expansion is purely a physical phenomenon and is not a result of any chemical reaction other than the normal setting expansion. Therefore variables that will have an effect on normal setting expansion will be liable for changes in hygroscopic expansion as well. For example, the higher hygroscopic expansion will be a result of a reduction in W/P ratio and increased speculation (Cesar et al., 2015; Cunningham et al., 1952; Garber et al., 1985). The hygroscopic expansion achieved during the setting of dental plaster or stone is double the time compared to the normal setting expansion. The higher hygroscopic expansion is required for gypsum-bonded casting investments used in the fabrication of cast gold alloy restorations. Other investment materials have also shown a similar expansion pattern when they set under hydrated conditions (Michalakis et al., 2009). It is important to note that hygroscopy is the ability of a substance to attract water molecules from the environment either by absorption or adsorption. For gypsum products, the hygroscopic expansion that has been described here is not a result of absorption or adsorption and, therefore, the term hygroscopic setting expansion is technically incorrect (Mahler and Bruce Ady, 1960; Michalakis et al., 2009). Nevertheless this term will be used whenever explaining the setting expansion of gypsum.
Dental gypsum and investments 49 3.6.1 Reproduction of detail The American National Standards Institute/American Dental Association Specification nos. 19 and 25 evaluate the compatibility of impression materials with dental stone by the presence of a 20-μm-wide line produced on a dental stone cast of calcium sulfate dehydrate (ADA, 1972, 1977). Surface details less than 20 μm are not recorded due to the porous nature of gypsum products, but macroscopic details are very well reproduced. There is some loss of surface details contributed by the production of air bubbles. It is advisable to use surface vibration and tapping techniques while pouring a cast to reduce air bubbles. Since the gypsum is soluble in water in minute degrees, care must be taken to avoid repeated drying and wetting of the surface when waxing up a mold for casting (Malaviya and Shrestha, 2016). It has been well established that gypsum dies do not reproduce surface details better than electroformed or epoxy dies due to their surface being porous at the microscopic level. There have been surface details record issues when silicon impressions are used as the gypsum does not wet these materials (Chang et al., 2016). Therefore nonionic surfactants are used to improve the wettability of these impression materials by water slurry. Contamination with saliva or blood can also affect surface detail reproduction. Therefore care must be taken when pouring impressions with gypsum to achieve good surface details. 3.6.2 Compressive strength The most common mechanical property for assessing the strength of gypsum products is the compressive strength. The compressive strength decreases from high-strength dental stone at 38 MPa to model stone at 30 MPa and down to plaster at 12 MPa (Noort and Barbour, 2013). The compressive strength is inversely proportional to the W/P ratio. The greater the amount of water used to mix, the lower is the compressive strength. Model plaster with a greater quantity of excess water exhibits a compressive strength of 12 MPa, whereas high-strength dental stone consisting of the least amount of excess water has a compressive strength around 38 MPa. The excess water only contributes to the volume and not the strength of the material (Earnshaw and Smith, 2009; Peyton et al., 1952; Noort, 2002). For example, if the stone is mixed with the same W/P ratio as model plaster its strength will be decreased to similar to that of the model plaster (Fig. 3.6). Thus the compressive strength of the set material is affected significantly if the recommended powder liquid ratio is not used. More water has the advantage of a uniform mix and air bubbles can be easily released by vibration action, but compressive strength becomes inferior. On the other hand, using less water makes a thick mix incorporated with air which in turn increases the porosity and significantly reduces the overall strength. Therefore it is highly recommended to follow the manufacturer’s guidelines. In general, a dry mix may increase the compressive strength up to twice that of the wet one, but the excessive reduction of water will result in inferior properties (Earnshaw and Smith, 2009; Peyton et al., 1952; Noort, 2002).
50 Advanced Dental Biomaterials Figure 3.6 Effect of W/P ratio on compressive strength of different gypsum products: plaster, model stone, and high-strength dental stone. W/P ratio, Water/powder ratio. Source: Data from Ronald, L.S., John, M.P., 2012. Craig’s Restorative Dental Materials, 13th ed. Mosby Elsevier. 3.6.3 Tensile strength Plaster’s tensile strength is very low (approximately 2 MPa) due to the porous and brittle nature of the material. This has the disadvantage of the teeth and margins being damaged if handled roughly. This is why dental stone, which has twice the tensile strength of plaster is preferred for the production of crown and bridge models and dies (Earnshaw and Smith, 2009; Padevět et al., 2011). 3.6.4 Surface hardness and abrasion resistance Gypsum products are highly susceptible to scratching and loss of surface details due to low surface hardness and abrasion resistance. The surface hardness is directly proportional to the compressive strength of the gypsum products. Therefore the greater the compressive strength, the greater the hardness. It has been noted that surface hardness increases at a much faster rate than the compressive strength for a final set material due to the surface mass of the hardened material reaching a dry state earlier than the mass of the inner portion (Peyton et al., 1952). Gypsum products have been tested again and again and attempts have been made to improve their hardness by incorporating set gypsum with epoxy or methylmethacrylate monomer that polymerizes when setting (Robert et al., 2004; Williams et al., 1983). Epoxy resin does show better detail reproduction, abrasion resistance, and transverse strength than gypsum products but undergoes polymerization shrinkage. This, in turn, can affect and compromise the fit of the castings unless taken into account before processing. Dental plaster with epoxy resin showed an increase in the strength but this was not the case for dental stone or high-strength dental stones (Paquette et al., 2000; Robert et al., 2004). Generally, abrasive resistance was improved by 15% 41% when epoxy resin was added to dental stone but compressive strength and surface hardness decreased. Drying molds, casts, or dies to achieve compressive strength and surface hardness rapidly has been rejected as it makes the gypsum dehydrated and therefore weakens the overall structure. Some other methods, such as soaking in glycerin or
Dental gypsum and investments 51 different oils that have been used to increase surface hardness, have also been rejected as this would not improve the surface hardness and would just smoothen the surface which makes the carver or wax knife slide over it when in contact. Surface hardness can be improved and increased using commercial hardening solutions containing colloidal silica (30%) (Robert et al., 2004; Williams et al., 1983). However, it should be noted that increased surface hardness does not mean increased abrasion resistance as that is only one of the many factors that can affect wear resistance. Studies for two-body abrasion resistance showed that there is no improvement in abrasion resistance when high-strength stone is commercially hardened using different solutions (Robert et al., 2004; Hollenback and Smith, 1967; von Fraunhofer and Spiers, 1983). Further studies and methods for the measurement of abrasion resistance are required as two-body abrasion tests are not well established for gypsum products. It is worth noting that disinfectants used on gypsum products have shown to adversely affect the surface hardness making the surface susceptible to erosion and wear. 3.6.5 Dimensional stability Dimensional stability of the set gypsum products is excellent; there is no or very little dimensional change of the set material (Chan and Darvell, 2001). These materials can be stored for a long period of time under dry conditions as the material is still slightly soluble in water. Therefore it is recommended not to wash the surface of the set material with hot water. 3.7 Dies and models produced from digital data Computer-assisted technologies such as 3D technologies have been increasingly used in dentistry, where a layer-by-layer replica or a designed restoration is formed. This is so-called additive manufacturing or rapid prototyping employed to make dental replicas as well as restorations. To manufacture the replicas a digital image of the oral tissue or an object is obtained either directly from the patient’s mouth or indirectly from a conventional impression and then sent to a 3D printer after processing the data. In order to make restorations, a virtual restoration is designed on the virtual model in the computer and sent to a 3D printer. This technology facilitates and accelerates manufacturing of replicas and restorations with higher accuracy and reduces material waste. However, the high initial investment makes adoption of this technology by dental practices or most dental labs very slow. Different materials have been used for 3D printing including ABS plastics, polyamide (nylon), glass-reinforced polyamide, epoxy resins, titanium, steel, photopolymers, waxes, and polycarbonates. Formlabs has recently introduced a long-term biocompatible 3D printing resin called Dental LT Clear for orthodontic appliances and devices (TESS, 2017). PolyJest dental materials are specially designed for dental application, and these include three types (Table 3.1).
52 Advanced Dental Biomaterials 1. Clear Biocompatible (MED 610) Used to produce orthodontic appliances as well as delivery and positioning trays and surgical guides. 2. Vero Dent Plus (MED 690) A dark beige material that creates amazingly fine features and finishes with great accuracy, strength, and durability. 3. Vero Glaze (MED620) An opaque material with A2 shading designed to provide the suitable color match for veneer try-ins and diagnostic wax-ups (Mahamood et al., 2016). These materials are used by various systems such as Object260 and Object500 Dental Selection (by Stratasys) to manufacture dental replicas and restorations. 3.8 Conclusion Gypsum products have been used in dentistry for many years, mostly to make replicas of oral tissue for various purposes, due to their desirable properties. However, computer-assisted technologies are enabling us to produce dental replicas and restorations by shortening the stages, increasing the accuracy, and reducing the wastes. Further work is required to reduce the initial cost and production time before they are routinely used in dental practices and dental laboratories. References Dental Science, 2017. Gypsum Products in Dentistry: Types, Uses, Properties. Available from ,http://www.dental-science.com/gypsum-materials/. (accessed 21.01.18.). ADA, 1972. New American Dental Association specification no. 25 for dental gypsum products. J. Am. Dent. Assoc. 84, 640 644. ADA, 1977. Revised American Dental Association specification no. 19 for non-aqueous, elastomeric dental impression materials. J. Am. Dent. Assoc. 94, 733 741. Anusavice, K.J., 2013. Philip’s Science of Dental Materials. Elsevier/Saunders, St. Louis, MO. Asaoka, K., Bae, J.Y., Lee, H.H., 2012. Porosity of dental gypsum-bonded investments in setting and heating process. Dent. Mater. J. 31, 120 124. Cesar, A.F., Tatiana, S.Z., Fabio, A.R., Yoshio, F., Marcia, F., 2015. Linear setting expansion of different gypsum products. RSBO 12, 61 67. Chan, T.K., Darvell, B.W., 2001. Effect of storage conditions on calcium sulphate hemihydrate-containing products. Dent. Mater. 17, 134 141. Chang, Y.-C., Yu, C.-H., Liang, W.-M., Tu, M.-G., Chen, S.-Y., 2016. Comparison of the surface roughness of gypsum models constructed using various impression materials and gypsum products. J. Dent. Sci. 11, 23 28. Cunningham, W.A., Dunham, R.M., Antes, L.L., 1952. Hydration of gypsum plaster. Ind. Eng. Chem. 44, 2402 2408. Docking, A.R., 1965. Gypsum research in Australia—the setting process. Int. Dent. J. 15, 372 376. Earnshaw, R., Smith, D.C., 2009. The tensile and compressive strength of plaster and stone. Aust. Dent. J. 11, 415 422.
Dental gypsum and investments 53 Earnshaw, R., Marks, B., 1964. The measurement of setting time of gypsum products. Aust. Dent. J. 1 (9), 17 26. Garber, D.K., Powers, J.M., Brandau, H.E., 1985. Effect of spatulation on the properties of high-strength dental stones. J. Mich. Dent. Assoc. 67, 133 136. Hollenback, G.M., Smith, D.D., 1967. A further investigation of the physical properties of hard gypsums. J. Calif. Dent. Assoc. 43, 221 227. Luk, H.W.K., Darvell, B.W., 1991. Strength of phosphate-bonded investments at high temperature. Dent. Mater. 7, 99 102. Luk, W.K., Darvell, B.W., 2003. Effect of burnout temperature on strength of gypsumbonded investments. Dent. Mater. 19, 552 557. Lyon, H.W., George, D., Schoonover, I.C., 1955. The mechanism of hygroscopic expansion in dental casting investments. J. Dent. Res. 34, 44 50. Mahamood, S., Abdul Khader, M., Ali, H., 2016. Applications of 3-D printing in orthodontics: a review. Int. J. Sci. Study 3, 267 270. Mahler, D.B., Bruce Ady, A., 1960. An explanation for the hygroscopic setting expansion of dental gypsum products. J. Dent. Res. 39, 578 589. Mccabe, J.F., Walls, A., 2013. Applied Dental Materials. John Wiley & Sons. Michalakis, K.X., Stratos, A., Hirayama, H., Pissiotis, A.L., Touloumi, F., 2009. Delayed setting and hygroscopic linear expansion of three gypsum products used for cast articulation. J. Prosthet. Dent. 102, 313 318. Malaviya, N., Shrestha, A., 2016. Comparative evaluation of surface detail changes and compressive strength of gypsum casts and dies after immersion in hypochlorite solution and microwave irradiation—an in vitro study. Int. J. Contemp. Med. Res. 3, 1547 1551. Noort, R.V., 2002. Introduction to Dental Materials. Mosby, Edinburgh. Noort, R.V., Barbour, M.E., 2013. Introduction to dental materials. Mosby Elsevier, Edinburgh; New York. O’Brien, W.J., 2008. Dental Materials and Their Selection. Quintessence Pub. Co., Hanover Park, IL; London. Padevět, P., Tesarek, P., Plachý, T., 2011. Evolution of mechanical properties of gypsum in time. Int. J. Mech. 5 (1), 1 9. Paquette, J.M., Taniguchi, T., White, S.N., 2000. Dimensional accuracy of an epoxy resin die material using two setting methods. J. Prosthet. Dent. 83, 301 305. Peyton, F.A., Leibold, J.P., Ridgley, G.V., 1952. Surface hardness, compressive strength, and abrasion resistance of indirect die stones. J. Prosthet. Dent. 2, 381 389. Robert, G.C., John, M.P., John, C.W., 2004. Dental Materials. Mosby, St. Louis, MO. Sanad, M.E.E., Combe, E.C., Grant, A.A., 1982. The use of additives to improve the mechanical properties of gypsum products. J. Dent. Res. 61, 808 810. Singh, M., Garg, M., 1997. Retarding action of various chemicals on setting and hardening characteristics of gypsum plaster at different pH. Cem. Concr. Res. 27, 947 950. TESS, 2017. Formlabs Releases New Dental LT Clear Orthodontic 3D Printing Resin, Updates Dental SG Resin. ,www.3ders.org. (accessed 05.05.18.). von Fraunhofer, J.A., Spiers, R.R., 1983. Strength testing of dental stone: a comparison of compressive, tensile, transverse, and shear strength tests. J. Biomed. Mater. Res. 17, 293 299. Williams, G.J., Bates, J.F., Wild, S., 1983. The effect of surface treatment of dental stone with resins. Quintessence Dent. Technol. 7, 41 45. Williams, K.R., Williams, G.J., 1984. The setting of dental plasters: an electron microscopical study. J. Mater. Sci. 19, 929 938.
54 Advanced Dental Biomaterials Winkler, M.M., Monaghan, P., Gilbert, J.L., Lautenschlager, E.P., 1998. Comparison of four techniques for monitoring the setting kinetics of gypsum. J. Prosthet. Dent. 79, 532 536. Worner, H.K., 1944. The effect of temperature on the rate of setting of plaster of Paris. J. Dent. Res. 23, 305 308.
Ceramic materials in dentistry 4 James K.H. Tsoi Dental Materials Science, Discipline of Applied Oral Sciences, Faculty of Dentistry, The University of Hong Kong, Pokfulam, Hong Kong SAR, P.R. China Chapter Outline 4.1 Introduction 55 4.1.1 Glass ceramics 56 4.1.2 Oxide ceramics 61 4.1.3 Polymer-containing ceramics 64 4.2 Ceramic bonding 65 4.2.1 Mechanism 65 4.2.2 Bond strength evaluation 71 4.2.3 Fatigue 73 4.3 Ceramic biological interaction 74 4.3.1 Surface chemistry 74 4.3.2 Physical parameters 76 4.3.3 Sterilization methods 77 4.4 Conclusion 77 References 77 4.1 Introduction Ceramic, being a composite material, is defined as a (IUPAC Goldbook, 2007): Rigid material that consists of an infinite three-dimensional network of sintered crystalline grains comprising metals bonded to carbon, nitrogen or oxygen Generally speaking, the term ceramic is applicable to any class of inorganic, nonmetallic product that consists of two or more phases, which is different to a glass (that is amorphous), and commonly requires a high temperature during the manufacturing process. Ceramics are capable of withstanding high strength and stiffness. However, they are often brittle and are low in tensile strength. These properties limit the application of ceramics. Nevertheless, ceramics can mimic natural teeth well esthetically. Thus, while limitations still need to be overcome, the application of ceramics in dentistry is somewhat multifold. Ceramics in modern dentistry can be generally classified into glass ceramics and oxide ceramics based on chemical composition. Glass ceramics, also termed Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00004-9 Copyright © 2019 Elsevier Ltd. All rights reserved.
56 Advanced Dental Biomaterials silica-based ceramics, are a group of materials that have been widely used in all-ceramic restorations since the 1970s. These can be further classified as porcelain-based (e.g. feldspathic and leucite-reinforced) or non-porcelain-based (e.g. fluormica glass and lithium disilicate) ceramics. Oxide ceramics are defined as a group of ceramics containing no more than 15% silica with little or no glass phase. They can be subclassified into (1) aluminium-oxide/alumina ceramics, including glass-infiltrated alumina ceramics, densely sintered alumina ceramic systems and (2) zirconium-oxide/zirconia ceramics, including glass-infiltrated and densely sintered zirconia ceramic systems. Because of their stable chemical structure, oxide ceramics have significantly improved mechanical properties and can be regarded as “high-strength ceramics.” 4.1.1 Glass ceramics 4.1.1.1 Feldspathic porcelain As aforementioned, the classification of glass ceramics can be based on “porcelain”. Porcelain was originally developed in China 12002000 years ago and then exported to Europe. Chemically speaking, porcelain contains a number common components such as kaolin, feldspar, and quartz. When these components are heated to a sufficiently high temperature, some components are melted and a silica-based glassy matrix embedding a crystalline core is created. This process is called vitrification. Thus, the porosity (i.e. the flaws) decreased and accompanied by general increase in hardness, mechanical strength, and translucency. In terms of dental feldspathic porcelain, various components exist, typically 70%75% potash feldspar (K2O Al2O3 6SiO2), few percentage of kaolin (Al2O3 2SiO2 2H2O) as binder, 15%20% quartz (SiO2) as crystalline phase, and minute amounts of pigments for different shades. These components are melted or react at various temperatures to form different compounds: Al2 O3 2SiO2 2H2 O ! Al2 O3 2SiO2 ðmetakaoliniteÞ12H2 OðgÞ @ 450o C k (4.i) 3Al2 O3 2SiO2 ðmulliteÞ 1SiO2 (4.ii) @ . 1000o C K2 O Al2 O3 6SiO2 ! K2 O Al2 O3 4SiO2 ðleuciteÞ12SiO2 @1150o C (4.iii) It should be noted that these reactions will never complete unless a sufficiently long enough time is given. Thus, the overall entity is a partially melted (softened, because the temperature is less than the melting point in its pure form) mass consisting of multispecies with reactants and products. This is a sintering process that is defined as (Darvell, 2011): Heating a powder below the melting point of any component such as to permit agglomeration and welding of particles by diffusion alone, with or without applied pressure.
Ceramic materials in dentistry 57 Continuation of heating of the mass would let the amorphous SiO2 [as generated in Eqs. (4.ii) and (4.iii)] to become liquid and chemically cross-link with metal ions in the system, for example, Al31 and K1, to form a 3D network. In addition, the Al atom could chemically react with SiO2 on the surface of crystalline quartz. This newly formed aluminosilicate can effectively form a chemical (covalent) bond between crystalline and amorphous phases. During the cooling process, crystallization would not happen on this partially melted mass due to (1) the high viscosity of the liquid phase and (2) crystallization that requires high activation energy. Ultimately, the final structure is a glassy silicate matrix embedded with (1) a core in mixture form that consists largely of quartz and unreacted feldspar and (2) a dispersion of the separated mullite. Despite quartz being in the crystalline phase, which is deemed to contribute to the overall strength, quartz per se is not a high-strength material. Some additives, such as alumina (Al2O3), can be added in order to improve the mechanical performance. However, caution should be taken for such an addition because the coefficients of thermal expansion (CTE) between different particles are different. Residual stress would happen at the particles’ interface if the CTEs are not matched. On the other hand, as natural feldspar is commonly used by manufacturers, other feldspar analogs that contain sodium, rubidium, and calcium are also found in dental feldspathic porcelain. These compounds will affect the transformation temperatures in the whole system, and thus quality control is an important aspect in making the final product. For example, soda feldspar (Na2O Al2O3 6SiO2) is able to decrease the softening temperature but increase the pyroplastic flow that will slump and distort the shape of the material. As you might notice, the sintering process can be closely controlled by time and temperature. These technique-sensitive procedural parameters allow (if proper!) diffusion between particles and phases, which can reduce the porosity, the total surface area, and the total surface energy, thereby producing ideally a pore-free high-fusing product. These temperature and time parameters, as well as the dimension of the porcelain, can be generally described using the Biot number (Bi): B i 5 Lc h kb (4.1) where Lc is the characteristic length (i.e. volume of body divided by the surface area of the body), h the heat transfer coefficient (i.e. heat flux divided by temperature difference between solid and surrounding fluid such as air), and kb the thermal conductivity. Bi is an index stating the ratio of heat transfer resistance inside and on the surface of the material. So, a small Bi means less heat transfer, while a large Bi means good heat transfer. When considering an object under heating with time, there is no way for heat to be evenly distributed at every single point inside and at the surface. Commonly, at the surface, maximum tensile stress would be encountered and at the center there would be compressive stress. The overall magnitude of the stresses increases with increasing Bi. This said, at the extreme case of Bi 5 0, every test point would be stress free; conversely Bi 5 N would be highly stressed. As a result,
58 Advanced Dental Biomaterials if there is a steep temperature gradient as a sudden deposition of energy from the object to the surrounding or vice versa, thermal stress would be developed. If the thermal stress is larger than the strength of material, failure will occur and this is called thermal shock. Another issue is shrinkage. To control the shrinkage, particle size distribution is important so as to pack the particles as tightly as possible. Nowadays, dental technicians can use manufacturer premixed, fused, and quenched porcelain powders, that is, "cold-shocked", to produce low- and medium-fusing dental porcelain frits. Then, there is no chemical reaction during the heating process since the powders simply melt about the glass transition temperature and fuse together. Upon cooling, a solid mass is produced. If cooling is carried out in the mold, a block can be produced and this is used as a blank CAD/CAM (computer-aided design/computer-aided machining) block. Porcelain, due to its lower strength than other ceramics, is commonly used in applications that have lower mechanical strength requirement, for example, veneer, external facing of metal framework, inlay, and onlay. 4.1.1.2 Leucite-reinforced porcelain As the name implies, this is a porcelain that is reinforced by leucite. With the incorporation of a high proportion of leucite, CTE is raised such that the chance of thermal mismatch with, say, metal during the application of PFM (porcelain-fused metal) can be decreased. Leucite-reinforced porcelain also has a CAD/CAM block version, and currently, there is a similar application with feldspathic porcelain. 4.1.1.3 Fluorine-containing glass ceramics The dental representative for this type of ceramics is known as fluormica, which is mainly based on glass with constituents of SiO2 K2O MgO Al2O3 ZrO2, and also some fluorine-containing compounds. So, the final product is tetrasilica mica crystals within the glass that has a casting temperature of 1380 C, together with a flexural strength of 120150 MPa and a Knoop microhardness value of 360 (cf. enamel: 340). Commercial examples include Dicor (Cast version) and Macor (CAD/CAM version). Interestingly, Leung et al. (2015) have reported that an experimental fluorine-containing ceramic (fluorophlogopite, with K2SiF6) can be a more reliable CAD/CAM material (as a result of Weibull modulus on flexural strength) and can also produce less wear than other glass ceramics (feldspathic and lithium disilicate) as well as polymer-infiltrated ceramic blocks (Table 4.1). It should be noted that although a material may contain fluorinated compounds, it does not imply it has a fluoride ion-releasing ability, although fluoride was detected with the immersion of fluorophlogopite into water. 4.1.1.4 Lithium silicate Lithium disilicate is a glass ceramic based on crystalline phase of Li2Si2O5, which makes up 70 vol.%, and also has low vol.% of Li3PO4 crystalline phase. It has an
Table 4.1 Flexural strength, Vickers hardness, and Weibull modulus for fluorophlogopite, feldspathic porcelain, polymer-infiltrated polymer network, and lithium disilicate computer-aided design/computer-aided machining blocks. Characteristic Experimental fluorophlogopite (1) Feldspathic porcelain (2) Polymer-infiltrated ceramic network (3) Lithium disilicate (4) Group Multiple comparison Three-point flexural strength (MPa), mean (SD) Vickers hardness (HV), mean (SD) Weibull modulus Composition from manufacturer 117.61 (7.62) 106.67 (18.50) 145.95 (12.65) 341.88 (40.25) ,0.001 (1), (2) , (3) , (4) 153.74 (23.62) 594.74 (25.22) 372.29 (51.23) 731.63 (30.64) ,0.001 (1) , (3) , (2) , (4) 18.34 K2SiF6, Al2O3, MgO, SiO2, B2O3 6.93 SiO2, Al2O3, Na2O, K2O, CaO, TiO2, coloring oxides 12.87 SiO2, Al2O3, Na2O, K2O, B2O3, CaO, TiO2, coloring oxides, methacrylate polymer 10.01 SiO2, Li2O, K2O, P2O5, ZrO2, ZnO, Al2O3, MgO, coloring oxides Source: Modified from Leung, B.T.W., Tsoi, J.K.H., Matinlinna, J.P., Pow, E.H.N., 2015. Comparison of mechanical properties of three machinable ceramics with an experimental flurophlogopite glass ceramic. J. Prosthet. Dent. 114 (3), 440446.
60 Advanced Dental Biomaterials unusual microstructure that contains many randomly oriented small, interlocking, plate-like, and needle-like crystals. Such a microstructure (Fig. 4.1) is good to arrest cracks and provide substantial increase of flexural strength. It is quite user-friendly in that the processing temperature is 920 C, and flexural strength is achievable to 350450 MPa. Both hot press and CAD/CAM versions are available. However, it is worth noting that these two versions of ceramics are different in crystal size. Thus, they have different mechanical properties (Fig. 4.1). In addition, a study (Fabian Fonzar et al., 2017) revealed different translucencies of the lithium disilicate would show a difference in mechanical strength, particularly in CAD/CAM blocks. The reason is unknown but the formulation might contribute to this difference. Another type of lithium silicate ceramic is zirconia-lithium silicate, which has added, and is claimed to be reinforced by zirconia fine particles (B10 wt.%) in lithium silicate. It could achieve a higher flexural strength than lithium disilicate. Figure 4.1 SEM images of lithium disilicate hot press and CAD/CAM versions. NB: The particle sizes of hot press are larger than CAD/CAM, which might yield a better mechanical performance. CAD, Computer-aided design; CAM, computer-aided machining; SEM, scanning electron microscope.
Ceramic materials in dentistry 61 Limited scientific literature about this material is available. The commercially available CAD/CAM blocks are 1. Vita Surprinity (unsintered) 2. Dentsply Celtra DUO (sintered) 4.1.2 Oxide ceramics 4.1.2.1 Glass-infiltrated aluminium-oxide ceramic The need for improved fracture strength is vital for all-ceramic restorations. So, increasing the alumina (Al2O3) content can be functional as a reinforcement of the glassy matrix, that is, comparable to leucite crystals. Hence, ceramics that are based on aluminium oxide have been produced and are indicated in multiple areas for restoration, such as copings, frameworks, crowns, and fixed partial denture (FPDs). To use such copings and frameworks the alumina should be veneered with feldspathic porcelain so as to serve two purposes: (1) obtaining better physical strength and (2) esthetics. Glass-infiltrated aluminium-oxide ceramic utilizes a dry-sintered B85% Al2O3 core that is infused with B15% molten lanthanum glass. Commonly, this kind of core can achieve a flexural strength of 400500 MPa after glass infiltration. The commercial representative product is In-Ceram Alumina (VITA). 4.1.2.2 Densely sintered aluminium-oxide ceramic Densely sintered high-purity aluminium-oxide ceramic, as the name implies, utilizes 80.0%99.5% Al2O3 to sinter and does not contain any silica. It should be noted that, given the alumina has a hexagonal close-packed structure with Al31 and other metal ions in two-third of octahedral sites, the alumina is elastic and not viscoplastic at temperature ,1000 C (Ionascu, 2008). However, at the alleviation of temperature, alumina (and many other polycrystalline ceramics such as zirconia) will reduce its brittleness and allow viscoplastic deformation like metal. In this case the grain boundaries sliding and diffusion creep are commonly recognized as the mechanism for this high-temperature viscoplasticity. Such a so-called high-temperature superplastic behavior in polycrystalline ceramics is essential in stress-related densification processes, such as “hot isostatic pressing” and “hot pressing.” As such, the creep rate (plastic flow), ε_ , for the superplasticity of ceramics can be generalized in Arrhenius relationship of ε_ ðσ; T Þ 5 A σ2σ n bp 0 G d D0 expQ=KT (4.2) where σ is the stress applied, σ0 is the threshold stress that is dependent of grain boundaries, shear modulus G, stress component n, Burger’s vector b that represents the magnitude and direction of the lattice distortion resulting from a dislocation in a crystal lattice, grain size d, grain size component p, preexponential constants A and
62 Advanced Dental Biomaterials D0 that are dependent on reaction and diffusion, respectively, activation energy Q, temperature T, and Boltzmann constant k. As shown in Eq. (4.2), grain dislocations and conditions are important for the flow rate. So, all kinds of small amounts of impurities, such as from CAD/CAM machining, trimming, and grinding, would contribute a lot in varying the plastic flow between grain layers (say, superficial vs bulk). In addition, Eq. (4.1) can also predict the heat transfer variation between these layers. Therefore, during hightemperature sintering and also in the cooling, the mismatch of grain layers’ flow would induce intrinsic flaws and internal stresses that are responsible for chipping and cracks. Commercial product Procera Allceram (Nobel Biocare) can achieve a flexural strength of B700 MPa and requires a high sintering temperature (1600 C1700 C). Another product Techceram, which claimed to have 80%90% Al2O3, requires a special plasma technique to utilize the high-temperature plastic flow as stated in Eq. (4.2) to produce the ceramic core, and sinters at 1170 C. 4.1.2.3 Zirconia Zirconia, ZrO2, is the oxide of zirconium and belongs to the group of oxide ceramics. It has various desirable mechanical properties: despite its brittleness, the flexural strength is around 1200 MPa and the hardness very high (more than 1200 HV). It has a high fracture toughness and a good fatigue resistance. From the chemical point of view, zirconia does not dissolve in acids, nor in alkalis. Its appearance is dull, whiteopaque, and it is a good chemical and thermal insulator. As a metal oxide, zirconia has a high radiopacity and is well displayed on X-ray pictures. Furthermore, it is tasteless and odorless. Normally, under low pressure the crystallographic structure of pure zirconia 5 (Fig. 4.2) at room temperature is monoclinic (m, space group C2h or P21/c). With increasing temperatures the crystal lattice of zirconia undergoes phase transformation. Above 1170 C the atoms arrange in the tetragonal phase (t, space group D15 4h or P42/nmc) that is a first-order displacive martensitic phase transformation. When Figure 4.2 Crystalline structure of different phases of zirconia.
Ceramic materials in dentistry 63 zirconia is heated to temperatures higher than 2370 C, a second transformation occurs, which turns tetragonal phase into the cubic form (c, space group O5h or Fm3m) (Botelho et al., 2018). The changes in the crystal lattice of the dimension change from m to t and from t to c is approximately 3%5% for each transformation. However, when cooling the heated zirconia to room temperature, the opposite dimensional changes cause high stresses. So, pure zirconia forms cracks while cooling. To prevent this, some chemicals such as magnesia (MgO), lime (calcium oxide, CaO), ceria (cerium trioxide, Ce2O3), and yttria (Y2O3) are added. These chemicals are able to inhibit crack formation and crack propagation by stabilizing the zirconia at t phase. In principle, when zirconia is subjected to an increase of external stress, microcracks would appear and the stress is the highest at the tip of a crack. If the crystalline structure of zirconia is stabilized in tetragonal form, phase transformation from t to m can happen. Due to the new arrangement of the atoms, the material volume around the tip of the crack increases by 3%5%. So, this makes the crack tip less sharp and also increases the crack path length (i.e. toughened). Therefore, effectively speaking, tetragonal phase stabilization can inhibit crack propagations by t to m transformation. Such a toughening mechanism by phase transformation is able to capture cracks under compressive stresses but is less effective under tension. Therefore yttria-stabilized tetragonal zirconia crystal (Y-TZP) has poor resistance to tensile stresses. Currently, the most commonly used zirconia-based polycrystalline ceramic in dentistry is room temperature stabilized at t phase by adding 36 wt.% or 23 mol % of Y2O3, which is called Y-TZP. Y-TZP is the only zirconia that is listed on the ISO standard for surgical applications and therefore can be used as an implant material. However, there are some negative effects of such phase transformation. In particular, different kinds of stress can cause the t to m transformation and one of them is temperature, which is so-called low temperature degeneration (LTD). As such, the mechanism to inhibit crack propagation no longer works and this effectively decreases the mechanical properties of Y-TZP. The particular reasons for LTD are still controversially discussed in the literature. In general, some reasons are sound, for example, the absorption of water into the surface of zirconia, stabilizers, grain size of the zirconia, the residual stresses after sintering, and any other external stresses. Apparently, the real LTD mechanism could be a combination of various factors. It should be also noticed that, under a pressurized environment the monoclinic phase zirconia can be transited to orthorhombic brookite phase at 311 GPa, or to orthorhombic cotunnite phase at 915 GPa. So, caution should be taken during the manufacturing of zirconia CAD/CAM discs or blocks using the press technique. As a matter of fact, the first generation of Y-TZP is a dull and white material, which cannot mimic the translucency of natural teeth, and is not even comparable to porcelain. The secondary generation of dental zirconia added some colorants to make the monolithic zirconia. Although the shade is there, the material is still opaque because the incident light is (1) reflected at the material interface between air
64 Advanced Dental Biomaterials and ceramic, (2) scattered at grain boundaries, and (3) scattered at pores. So, the final in-line transmission (TIT) would become weaker and thus reduce the translucency. To allow zirconia to achieve the translucency of dental porcelain, a Rayleigh approximation could be used ( TIT 5 ð1 2 RÞexp 128π4 r 3 2 18λ4 ) Δnave ð2n1Δnave Þ 2 t 2 (4.3) where Δnave is the average birefringence, 2n is the grain size, t is the thickness, and λ is the wavelength. The loss of reflection at two surfaces is given by R 5 1 2 2n/ (n2 1 1) where n is the refractive index. Using this calculation, a study (Zhang, 2014) revealed that the grain size should be down to 82 nm (for 1.3 mm thickness), 77 nm (for 1.5 mm thickness), and even 70 nm (for 2 mm thickness). Obviously, such small sizes of zirconia grains are not easy (or not economical) to obtain. So, the easiest way to tackle this is to laminate a layer of porcelain (with similar CTE) such as the PFM. However, chipping might occur which can ultimately lead to fracture. Removing bonded zirconia prostheses, such as crowns or bridges, is difficult and causes damage to the ceramic. Recently, another approach that has been put into practice was the addition of small amounts of the cubic phase zirconia (with isotropic refractive index) into YTZP so as to avoid scattering from the grain boundaries. In fact the manufacturers are not simply adding cubic phase zirconia powders into TZP, but they add the preprocessed B1 μm crystal powders with high amounts of (B50%) cubic zirconia and high yttria content (B5 mol%) into TZP. While the third generation of zirconia seems to have high translucency that matches with current dental porcelain, the strength (flexural strength: 650800 MPa) was compromised because the grain boundaries might not be well attached [in particular during high temperature, the grain boundaries slide, as shown in Eq. (4.2)], and also the t to m toughening may not be so effective. 4.1.3 Polymer-containing ceramics As the name implies, this type of ceramic (in CAD/CAM block form) does not contain solely the ceramic. Commonly, polymeric phase is added so as to increase the crack length, that is, it gives the onset of plasticity under the load and demonstrates a Dugdale cracking model in a way that displacement happens around the crack tips. So, the crack resistance is increased and this effectively lowered the scattering of the strength data. Thus, the Weibull modulus is increased (Table 4.1). Currently, in the market, there are some brands of products named “Hybrid Ceramics.” Indeed, first, whether the CAD/CAM block can be claimed ceramic depends on the existence of an infinite three-dimensional network of crystalline grains. Second, whether it can be named “hybrid” material depends on the interfacial intimate interactions at the nanoscale that has happened between the phases. If the polymeric and ceramic components are separated, we should term these blocks
Ceramic materials in dentistry 65 as “composites.” Therefore if the CAD/CAM block is made by compression of resin-coated ceramic powders, then this is a “composite” material, neither hybrid nor ceramic. One more issue is, in the literature both CAD/CAM and CAD-CAM are used. Although they look the same, slash (/) and hyphen (-) conventionally have different meanings in instrumentation. Accordingly, in the process, if the CAD and CAM are different, for example, operated by different equipment, a slash is preferred. If CAD and CAM are hyphenated (-) that means this is a combined or integrative process. However, in dentistry, it seems to be that both CAD/CAM and CAD-CAM are used interchangeably. 4.2 Ceramic bonding Nowadays, various ceramic restorations, for example, veneers, inlays/onlays, crowns, implants, and resin-bonded bridges exist and a long-term success of the restoration relies heavily on the quality of bonding. To achieve this, resin luting cements were introduced to aid all-ceramic restoration retention. Resin cements not only provide stronger and more durable bonding between ceramics and teeth but can also achieve better esthetic outcomes and maintain higher ceramic strength (Tian et al., 2014). 4.2.1 Mechanism It is generally accepted that adhesion between ceramics and resin cements is provided by two major mechanisms: micromechanical attachment and chemical bonding. Micromechanical attachment can be created by etching and/or grit-blasting, while coupling agents can provide a chemical bond. All these mechanisms rely on the surface treatments/conditioning methods in order to provide a suitable surface for adherent and substrates. As various ceramics differ in chemical conformation and microstructure, it is necessary to establish bonding procedures according to the ceramic type. 4.2.1.1 Chemical surface conditioning Hydrofluoric acid etching Hydrofluoric acid (HF) is an aqueous solution of hydrogen fluoride. HF has a long history in the surface pretreatment of silica-based ceramics before bonding. Such an etching process can increase the roughness of the ceramic surface and consequently creates a micromechanical interlock between the ceramic and resin luting cement. HF is a weak acid compare to the same concentration of, say, HCl. In fact, HF can react with SiO2: SiO2 ðsÞ 1 4HF ðaqÞ ! SiF4 ðgÞ 12H2 O (4.iv) 4SiF4 ðgÞ 1 3H2 OðlÞ 1 2HFðaqÞ ! 3H2 SiF6 ðaqÞ 1 H2 SiO3 ðaqÞ (4.v)
66 Advanced Dental Biomaterials So, the etching process induced by HF does not rely on the “acidic” property of HF but rather the fluoride substitution to oxygen due to electronegativity in glass that forms the Si F glass. Thus, the term “HF etching” should be used, as “acid etching” is misleading. Various glass ceramics and composite can be HF etched to create the roughness for adhesion, as illustrated in Fig. 4.3. For feldspathic porcelain, it was found that the longer the etching time, the rougher the surface will become (Table 4.2), but the biaxial flexural strengths are no different (Liu et al., 2015). So, this result Figure 4.3 SEM images of various CAD/CAM materials before and after HF etching. Mark II: feldspathic porcelain; e.max: lithium disilicate; Suprinity: zirconia-reinforced lithium silicate; Enamic: PICN. CAD, Computer-aided design; CAM, computer-aided machining; HF, hydrofluoric acid; PICN, polymer-infiltrated ceramic network; SEM, scanning electron microscope.
Ceramic materials in dentistry 67 Table 4.2 Comparison between different hydrofluoric acid etching time on feldspathic porcelain. Surface treatment Surface roughness (µm) Biaxial flexural strength (MPa) Shear bond strength (MPa) Weibull modulus of shear bond strength Control Etching for 30 s Etching for 1 min Etching for 2 min 0.11 6 0.02a 0.25 6 0.04b 0.50 6 0.07c,d 0.52 6 0.09d 90.0 6 5.5e 89.0 6 4.8e 86.3 6 7.4e 86.8 6 9.6e 11.60 6 1.93f 10.58 6 1.27f,g 10.37 6 2.32f,g 10.26 6 1.46g 7.38 9.99 5.36 8.11 Different superscript letters indicate the significant differences (P , .05). Source: From Liu, D., Tsoi, J.K.H., Pow, E.H.N., Wong, H.M., 2015. Influence of different etching protocols on the reliability of resin bonding to CAD/CAM feldspathic porcelain. Int. J. Adhes. Adhes. 62, 1824. suggested that the HF etching would change the superficial surface only. On the other hand, the longer etching time, despite increasing the surface roughness, would decrease the shear bond strength. Again, the Weibull moduli as shown in Table 4.2 can demonstrate the bond reliability. Thus, there is a need to follow the manufacturers’ recommendation for the HF etching procedure. Nonetheless, different dosage forms (liquid vs gel), HF concentrations, and components in the formulation would affect the overall outcome. In addition, different test methods, due to the variation in setup, would have different sensitivity for the effects. For example, HF etching on lithium disilicate for 60 seconds may show the same tensile bond strength as 20 seconds. However, a bending test setup would show 60 seconds of HF etching to have significantly lower bond strength than 20 seconds (Wong et al., 2017). Therefore, the test setup would be influential on test sensitivity. Early studies have shown that zirconia surface cannot be etched by HF, and zirconia is termed an “acid-resistant ceramic.” However, recent studies have revealed that some drastic (and very dangerous!) conditions, for example, hot HF, can etch the zirconia. Such a condition can create a higher roughness (Fig. 4.4), comparable bond strength, and no t to m transformation as revealed in grit-blasting. HF-etched zirconia is also shown to be bacteriostatic according a study (Han et al., 2017), possibly due to the sharp and spiky micro-/nanotopographical features on the surface being not favorable for bacterial attachment. Primer Primer, by definition is a (Darvell, 2011): Material used to treat a surface in order to prepare it for subsequent coating or bonding, especially through the deposition of an intermediate layer. In general, primers are used to unite two dissimilar materials by providing chemical bonds between them. In dental ceramics the most commonly used are silane and phosphate coupling agents. These two coupling agents can change the surface chemistry of ceramics, so that the resin cements are able to chemically bond to the
68 Advanced Dental Biomaterials Figure 4.4 SEM images of zirconia without treatment, grit-blasted by silica-coated alumina, and HF etched. XRD on bottomright revealed the monoclinic phase was presented in gritblasted zirconia surface, but only tetragonal phase was found in untreated and HF-etched zirconia. NB: HF-etched zirconia created some nanospiky structures that are bacteriostatic. HF, Hydrofluoric acid; SEM, scanning electron microscope; XRD, X-ray diffraction. Source: From Liu, D., Tsoi, J.K.H., Pow, E.H.N., Wong, H.M., 2015. Influence of different etching protocols on the reliability of resin bonding to CAD/CAM feldspathic porcelain. Int. J. Adhes. Adhes. 62, 1824. ceramics. Given the difference in chemical nature, these coupling agents are used for different ceramics. Recently, a study (Lung et al., 2018) has comprehensively reviewed the dental silane adhesion mechanism. The most commonly used dental silane is 3-(methacryloyloxy)propyltrimethoxysilane (MPS). To have a successful silane adhesion on ceramics the hydroxyl group (OH) on the ceramic surface plays an important role (Fig. 4.5). In terms of glass ceramics, hydroxyl groups come from the equilibrium reaction between the chemisorbed water in atmosphere and SiO2. However, the amount of OH is low. So, another aspect of HF etching is about the formation of silica gel from the metasilicic acid (H2SiO3) upon dehydration [as revealed in Eq. (4.iv)]. The thin silica gel is highly active with a high amount of hydroxyl group. For an oxide ceramic such as zirconia, as revealed in Fig. 4.4, it could not be easily etched by HF, but it can be grit-blasted. The types of grits are influential for
Figure 4.5 Schematic diagram to illustrate the mechanism of silane coupling agent on ceramic surface.
70 Advanced Dental Biomaterials the selection of a proper coupling agent. For example, to use silane as a primer for zirconia, the grits should be silica-coated alumina. Such grits would enable a formation of silica layer on the zirconia surface which is reactable with silane. If ordinary alumina is used for grit-blasting, then the primer of choice should be 10-MDP (10methacryloyloxydecyl dihydrogen phosphate), which is a phosphate coupling agent. The reason is simple: 10-MDP can chemically react with ZrO2, ZrOH, and even Zr41 as revealed on the zirconia surface in a study (Nagaoka et al., 2017). Ideally, the primer layer should be as thin as possible because the primer is shown to be the weakest (in strength) in an adhesive joint system (substrateprimersubstrate). So, to control the thickness, various techniques could be used. For example, using a flow of hot air on the primer layer deposited on lithium disilicate was shown to yield a higher bond strength (Tian et al., 2014). However, it should be noted that, achieving the single primer monolayer (i.e. only one layer) in ordinary dental surface is not feasible. We should term the dental primer layer as ultrathin film. 4.2.1.2 Mechanical surface conditioning Grit-blasting There are various reasons for blasting surfaces with grits in dentistry, for example, to cleanse and to increase their surface roughness. Commonly, surface roughness is one of the surface topographical parameters, which is measurable by profiling at a microlevel (e.g. profilometer) or at a nanolevel (e.g. atomic force microscopy). Grit-blasting can effectively increase the surface roughness and (also increase) the surface area. As such, the micromechanical interlocking between two dissimilar materials could be enhanced. Various grits are available for blasting the dental ceramic surfaces, and the most commonly used are alumina (Al2O3) and silicacoated alumina. These grits are available in different sizes, and in dentistry, the most commonly used are “marked” as 25, 50, and 110 μm, which indicate the average particule size within an acceptable range. These particles also come with different shapes. Obviously, spherical grits are effective in transferring the kinetic energy but not effective for abrasive actions (plowing, cutting, and gouging) compared to sharp and angular grits. Also, the angle and the distance between the blasting tip and ceramic substrate would also yield a difference in adhesion (Ho et al., 2015). Thus, care should be taken in operation and selection of grit-blasting parameters. For example, in the case of glass ceramics, 2550 μm alumina powders are commonly used at a pressure of 0.28 MPa. Zirconia might need a higher pressure, say 23.5 bar, and be blasted with 30 μm (chairside) or 110 μm (laboratory) grits that should be coupled with the suitable primers. The ultimate drawback of grit-blasting is that it could also create sharp micro- or nanocracks on the surface, and the heat generated during the blasting might be able to partially melt the tip of cracks. One possible solution is to use HF to further etch the surface and hence smoothen the crack tips. For zirconia the t to m transformation on zirconia happens during the blasting and is not easily detected (Liu et al., 2013). Nevertheless, normally zirconia cracks from stress concentration areas, for
Ceramic materials in dentistry 71 example, cusps, thin margin, and also connectors. Therefore for the adhesion area, grit-blasting appears to have no or little/minimal damaging effects. Laser Laser (light amplification by stimulated emission of radiation) has also been utilized as one of the surface treatment methods in ceramics. Various parameters, for example, mode of action, irradiation (power), and wavelength, can be adjusted in order to optimize the operation. A study (Liu et al., 2013) has successfully used a high-power CO2 laser to make certain patterns on zirconia surface and improve porcelainzirconia bonding without t to m transformation. It should be noticed that chipping from zirconia or porcelain is a crucial issue that is a catastrophic failure and not reversible. Polishing/grinding A recent study (Ho et al., 2018) has revealed that different surface grinding procedures on zirconia would affect the surface roughness, thickness, fracture strength, and biaxial flexural strength in various ways (Table 4.3). Obviously, tungsten carbide should not be used in trimming or grinding zirconia because the treatment will weaken the material. Diamond is preferred, and it does not cause a huge reduction in zirconia thickness. 4.2.2 Bond strength evaluation The oral cavity is a complex environment that is influenced by numerous intrinsic and extrinsic factors, for example, temperature change, saliva, daily food and drink intake, biting force, and other habits. Thus, in the evaluation of ceramic and its adhesion, laboratory testing should take these variables into account so as to provide additional insights about the problems and solutions. Common methods to evaluate resinceramic bond strength include shear, tensile, and four-point flexural bond strength tests. In particular the shear motion of the crosshead applied on the resin at the resinceramic joint cannot give a “pure shear” at the interface between resin and ceramic surfaces. Although it is true that the tensile component could be reduced by placing the crosshead as close as possible to the ceramic substrate, there is no way to get a zero distance between the crosshead and the ceramic substrate. Thus, “brittle shear” between two substrates cannot be obtained, and “ductile” or “brittle-ductile” shears are observed, which might contribute nonadhesive failures (Jin et al., 2016). So, despite the simplicity in preparing the sample, the shear bond strength test has become less popular now. To test the shear component more accurately, mold-enclosed shear bond strength and lever-induced mold-enclosed shear bond strength tests are recommended. Tensile bond strength test is currently a popular test to evaluate resinceramic bond, and it is so far the best to evaluate tensile component (as the shear component is minimal). Various bond configurations can be set, for example, two-substrate (resinceramic) and three-substrate (ceramicresinceramic or ceramicresin dentine) systems. A report (Tian et al., 2014) has shown the three-substrate
Table 4.3 Effects of different surface grinding procedures on zirconia’s thickness, roughness, breaking force, and biaxial flexural strength. Groups Mean thickness reduction (mm) Mean surface roughness (µm) Mean breaking force (N) Mean biaxial flexural strength (MPa) Coarse diamond Fine diamond Tungsten carbide Fine tungsten carbide Coarse diamond 1 fine tungsten carbide Coarse diamond 1 air particle abrasion Air particle abrasion Control 0.14 6 0.03 0.05 6 0.02 0.07 6 0.01 0.13 6 0.04 0.27 6 0.04 0.21 6 0.04 0.09 6 0.02 0.88 6 0.01 (mean thickness) 5.76 6 0.75 2.52 6 1.22 3.52 6 0.53 2.27 6 0.87 4.49 6 0.40 6.05 6 0.79 3.28 6 0.67 0.44 6 0.16 337.39 6 69.17a 301.97 6 117.86 202.83 6 27.02a 145.26 6 30.72a 137.85 6 9.53a 349.52 6 66.76 372.63 6 68.96 453.25 6 56.59 840.57 6 181.37 598.89 6 241.19 412.01 6 43.13a 352.07 6 65.37a 495.15 6 62.28a 1070.41 6 199.59 840.38 6 176.62 788.34 6 92.32 a The mean difference is significant at the 0.05 level (GamesHowell post hoc test) compared to dry control group. Source: From Ho, C.M.B., Ding, H., Chen, X., Tsoi, J.K.H., Botelho, M., 2018. The effects of dry and wet trimming on the strength of dental zirconia. Ceram. Int. 44 (9), 1045110462.
Ceramic materials in dentistry 73 (ceramicresinceramic) system is a more reliable method to evaluate resinceramic bond strength than the two-substrate (resinceramic) system, since the bond strength value is higher and the Weibull modulus slightly higher. To “make” the bond test sensitive the area of tensile test should be small, say B1 mm2. This dimension is easily achievable for lower strength ceramics such as glass ceramics. However, for highstrength ceramics, such as zirconia, cutting the specimen would not be that easy. The four-point flexural bond strength is a new concept, which uses the threesubstrate system (e.g. ceramicresinceramic) and places the specimen on the four-point bending test jig. The results are promising for the evaluation of resinzirconia (Sanli et al., 2015) and resinlithium disilicate (Wong et al., 2017) bonding. In particular, four-point flexural strength is more sensitive in the evaluation of the effects of different surface treatments on bonds. The comparison of Weibull moduli and relative standard deviation has shown this method to be more reliable than tensile bond strength test (Wong et al., 2017), even when the physical size of the specimen is large. So, the sample preparation (especially for zirconia) is easier. 4.2.3 Fatigue A study (Homaei et al., 2016a) noticed that dental ceramics have a fatigue life distribution that can be fitted with Basquin formula: σa 5 AðN ÞB (4.4) where σa is the stress amplitude, N is the cycles of failure, A is the coefficient constant, and B is the exponent constant. In addition, the fatigue limit of the maximum load (σmax) can be calculated: σmax 5 2σa 1 σmin (4.5) where σmin is the minimum load applied. Table 4.4 revealed a summary between static and fatigue properties of zirconia, lithium disilicate, and polymer-infiltrated ceramic network (PICN). It could be seen that the fatigue limit of ceramics was approximately 45%55% of mean flexural strength. It should be noted that the ceramic fracture surfaces due to fatigue are much smoother than under static loading. This is because of the second stage of fatigue failure in which the crack propagates perpendicular to the tensile stress. Despite the fact the ceramic fatigue happens in laboratory tests, in clinical situation, the ceramic is bonded on the teeth. So, to evaluate the ceramic fatigue or fracture, ceramics bonded on extracted teeth could give a good prediction on the materialteeth interaction. In another study (Homaei et al., 2016b), a ball indentation model on cyclic fatigued CAD/CAM ceramic crowns on human premolars revealed that the average fatigue limit for lithium disilicate is B900k cycles (for 1400 N), and B380k cycles
74 Advanced Dental Biomaterials Table 4.4 Static and fatigue properties of zirconia, lithium disilicate, and polymerinfiltrated ceramic network (PICN). vis-v Four-point flexural strength σ 0 (MPa) Fatigue of stress amplitude σa (MPa) Fatigue limit of maximum load σ max (MPa) Zirconia Lithium dilisilicate PICN 886.9 356.7 135.8 247.1 81.2 33.9 500.1 168.4 73.8 Source: From Homaei, E., Farhangdoost, K., Tsoi, J.K.H., Matinlinna, J.P., Pow, E.H.N., 2016a. Static and fatigue mechanical behavior of three dental CAD/CAM ceramics. J. Mech. Behav. Biomed. Mat. 59, 304313; Homaei, E., Farhangdoost, K., Pow, E.H.N., Matinlinna, J.P., Akbari, M., Tsoi, J.K.H., 2016b. Fatigue resistance of monolithic CAD/CAM ceramic crowns on human premolars. Ceram. Int. 42, 1570915717. for PICN (for 870 N). This said, both materials can survive under normal masticatory forces, and interestingly both materials would mostly fracture above the cementoenamel junction which means it is restorable. In addition, a new attempt (Homaei et al., 2018) using finite element method with the fatigue Eq. (4.4) has demonstrated that the development of shear stress concentration in the margin is important in resinceramic bonding failure because the enamel failed in the adjacent area to the interface. This is in agreement with some of the experimental data and also signifies the importance of shear component in bond strength test. 4.3 Ceramic biological interaction Biological substances, such as cells and bacteria, can attach on any surface in any favorable conditions, such as temperature and humidity. The surface characteristics, cell/bacterial types and properties, as well as the serum proteins and oral environments play significant roles in the growth of these biological substances. In particular, in the case of oral bacteria, there are more than 700 species of microbes existing in the oral cavity, and they can form biofilm that is defined as (Samaranayake, 2018) Complex, functional community of one or more species of microbes, encased in an exopolysaccharide matrix and attached to one another or to a solid surface In terms of ceramics, various surface parameters would contribute to the bacterial attachment. These are either (1) chemically or (2) physically related. 4.3.1 Surface chemistry As illustrated in Fig. 4.5, the ceramic surface has a lot of hydroxyl (OH) groups. Under different pH together with different content of salivary ions and molecules, 2 the hydroxyl groups would be protonated (2OH1 2 ), deprotonated (O ), or stay as
Ceramic materials in dentistry 75 they are, all due to chemical equilibrium. These chemical groups can attract various bacteria with van der Waals forces, electrostatic or Coulomb interactions. Most of the time, the ceramic substrate would be firstly coated with proteins and then bacteria. In terms of cells, the OH is attractive to osteoblasts. For example in zirconia, due to the adsorption of water molecules, a partially hydroxylated surface is formed (Fig. 4.6). With the addition of the OH on zirconia, the crystalline structure at the surface is no longer intact. Chemically speaking, the bond lengths between ZrO and ZrOH are different. Thus, the volume (V) at the surface that contains both ZrO and ZrOH is different from its original form. According to thermodynamic principle, that is, G 5 U 1 pV TS (4.6) where G the Gibbs free energy would be increased with increase of V, given that U the internal energy, p the pressure, T the temperature, and S the entropy are unchanged. This said, the partially hydroxylated surface is not that stable and therefore “activated.” It is also interesting to note that such a hydroxylated surface contains “clusters”, such that different sizes of the clusters would yield different energy (Eg) between the highest occupied (HOMO) and the lowest unoccupied molecular orbitals (LUMO). Fig. 4.7 illustrates the HOMO and LUMO for ZrO2 clusters for Zr number of 12. In fact the zirconia surface shows Eg with 4.55.5 eV, which is a semiconductor range and similar to other chemicals that have osseointegration effect such as titania (TiO2) and hafnia (HfO2). Although the biological knowledge on this aspect is Figure 4.6 Schematic diagram to illustrate the partially hydroxylated (in red square) surface on zirconia.
76 Advanced Dental Biomaterials Figure 4.7 HOMOLUMO illustration for ZrO2 clusters (Zr atom number 5 12). Redblue paired dumbbell-shaped orbitals are the orbitals. We can see the orbitals are confined mainly at the oxygen atoms (in red) rather than zirconium atoms (gray). HOMO, Highest occupied molecular orbitals; LUMO, lowest unoccupied molecular orbitals. limited, such a capacitor-like material would be able to show a certain kind of electrical stimulation that provides another pathway for osseointegration or at least induces the levels of proteins via van der Waals force guidance (Zhuang et al., 1997). Therefore, the biological interaction is enhanced. Some silanes, such as 3-MPS, have been proven to diminish the amount and viability of Candida albicans on zirconia. However, some other silanes, such as 3acryloxypropyltrimethoxysilane (ACPS), may not give the same effects (Villard et al., 2014). Interestingly, MPS on zirconia would make the surface hydrophilic and decrease the surface free energy, but not ACPS. The reason is unknown and worth further investigation. 4.3.2 Physical parameters Various factors, such as roughness and texture, belong to the physical parameters that will contribute to the bacterial and cellular adhesion, in particular during the initial stage. In general a rougher surface would attach more biological substances because the surface can easily trap the substances physically. Therefore, the ceramics should be treated in different ways depending on the application. If applied for veneer, crown, and bridges, the inner surface should be rough in order to bond well. However, the esthetic surface should be smooth or even glazed to make sure the surface cannot be easily attached by bacteria. For an implant body (such as screw), obviously the zirconia should be roughened in order to promote the
Ceramic materials in dentistry 77 osteoblastic reaction. Nevertheless, to prevent periimplantitis, the screw neck part should be smoothened. Another approach could be using hot HF to etch the neck part so as to actively shred the bacteria (Han et al., 2017) by the nanospiky surface. 4.3.3 Sterilization methods In implant application, a sterilization procedure is necessary for all implant components. Ultraviolet light (UV), autoclave, and gamma waves are commonly used in dental industry for the sterilization of implants. A laboratory study (Han et al., 2018) has compared the biofilm formation on different sterilization methods on zirconia. The results revealed that a dry heat-autoclaved sample has a significantly lower bacteria content than other groups, even better than UV-C and gamma ray. High-energy UV-C and gamma ray would even change the color of zirconia, possibly due to the electrons in the quantized energy levels of ZrO being excited. This is also an important indication for a laboratory that a proper sterilization method should be adopted for testing the antibacterial property. 4.4 Conclusion This chapter has provided a quick summary on the classification of dental ceramics, adhesion, and biological interaction. In particular, glass ceramics and zirconia play a major role nowadays with some new developments from polymer-infiltrated ceramic networks. In terms of resinceramic adhesion, HF is useful to etch glasscontaining ceramics and grit-blasting can roughen the oxide ceramic surfaces. A successful adhesion also relies on primer (coupling agents). Bacteria and cells can attach onto ceramic surfaces. We can conclude that ceramic materials are very useful in dentistry, and a proper understanding is essential. References Botelho, M.G., Dangay, S., Shih, K., Lam, W.Y.H., 2018. The effect of surface treatments on dental zirconia: an analysis of biaxial flexural strength, surface roughness and phase transformation. J. Dent. (in press). Darvell, B.W., 2011. A Glossary of Terms in Dental Materials Science, eleventh ed. Fabian Fonzar, R., Carrabba, M., Sedda, M., Ferrari, M., Goracci, C., Vichi, A., 2017. Flexural resistance of heat-pressed and CAD-CAM lithium disilicate with different translucencies. Dent. Mater. 33, 6370. Han, A., Tsoi, J.K.H., Matinlinna, J.P., Chen, Z., 2017. Influence of grit-blasting and hydrofluoric acid etching treatment on surface characteristics and biofilm formation on zirconia. Coatings 7 (8), 130. Han, A., Tsoi, J.K.H., Matinlinna, J.P., Zhang, Y., Chen, Z., 2018. Effects of different sterilization methods on surface characteristics and biofilm formation on zirconia in vitro. Dent. Mater. 34 (2), 272281. Ho, B.J., Tsoi, J.K.H., Liu, D., Lung, C.Y.K., Wong, H.M., Matinlinna, J.P., 2015. Effects of sandblasting distance and angles on resin cement bonding to zirconia and titanium. Int. J. Adhes. Adhes. 62, 2531.
78 Advanced Dental Biomaterials Ho, C.M.B., Ding, H., Chen, X., Tsoi, J.K.H., Botelho, M., 2018. The effects of dry and wet trimming on the strength of dental zirconia. Ceram. Int. 44 (9), 1045110462. Homaei, E., Farhangdoost, K., Tsoi, J.K.H., Matinlinna, J.P., Pow, E.H.N., 2016a. Static and fatigue mechanical behavior of three dental CAD/CAM ceramics. J. Mech. Behav. Biomed. Mat. 59, 304313. Homaei, E., Farhangdoost, K., Pow, E.H.N., Matinlinna, J.P., Akbari, M., Tsoi, J.K.H., 2016b. Fatigue resistance of monolithic CAD/CAM ceramic crowns on human premolars. Ceram. Int. 42, 1570915717. Homaei, E., Jin, X.Z., Pow, E.H.N., Matinlinna, J.P., Tsoi, J.K.H., Farhangdoost, K., 2018. Numerical fatigue analysis of premolars restored by CAD/CAM ceramic crowns. Dent. Mater. (in press). Ionascu, C., 2008. High Temperature Mechanical Spectroscopy of Fine-Grained Zirconia and Alumina Containing Nano-Sized Reinforcements (Ph.D. thesis). Lausanne, EPFL. IUPAC Goldbook, 2007. Definitions of Terms Relating to the Structure and Processing of Sols, Gels, Networks, and Inorganic-Organic Hybrid Materials (IUPAC Recommendations 2007), vol. 79, p. 1801. Jin, X.Z., Homaei, E., Matinlinna, J.P., Tsoi, J.K.H., 2016. A new concept and finite-element study on dental bond strength tests. Dent. Mater. 32, e238e250. Leung, B.T.W., Tsoi, J.K.H., Matinlinna, J.P., Pow, E.H.N., 2015. Comparison of mechanical properties of three machinable ceramics with an experimental flurophlogopite glass ceramic. J. Prosthet. Dent. 114 (3), 440446. Liu, D., Matinlinna, J.P., Tsoi, J.K.H., Pow, E.H.N., Miyazaki, T., Shibata, Y., et al., 2013. A new modified laser pretreatment for porcelain zirconia bonding. Dent. Mater. 29, 559565. Liu, D., Tsoi, J.K.H., Pow, E.H.N., Wong, H.M., 2015. Influence of different etching protocols on the reliability of resin bonding to CAD/CAM feldspathic porcelain. Int. J. Adhes. Adhes. 62, 1824. Lung, C.Y.K., Matinlinna, J.P., Tsoi, J.K.H., 2018. Silane adhesion mechanism in dental applications and surface treatments: a review. Dent. Mater. 31 (1), 1328. Nagaoka, N., Yoshihara, K., Feitosa, V.P., Tamada, Y., Irie, M., Yoshida, Y., et al., 2017. Chemical interaction mechanism of 10-MDP with zirconia. Sci. Rep. 7, 45563. Samaranayake, L.P., 2018. Essential Microbiology for Dentistry (5th Edition), Elsevier. Sanli, S., Comlekoglu, M., Comlekoglu, E., Sonugelen, M., Pamir, T., Darvell, B.W., 2015. Influence of surface treatment on theresin-bonding of zirconia. Dent. Mater. 31, 657668. Tian, T., Tsoi, J.K.H., Matinlinna, J.P., Burrow, M.F., 2014. Evaluation of microtensile bond strength on ceramic-resin adhesion using two specimen testing substrates. Int. J. Adhes. Adhes. 54, 165171. Villard, N., Seneviratne, C., Tsoi, J.K.H., Heinonen, M., Matinlinna, J.P., 2014. Candida albicans aspects of novel silane systemcoated titanium and zirconia implant surfaces. Clin. Oral Imp. Res. 26, 332341. Wong, A.C.H., Tian, T., Tsoi, J.K.H., Burrow, M.F., Matinlinna, J.P., 2017. Aspects of adhesion tests on resin-glass ceramic bonding. Dent. Mater. 33, 10451055. Zhang, Y., 2014. Making yttria-stabilized tetragonal zirconia translucent. Dent. Mater. 30, 11951203. Zhuang, H., Wang, W., Seldes, R.M., Tahernia, A.D., Fan, H., Brighton, C.T., 1997. Electrical stimulation induces the level of TGF-b mRNA in osteoblastic cells by a mechanism involving calcium/calmodulin pathway. Biochem. Biophy. Res. Commun. 237, 225229.
Acrylic denture base materials 5 Touraj Nejatian1,2, Sajjad Pezeshki3 and Azeem Ajaz4 1 Eastman Dental Institute, University College of London, London, United Kingdom, 2 Nottingham Dental Clinic, Nottingham, United Kingdom, 3Babol University of Medical Sciences, Babol, Iran, 4Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Kingdom of Saudi Arabia Chapter Outline 5.1 Introduction 79 5.2 Ideal properties of a denture base material 80 5.3 Acrylic denture base materials 80 5.3.1 Development of denture base materials 80 5.3.2 Chemical structure and mechanism of polymerization 5.3.3 Commercial forms and composition 85 82 5.4 Modified and novel denture base materials and manufacturing technologies 86 5.4.1 5.4.2 5.4.3 5.4.4 5.4.5 5.4.6 5.4.7 Rubber-reinforced resins 87 Fiber-reinforced resins 87 Particulate-reinforced resins 89 Hybrid reinforcement 91 Hypoallergenic resins 91 Thermoplastic resins 92 Novel technologies in manufacturing removable denture base 5.5 Denture lining materials 5.5.1 5.5.2 5.5.3 5.5.4 Clinical indication 94 Hard relining 95 Soft relining 96 Tissue conditioners 98 5.6 Acrylic artificial teeth 5.7 Conclusion 100 References 100 Further reading 104 5.1 93 94 99 Introduction A denture base is a part of a denture that rests on soft tissue, replaces lost soft tissue, and retains the artificial teeth, which require the endurance of various types of stresses. Various types of materials have been used in the construction of denture Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00005-0 Copyright © 2019 Elsevier Ltd. All rights reserved.
80 Advanced Dental Biomaterials bases including cellulose products, epoxy resins, polystyrene, polycarbonates, Bakelite, vinyl resin, nylon, and vulcanite. From 1840 to 1940 different kinds of materials were introduced as denture base materials including wood, ivory, and metal, along with porcelain or natural teeth. The appearance and durability of these materials were not desirable, and the use of mechanical retentive devices was the main method of securing dentures. Vulcanite was produced by heating natural rubber and sulfur, and was the first material to be used for mass production of dentures; however, the undesirable color was its disadvantage. Vinyl resins have low fracture resistance, probably because of their inferior fatigue resistance. The introduction of phenol-formaldehyde (Bakelite) and cellulose nitrates at the beginning of the 20th century was another step in the development of denture bases. However, because of the difficulties in processing and unstable color of Bakelite and warpage of cellulose nitrates as well as its bad taste, they were still far from a desirable denture base material. From 1937 vulcanite was replaced by polymethyl methacrylate (PMMA). Acrylic polymers are now used in prosthetic dentistry as artificial teeth, for denture repair, as facings in crown and bridges, impression trays, record bases, temporary crowns, and obturators of cleft palates. 5.2 Ideal properties of a denture base material A denture should function for a reasonable period in a hostile oral environment without having adverse effects on oral and general health of patients. This means a denture base should have certain characteristics which are summarized in Table 5.1 (Anusavice et al., 2012; Alla, 2013; Rickman et al., 2012). 5.3 Acrylic denture base materials 5.3.1 Development of denture base materials From 1840 to 1940 different kinds of materials were employed as denture base materials, but the appearance and durability of these materials were not desirable, and the use of mechanical retentive devices was necessary. Vulcanite: It was the first material which was used for mass production of dentures and produced by heating natural rubber and sulfur. Its disadvantage was the unfavorable color and porosity that led to an unhygienic denture base (Tandon et al., 2010). Celluloid: It is produced by plasticizing cellulose nitrate with camphor after which the pigmentation is accomplished. A denture base is fabricated by pressing the celluloid blank into a dry, heated mold. Rapid discoloration, water absorption, persistent residual camphor taste, and difficulties in repair are the disadvantages of celluloid (Khindria et al., 2009). Bakelite: The introduction of phenol-formaldehyde (Bakelite) at the beginning of the 20th century was another step in the development of denture bases. However,
Acrylic denture base materials 81 Table 5.1 Requirements of a denture base material. Property Biological Chemical Requirements G G G G Mechanical G G G G G G G G Thermal G G G Esthetic G G Other G G G G G G G Should be neutral, nontoxic, nonirritant, and noncarcinogenic Should be insoluble and nonreactive in the oral fluids or any other food ingredient Should not absorb fluids which cause dimensional changes Should be compatible and bond very well with artificial teeth and liners Should have a high modulus of elasticity and rigid enough against masticatory forces Should be resilient enough to absorb masticatory force to protect the underlying soft tissues Should have a high elastic limit and proportional Limit to prevent the permanent deformation against stresses Should have adequate fatigue resistance and impact strength to withstand the cyclic masticatory force and impact forces, respectively Should be dimensionally stable Should have adequate abrasion resistance Specific gravity should be low (especially for maxillary dentures) Should be a good thermal conductor to encourage natural stimuli and avoid accidental burning of oral tissue COTE should match with that of the artificial teeth to avoid concentration of stress on the toothdenture interface Should have a softening temperature above boiling temperature of water Should exhibit sufficient translucency so that it can be made to match the appearance of the oral tissues Should be capable of being tinted or pigmented Should maintain the desirable properties for an extended period after manufacture Should be inexpensive Should be easy to manipulate Should be radiopaque so that they can be detected if some part of the denture is accidentally swallowed Should be easy to repair Should be easy to clean Should have a long shelf life COTE, Coefficient of thermal expansion. because of the difficulties in processing and repair, and unstable color of Bakelite and its bad taste, it was still far from a desirable denture base material (Noort and Barbour, 2013). From 1937 vulcanite was replaced by PMMA. Polyvinyl chloride (PVC): PVC is still being used as a denture lining material and for the construction of athletic mouth guards. Vinyl resins have low fracture resistance, probably because of their inferior fatigue resistance. They harden over time because the plasticizer leaches out during service.
82 Advanced Dental Biomaterials Its polishing is difficult; therefore, it results in poor denture hygiene and acts as an irritant to the oral mucosal tissues (Munksgaard, 2004). PMMA: In prosthetic dentistry, there are multiple applications for acrylic resins such as artificial teeth, for denture repair, as facings in crown and bridges, impression trays, record bases, temporary crowns, and obturators of cleft palates. Completely polymerized PMMA is highly transparent. Heat-cured acrylic resin dentures contain 0.2%0.5% and the self-cured acrylic resin dentures contain 2%5% (Bahrani et al., 2012) methyl methacrylate as a residual monomer. This residual monomer has the potential to evoke irritation, inflammation, and allergic response of the oral mucosa. Further, the residual monomer can produce both stomatitis and angular cheilitis. Toxic and allergic effects of acrylics depend on several variables, including polymermonomer ratio, storage time and water immersion, type of polymerization cycle, and polymerization method used (Jorge et al., 2003). Acrylic resins are soluble in aromatic hydrocarbons, ketones, and esters and much less soluble in water comparatively. Alcohol functions as a plasticizer and reduces the glass transition temperature. Therefore solutions containing alcohol should not be used for cleaning or storing the dentures. The water absorption of the PMMA changes the dimension of the denture when alternately soaked in water and dried which leads to crazing (Rimple et al., 2011). Due to ease of processing, low cost, light weight, excellent esthetic properties, low water sorption and solubility, and ability to be repaired easily, PMMA is a popular denture base material. However, low thermal conductivity, inferior mechanical strength, brittleness, high coefficient of thermal expansion, and relatively low modulus of elasticity are its weakness during the clinical service (Sakaguchi and Powers, 2007). Research in this area is aimed at modifying the composition or reinforcing the PMMA with stronger materials and developing new materials with better properties. The approaches include modifying the composition with copolymers, reinforcing the dentures with various fibers such as aramid fibers, carbon/graphite fibers, polyethylene fibers, and glass fibers, and also reinforcing with various metallic fillers and nanoparticles (NPs), such as silica and titania NPs. Almost all denture materials are radiolucent. An addition of bismuth (10%15%) or uranyl salts provides adequate radiodensity but at the cost of the increased transverse deflection and water sorption. Producing clinically acceptable levels of radiopacity without degrading physical and esthetic properties of the acrylic resin still remains a challenge (Bloodworth and Render, 1992). 5.3.2 Chemical structure and mechanism of polymerization Methyl methacrylate monomers (the formula is shown below), by using the process of free-radical addition polymerization, make PMMA, which is the main component of acrylic resin denture bases. This process includes the sequence of activation, initiation, propagation, and termination. Depending on the activation process, they are classified as heat-cured, cold-cured, light-activated, and microwave-cured acrylic resins (Sakaguchi and Powers, 2007).
Acrylic denture base materials 83 H Me C=C H C=O O Me The methyl methacrylate structure is formed by additional polymerization in which free-radical polymerization reactions occur with molecules containing carbon double bonds, which is a reaction without a by-product accelerated by heat, light, chemicals, and traces of peroxides. The reaction is initiated by a free radical made by one of these methods. The initiation stage is followed by the propagation stage in which other monomers bound to the free radical speedily shift the free electron to the end of the chain. The chain continues growing until free radicals are terminated. The termination stage can take place in several ways. Any material that can react with a free radical can reduce the initiation or increase the rate of termination. Decreasing the initiation retards polymerization and increasing the rate of termination reduces the degree of polymerization or molecular weight of the final polymer. Hydroquinone, eugenol, or large amounts of oxygen will inhibit or retard the polymerization. Therefore a small amount of hydroquinone is used to increase the shelf life of the monomer (Bartoloni et al., 2000). 5.3.2.1 Curing mechanisms of acrylic denture base materials Heat-cured acrylic resins These materials are usually supplied as a powder and liquid. The composition of the powder includes beads or granules of PMMA, benzoyl peroxide as an initiator, pigment/dyes, opacifiers—titanium/zinc oxides, plasticizer (dibutyl phthalate), and synthetic fibers (nylon/acrylic). Components of the liquid are methyl methacrylate monomer, inhibitor (hydroquinone), and cross-linking agent (ethylene glycol dimethacrylate). Good physical properties are achieved by the high degree of polymerization, but the adaptation to the tissues is poor due to polymerization contraction, thermal contraction, and the strain accompanying stress release during deflasking. Although heat-activated acrylic resins have certain flaws, they are still one of the most commonly used materials for denture fabrication (Anusavice et al., 2012). There is another type of heat-cured resin named as rapid heat-polymerized resin which is a hybrid acrylic that is polymerized within 20 minutes in boiling water. The initiator is a combination of both chemical and heat-activated initiators used to speed up the processing (Jerolimov et al., 1989).
84 Advanced Dental Biomaterials Cold-cured acrylic resins The chemistry of this kind of acrylic resin is similar to that of heat-cured resins except that the cured is initiated by a tertiary amine (e.g., dimethyl-p-toluidine or sulfinic acid) instead of heat. The size of the polymer beads is smaller than the heatcured resin to ease the dissolution in the monomer to produce the dough before going to that level of curing reaction, which prevents proper adaptation of the mixture to the mold. Also, because the method of curing is not as efficient as the heat curing technique, the outcome is lower molecular weight polymers which in turn results in a lower strength property, higher residual monomer, lower color stability, and lower glass transition temperature of the resin (with Tg being typically 75 C80 C). Therefore these materials are highly susceptible to creep, and this can contribute significantly to the eventual distortion of the denture when in use (van Noort, 2012). The radical polymerization of the surface layer is inhibited by oxygen due to deactivation of free radicals produced by the initiator. It was proposed that the polymerization process should be carried out in water in order to obtain an adequate degree of polymerization on the surface of autopolymerizing PMMA resin (Vallittu, 1999). Advantage: Greater dimensional accuracy due to reduced polymerization shrinkage. Visible lightcured resins The composition of these resins is more similar to restorative composites than denture base materials because of their matrix, which is urethane dimethacrylate (UDMA), which contains a small amount of colloidal silica to control the flowability. Their filler consists of acrylic beads that become part of an interpenetrating polymer network structure when cured. Camphorquinone, as a photoinitiator, is included in this composition, so it needs special equipment for polymerization (Anusavice et al., 2012). The acrylic is polymerized in a light chamber with blue light at the wavelength of 400500 nm (Sakaguchi and Powers, 2007). According to Takamata and Setcos (1989), mechanical characteristics of these denture base materials are slightly lower than those of the conventional PMMA. Visible lightactivated acrylic resins provide dentures that have an accurate fit and superior strength. They allow for complete polymerization without residual compounds, ease of fabrication, and manipulation ability to bond with other denture base resins and low bacterial adherence (Ogle et al., 1986). However, visible lightactivated acrylic resin is more brittle than PMMA denture resin. Therefore it is commonly used for relines and repairs rather than for laboratory fabrication of the entire denture base. Advantages: Nontoxic and methyl methacrylate-free denture base resins, less porous than chemically activated denture base resins, facilitate fabrication and final adjustment of the mouth, decreased polymerization shrinkage. Disadvantages: In the conventional manner the resins are opaque and cannot be used, and this system is technique sensitive (Tandon et al., 2010). Microwave-cured acrylic resins In 1968 a microwave-activated denture base material was introduced. Metallic flasks cannot be used as microwaves do not pass through them. Specially designed
Acrylic denture base materials 85 plastic flasks and special microwave machines are needed for this procedure. Some studies showed comparable results to heat-cured acrylic resins, and some even showed more accuracy of fit and lower residual monomer (Sakaguchi and Powers, 2007). Advantages: Reduced curing time (3 minutes), and dough-forming time, color stability, less fracture of artificial teeth and resin bases, and superior denture base adaptation, lower residual monomer, most stable. Disadvantages: Less bond strength to the denture teeth. Increased porosity due to heat entrapment in the nonmetallic flasks. The plastic flasks and polycarbon bolts are relatively expensive and tend to break down on exceeding packing pressure (1200 psi) and after processing several dentures (Tandon et al., 2010). Pour-type denture resins This is a type of cold-cured resin, which commonly named as fluid resins. After mixing, the resulting slurry is so flowable that it can be poured into agar hydrocolloid or silicone or modified plaster molds. The small sizes of beads or powder of pour-type resins are the main difference between them and cold-cured resins (Sakaguchi and Powers, 2007). Centrifugal casting and injection mold methods are techniques used to inject the slurry into the mold. In comparison with heat and cold-cured ones, these offered improved adaptation, dimensional stability, reduced cost, and simple procedure but had low strength, higher solubility, and high residual monomer levels (Tandon et al., 2010). 5.3.3 Commercial forms and composition 5.3.3.1 Powderliquid form There are three reasons for the particular formulation of a powderliquid system: 1. By reaching the dough stage, the processing and packing of the resin under pressure is possible. In addition, in cold-cured types the manipulation and adaptation of the dough on the model and trimming off any excess to produce special trays are more convenient. 2. The polymerization shrinkage is reduced in comparison with monomer because the material being used (i.e., the beads or granules) has already been polymerized. Granules dissolve more readily in monomer than beads and go quickly to the doughy stage. 3. Exothermic reaction of polymerization caused by converting CQC to CC is reduced due to the previous substantial polymerization (Noort and Barbour, 2013). Powder 1. Acrylic polymer (PMMA) may be modified with ethyl, butyl, or other alkyl methacrylates to improve its strength. 2. Initiator such as benzoyl peroxide or diisobutylazonitrile to initiate the polymerization of monomer liquid after being added to the powder. 3. Tissue-like shades of pigments such as mercuric sulfide, ferric oxide, or carbon black. These pigments may be combined into the polymer beads by adding them during the polymerization or mechanically mixed with the polymer beads after polymerization. 4. Zinc or titanium dioxides are used as opacifiers.
86 Advanced Dental Biomaterials 5. Dyed synthetic fibers made from nylon or acrylic are usually added to make natural oral mucosa and simulate the shape of blood vessels. 6. Dibutyl phthalate may be added to powder or monomer as a plasticizer. 7. Heavy metal compounds of elements such as barium or radiopaque glass fillers are added to improve the radiopacity. Desirable radiopacity can be achieved by adding 20% by weight of these compounds, which leads to a reduction in the strength of the material and changes in the appearance of the denture. Also, bismuth or uranyl salts at concentrations of 10%15% and zirconyl dimethacrylate at 35% have been used. Recently, a new radiopaque terpolymer has been synthesized containing 20 30 50 -triodobenzyoyl ethyl methacrylates, methyl methacrylates, and 2-hydroxyethyl methacrylates, which may be used as denture-based materials (Sakaguchi and Powers, 2007). Liquid 1. The liquid part of the powderliquid-type acrylic resin is methyl methacrylate. 2. The inhibitor, which is added to prevent early polymerization, is generally hydroquinone in a concentration of 0.003%0.1%. 3. In self-curing resins, tertiary amines, sulfinic acids, or the more stable salts of sulfinic acid are used as an accelerator and may be added to increase the speed of peroxide decomposition and enable the polymerization of the monomer at room temperature. The commonly used amines are N,N-dimethyl-para-toluidine and N,N-dihydroxyethyl-paratoluidine. These types of resins may be called cold-curing or autopolymerizing resins. 4. Plasticizers are sometimes included to make the polymer softer and more resilient and are commonly low-molecular-weight esters, such as dibutyl phthalate. The mechanism of the effect of these molecules is that they interfere with the interaction between polymer molecules not entering the polymerization reaction. Unfortunately, the release of these materials into the oral fluid hardens the denture base. Some higher esters such as butyl or octyl methacrylate have been used to cope with the problem. They polymerize making an internal plasticizer and do not leach out in oral fluids. Thus the denture remains flexible. 5. Some cross-linking materials that have reactive CR 5 CH groups at opposite ends of the molecules can link long polymers together which is desirable. Glycol dimethacrylate is one of these “cross-linking” materials. Cross-linking materials increase resistance to minute surface cracking, termed crazing (Sakaguchi and Powers, 2007). 5.3.3.2 Gel form Some of the denture materials such as vinyl acrylics can be supplied in a gel form; their composition is the same as powderliquid types except for their chemical accelerators, which cannot be used with the initiator in the monomer. They need a light resource as an alternative curing route. 5.4 Modified and novel denture base materials and manufacturing technologies PMMA is the most popular denture base material because of its excellent properties, such as good esthetics, ease of processing and repair, and low cost. However, it is still far from an ideal denture base material because of residual monomers
Acrylic denture base materials 87 which may cause allergies (Danilewicz-Stysiak, 1971), poor mechanical strength, low fatigue strength, brittleness, poor conductor of heat, low hardness, high coefficient of thermal expansion, thermal shrinkage, poor color stability of self-cured resins, porosity, crazing, warpage, and poor adhesion to metal and porcelain teeth (Nandal et al., 2013). Addition of fillers and fibers to PMMA is a commonly used method to improve both its physical and mechanical properties. Many additives have been suggested to enhance the properties of denture base resin, such as fibers, fillers, or nanofillers. 5.4.1 Rubber-reinforced resins Incorporation of rubber, a low-molecular-weight butadiene styrene, either in polymer beads or matrix of PMMA makes “high impact” resins. This group has greater impact strength and fatigue resistance. Rubber is thought to slow down crack propagation through the denture base material, which improves their fracture resistance. Despite the improvement of the impact strength, Young’s modulus drops, making this a flexible denture. This type of the denture base resin is indicated for patients who drop their dentures repeatedly, for example, Parkinsonism, senility. These materials are available as powderliquid with processing similar to heat-cured resins (Rodford, 1990). 5.4.2 Fiber-reinforced resins Fibers are used to improve the flexural and impact strength as well as the fatigue resistance of the resin. However, the mechanical behavior of the material is affected by the orientation of the fibers. It is suggested that parallel arrangement of the fibers gives the optimum reinforcement (Jagger et al., 2002). Reinforcement of the denture base can be done by fiber weave, namely total fiber reinforcement, or by fiber bundles at the weak areas of the denture, named as partial fiber reinforcement (Figs. 5.15.3). Factors that affect the strength of these composites are (1) adhesion between the fibers and the polymer (Vallittu and Lassila, 1992); (2) impregnation of the fibers with the polymer (Jagger and Harrison, 1999); (3) the concentration of the fibers (Vallittu et al., 1994); (4) test condition; (5) composition of resin; and (6) fiber form, fiber fraction, fiber position, and orientation (Kim and Watts, 2004). Carbon fibers (6570 mm length, 5% by weight and treated with silane coupling agent) are inserted during packing. These fibers have high elastic properties and high resistance to fatigue, which permits the denture to behave like a spring under the occlusal stress and can spread forces evenly throughout the denture. Their strengthening effects can be improved by silanation of the fibers (Vallittu and Lassila, 1992). After the 1980s, despite early promising results, carbon fibers lost favor because of their disadvantages, such as potential toxicity, poor esthetics, polishing problems, difficult handling, and precise placement of the fibers inside the resin (Jagger and Harrison, 1999; Vallittu, 1996).
88 Advanced Dental Biomaterials Figure 5.1 Fiber placed in a tension zone to stop crack propagations. Figure 5.2 Fibers placed at the fracture line. Figure 5.3 Total fiber reinforce placed superficially. Aramid fibers are aromatic polyamide fibers, generally called Kevlar fibers, that have superior wettability and do not require any treatment with a coupling agent. They enhanced the fracture resistance of acrylic resin denture base material (Vallittu and Lassila, 1992); however, the hardness of the resin decreased with increasing fiber concentration. Their unsuitable yellow color and exposure on the surface of the denture makes the denture inconvenient for patients and creates polishing problems for the technician (Kanie et al., 2000). Nylon fibers are polyamide fibers primarily based on aliphatic chains and their main advantage is resistance to shock and cyclic stress. However, water absorption
Acrylic denture base materials 89 has an adverse effect on their mechanical properties. John et al (2001) showed that nylon fibers can improve the flexural strength of PMMA. Nylon increased the fracture resistance of PMMA, as it has high resistance to continual stress. Therefore incorporating nylon fiber in PMMA increased its structural elasticity (Soygun et al., 2013). Polyethene and polypropylene fibers have been used because of their better esthetic characteristics than other fibers. However, these fibers require a surface treatment to increase the impact strength of PMMA significantly (Yu et al., 2012). Glass fibers are mainly used because of their excellent esthetic appearance as well as reinforcement. These fibers have been used in different forms such as woven, loose and continuous (John et al., 2001), or bundles. Increasing fracture resistance of PMMA has been shown to depend on the strong adhesion between glass fibers and the acrylic resin matrix. The silanized glass fibers provide better flexural strength and the position and concentration of glass fibers within the polymer greatly affects the strength (Vallittu et al., 1994). Continuous fibers provide superior reinforcement but orienting continuous fibers at weak parts of the denture is difficult, and there is a risk of void formation inside the fiberpolymer matrix system. Glass fibers may be modified by plasma polymerization techniques using hydroxy-ethyl methacrylate (HEMA), ethylenediamine (EDA), and tetraethylene glycol dimethyl ether (TEGDME) (Ladizesky et al., 1992). 5.4.2.1 Position and placement of fibers 1. Placing the fibers in tension zone in a denture during mastication, which is considered as the weakest area. 2. Placing reinforcement fibers at 90 degrees to the fracture line when repairing the dentures. 3. Mesh should be placed superficially and not buried in the denture. Reinforcement of PMMA with metal oxides improved the physical and mechanical properties of the material, as well as the patients’ sensation of food temperature, which resulted in healthier oral mucosa. The properties of resin reinforced by nanofillers depend on the size, shape, type, and concentration of the added particles (Safi, 2014). 5.4.3 Particulate-reinforced resins 5.4.3.1 Alumina A recent study reported that the addition of alumina (Al2O3) NPs to microwavetreated and untreated PMMA powder has a good level of biocompatibility (Abdulkareem and Hatim, 2015). The main disadvantage of aluminum-reinforced PMMA is discoloration of the resin, which limits its use to areas where it is not visible (Jasim and Ismai, 2014). The addition of Al2O3 to PMMA significantly increased thermal conductivity, but the flexural strength of PMMA did not change
90 Advanced Dental Biomaterials significantly (Kul et al., 2016). Also, the addition of silane-treated aluminum particles to PMMA powder significantly increased the compressive, tensile, and flexural strength and the wear resistance of denture base resin (Chaijareenont et al., 2012). Surface roughness and water sorption of aluminum-reinforced PMMA were not significantly changed (Vojdani et al., 2012). 5.4.3.2 Zirconia Incorporating zirconia (ZrO2) NPs in PMMA increased its thermal conductivity, impact strength, and flexural strength (Gad et al., 2016). Also, compressive strength, fatigue strength, as well as fracture toughness and hardness of the polymer were improved (Salih et al., 2015; Nejatian et al., 2006). The adhesion between the resin matrix and filler particles is very important to enhance the composite’s properties. Therefore a silane coupling agent could be useful to improve the bond strength between zirconia NPs and PMMA. The flexural strength and impact strength of acrylic resin increased, but its tensile strength was not improved. The effect of ZrO2 on the water sorption and solubility of PMMA was different (Bian et al., 2007). In addition, zirconia nanotubes showed a better reinforcing effect than zirconia NPs but, in contrast to zirconia NPs, surface treatment would decrease the reinforcing effect of zirconia nanotubes (Yu et al., 2014). 5.4.3.3 Titanium Adding titanium (TiO2) particles could improve fracture toughness and hardness of PMMA as well as thermal conductivity at the cost of flexural strength (Nejatian et al., 2006). In addition, there is a significant increase in impact strength (Ahmed et al., 2016). The improvement of the physical and mechanical properties of PMMA is limited by poor wettability between fillers and the resin matrix. For this reason, a titanium coupling agent was used for improving the properties of titanium-reinforced PMMA (Elshereksi et al., 2014). The incorporation of silanized TiO2 NPs in PMMA improved the impact strength, transverse strength, and surface hardness of the resin but decreased its water sorption and solubility. Moreover, surface roughness increased with the addition of 3 wt.% of silanized TiO2 NPs to the acrylic resin (Alwan and Alameer, 2015). Adding apatite-coated titanium dioxide and fluoridated apatite-coated titanium dioxide inhibited Candida adhesion due to the antifungal effect (Sawada et al., 2014; Shibata et al., 2007). The addition of barium titanate (BaTiO3) as a radiopacifier to PMMA showed a slight decrease in fracture toughness properties (Elshereks et al., 2009). Silver The addition of silver (Ag) in PMMA could reduce microbial adhesion and colonization because silver has an antimicrobial effect. Also, the addition of silver NPs (Ag NPs) to denture base acrylic resin showed antifungal properties (Casemiro et al., 2007; Monteiro et al., 2012). Incorporating silver powder to PMMA significantly increased thermal conductivity, but the flexural strength values of PMMA were not significantly changed (Kul et al., 2016). The physical and mechanical
Acrylic denture base materials 91 properties of PMMA were improved by the addition of silver NPs, including the increase of thermal conductivity and compressive strength. Therefore it is recommended to be used in the palatal area of maxillary acrylic resin dentures (HamediRad et al., 2014; Ghafari et al., 2014). Hydroxyapatite fillers Hydroxyapatite fillers added to PMMA enhanced the mechanical properties of PMMA. They increased the flexural strength as well as the flexural modulus of PMMA (Tham et al., 2010). Silicon dioxide The mechanical and thermal properties of PMMA were improved by the incorporation of silicon dioxide (SiO2) NPs (Safi, 2014). The addition of SiO2 NPs has increased PMMA impact and transverse strength. Surface hardness also increased with a higher SiO2 NP concentration (Alnamel and Mudhaffer, 2014). Silica-based filler Incorporation of mica in PMMA increased the hardness of acrylic resin and improved the mechanical and thermal properties, as well as the dimensional stability of PMMA, but its flexural strength was reduced because of the weak bond between mica and acrylic resin (Mansour et al., 2013). The incorporation of fluoride glass fillers to PMMA decreased its microbial adhesion at the cost of surface roughness (Al-Bakri et al., 2014; Tsutsumi et al., 2016), whereas nanoclay particles improved the thermal conductivity, but negatively affected the flexural strength (Ghaffari et al., 2016). 5.4.4 Hybrid reinforcement Reinforcement of PMMA by more than one type of fiber was first suggested by Vallittu in 1997. A combination of fibers improves mechanical and physical properties of the material, such as surface roughness (Aljafery and Basima, 2015), tensile strength, flexural modulus (Alhareb and Ahmad, 2011), hardness, thermal conductivity, and radiopacity (Salman and Khalaf, 2015), as well as reducing polymerization shrinkage (Safarabadia et al., 2014; Alhareb et al., 2015). 5.4.5 Hypoallergenic resins Hypoallergenic denture base materials such as diurethane dimethacrylate, polyurethane, polyethylene terephthalate, and polybutylene terephthalate exhibit significantly lower residual monomer content than PMMA (Uzun and Hersek, 1999); thus they could act as an alternative denture base material in allergic patients. Enterephthalate-based dentures show low water solubility than PMMA (Pfeiffer and Rosenbauer, 2004). Light-cured indirect composite containing urethane dimethacrylate (UDMA) is another alternative to PMMA for patients hypersensitive to PMMA (Pfeiffer and Rosenbauer, 2004).
92 Advanced Dental Biomaterials 5.4.6 Thermoplastic resins In this procedure, a fully polymerized basic material is softened by heat (without chemical changes) and injected afterward. These materials have opened a new chapter in making dentures (Philip, 1998) Thermoplastic resins have many advantages over the conventional powderliquid systems. They are very resistant to heat, solvents, and chain degradation as well as being comfortable for the patients. They are also esthetically pleasing and have high fatigue, creep, and wear resistance. Nonporous structure of the material makes them more resistant to bacterial growth. Relining and repairing through repressing is possible. This group includes thermoplastic nylon (polyamide), thermoplastic acetal, thermoplastic acrylic, and thermoplastic polycarbonate (Rawls et al., 1992). 5.4.6.1 Thermoplastic nylon The basic material of thermoplastic nylon is polyamide (derived from diamine and dibasic acid monomers). Thermoplastic nylon, which was introduced to dentistry in the 1950s, was later used by Rapid Injection System (currently known as The Flexite Company, United States) in 1962 as the first flexite thermoplastic (a fluoropolymer—a Teflon-type plastic). Thermoplastic nylon is injected at temperatures of 274 C293 C. This material generally replaces the metal and the pink acrylic denture material used to build the framework for removable partial dentures. Valplast and Flexiplast are two polyamide materials used to make partial dentures. In 1992 The Flexite Company developed the first preformed nylon clasps known as ClaspEze, which are available in pink and clear shades. Nylon is not brittle, has a light weight, does not warp, and provides better esthetics. Nylon is nonporous and more durable against discoloration, and it still retains a slight amount of moisture to keep it comfortable against mucosa (Phoenix et al., 2004). However, nylon is not strong enough to be used as partial denture rests or in occlusal force bearing areas and is also not easy to adjust and polish (Lowe, 2004). 5.4.6.2 Thermoplastic acetal Acetal was first introduced as an unbreakable thermoplastic denture resin in 1971. Rapid Injection System developed the first tooth-colored clasps with thermoplastic fluoropolymer (Keenan et al., 2003). In 1986 Dental “D” reintroduced a homopolymer called acetal as tooth-colored clasps with good short-term mechanical properties, better long-term stability, that were hypoallergic (monomer-free), and had desirable esthetics. The artificial teeth and clasps can be injected together in resin or into the existing denture and polished by a dedicated polishing kit. However, overtrimmed dentures cannot be repaired easily as it requires injection, and poor adhesion between the denture base and the artificial teeth may result in leakage of saliva around the teeth, resulting in debonding (Keenan et al., 2003). Also, a major connector is required to be rigid. Therefore an injected resin major connector for temporary partials is no longer recommended due to high flexibility.
Acrylic denture base materials 93 5.4.6.3 Thermoplastic acrylic Heat-cured PMMA has high porosity, high water absorption, volumetric changes, and residual monomer. Thermoplastic acrylic demonstrates poor impact resistance but has adequate tensile and flexural strength and also provides excellent esthetics due to good translucency and vitality. In addition, it is easy to adjust, reline, repair, and polish at the chair-side. However, acrylic wear resistance is less than acetal especially when subjected to occlusal forces and as a consequence it cannot maintain its vertical dimension for as long. Flexite M.P., a thermoplastic acrylic, is a special blend of polymers with the highest impact strength among this group and resistance to fracture against impact forces, for example, when the dentures are dropped on the floor (Negrutiu et al., 2005). 5.4.6.4 Thermoplastic polycarbonate Polycarbonate is a polymer chain of bisphenol-A carbonate. Similar to acetal resin, polycarbonate resin is also very strong, fracture resistant, and flexible. However, due to low wear resistance, it should not be used on occlusal surfaces. Polycarbonate is not suitable for full or partial dentures, but it is ideal for provisional crown and bridges. Having natural translucency and good finishing, polycarbonate provides excellent esthetics and function in the short or mid-term. This material exhibits little water absorption and no unpleasant smell of saliva absorption. Good adherence and coherence of this material with saliva provides a good affinity with mucous membrane. In addition, it has high impact, crack, and abrasive resistance (Negrutiu et al., 2005). 5.4.7 Novel technologies in manufacturing removable denture base Computer-aided design/computer-aided manufacturing (CAD/CAM) techniques are employed for fabricating fixed and removable partial dental prostheses. Generally, two technological approaches were used for fabricating dentures: CAD/CAM and rapid prototyping (RP) (Lima et al., 2014). In both approaches, a clinical impression procedure for obtaining the morphology of denture bases and the muscular and phonetic positions of the denture teeth is necessary. After making gypsum models, acrylic bases are made, and occlusal rims for each model are fabricated with previously determined standard vertical and centric relationships. Subsequently, the recorded information is scanned, and the complete denture bases are virtually designed. The maxillary and mandibular teeth alignment is verified, and the teeth in each arch are fabricated in one piece, or set, by CAM or RP. Of course, CAD/CAM systems have not completely replaced traditional impression taking, because dynamic movements of the soft tissues influence the quality of the direct digital impression. At present, there are five CAD/CAM systems available for production of dentures: AvaDent (Global Dental Science), Baltic Denture System (Merz Dental GmbH), Ceramill Full Denture System (Amann Girrbach AG), DENTCA/Whole
94 Advanced Dental Biomaterials You (DENTCA, Inc.; Whole You, Inc.), and Wieland Digital Denture (Ivoclar Vivadent, Inc.). Advantages of digital fabrication of dentures: G G G G G G G decreased number of appointments, reduced shrinkage and improved strength, decreased manufacturing time, decreased risk of microorganism colonization on the denture surfaces, advances in standardization for clinical research on removable prostheses, easy reproduction of the denture and manufacture of a trial denture using stored digital data, and superior quality control by clinicians and technicians (Min et al., 2014; Kattadiyil et al., 2015). Limitations and disadvantages of digital fabrication of dentures: G G G G manufacturing challenges used by impression-taking and OVD-recording procedures, MMR transfer, and maintenance of lip support, which is all similar to the procedures used in the conventional process, inability to define the mandibular occlusal plane, expensive materials and increased laboratory cost compared with those for conventional methods, and lack of trial denture, which precludes the evaluation of dentures by patients and dentists before final denture fabrication (Sirirungrojying et al., 1998). 5.5 Denture lining materials The anatomy of the denture bearing area remains under constant change due to variables such as the load, physiological conditions, and tissue tolerance. Therefore edentulous patients should be made aware of the residual ridge resorption which is inevitable, and the prosthesis needs to be modified in a timely manner. Ill-fitting dentures due to the bone and soft tissue changes result in significantly abused tissues. Traumatized and abused tissues are conditioned in a time-dependent manner providing a cushioning effect with the provision of denture lining materials. This not only involves the refitting of the surfaces within complete dentures but also aids in retention of the extra and intraoral prosthesis. 5.5.1 Clinical indication The general indications for relining are as follows: G G G G loss of fit surface, retention, and stability due to alveolar bone resorption, loss of vertical dimension of occlusion (VDO), degradation of denture base, denture extension not covering the mucobuccal fold areas, retromolar pad in the mandibular region, and hamular notch area in the maxillary region,
Acrylic denture base materials G G 95 for elderly patients where new dentures are suspected not to work due to old denture habituation, and patients with the psychological unwillingness to new dentures. The earliest lining materials were made from natural rubber until the 1945s when the synthetic resin was first used as a resilient liner. By 1958 silicone rubbers were introduced as liners. Relining materials are selected and used according to the clinical need. These materials can be classified into three types: G G G Hard reline materials: permanent type (Requirements ADA 13 and 17) Soft liners: semipermanent type (Requirements ADA 17) Tissue conditioners: temporary soft relining materials (Requirements ADA 17). 5.5.2 Hard relining A complete denture may be expected a mean life of about 45 years, depending on the rate of resorption of the alveolar bone. When soft tissue changes arise due to bone resorption, the fit surface of the dentures can be improved by employing a hard reline material. This becomes more challenging with mandibular dentures than maxillary dentures. There are principally two techniques of lining the dentures: direct and indirect. The direct technique involves the relining of the denture with a cold-cured acrylic resin at the chair-side, whereas the indirect technique involves the denture being sent to the laboratory for relining with a heat-cured acrylic. Heat-cured acrylics (polymer powder and monomer liquid) used for relining are identical to the ones used for denture construction (Table 5.2). However, the fabrication differs from the normal construction of dentures. The impression is made using the defective or old denture as a special tray where changed oral tissues are recorded and a stone cast is prepared. The stone cast along with the impression and denture is invested in the flask; after setting, the two halves are separated and the impression material removed. Separating media is applied and polymer powder (PMMA) and liquid (methyl methacrylate) are mixed (Table 5.2). A resin dough is achieved which is packed into the old denture space (previously occupied by impression material) and curing is performed. After curing, the modified denture is obtained for which slight finishing is done. Cold-cured resins are available in two types (Table 5.2). The second type is preferred over the first due to MMA being an irritant to the soft tissues, and it also can sensitize the patients. The only disadvantage of using PEMA and BMA is that they reduce the Tg, and this causes the possibility of dimensional instability. The chairside procedure is straightforward and just needs to follow the recommended manufacturer’s guidelines. Care must be taken, as poorly controlled chair-side relining may result in increased VD. Also, these materials have shown to have an unacceptable taste, high exothermic reactions, and diminished color stability over time.
96 Advanced Dental Biomaterials Table 5.2 The heat-cured and cold-cured resins used for relining dentures. Heat-cured resin Powder Initiator Beads or granules of PMMA Benzoyl peroxide Pigments/dyes Opacifiers Plasticizer Synthetic fibers Liquid Inhibitor Cross-link agent Titanium, zinc oxides Dibutyl phthalate Nylon/Acrylic MMA Hydroquinone EGDMA Cold-cured resin Type 1 Type 2 Powder PMMA Benzoyl peroxide Pigments PEMA Benzoyl peroxide Pigments Liquid MMA Di-n-butyl phthalate Amine BMA Amine BMA, Butyl methacrylate; EGDMA, ethylene glycol dimethacrylate; MMA, methyl methacrylate; PEMA, polyethyl methacrylate; PMMA, polymethyl methacrylate. 5.5.3 Soft relining Long-term soft liners are recommended as a therapeutic measure for patients who are unable to cope up with their existing dentures due to either uncomfortable denture bearing area or prolonged contact between the denture base and the soft tissues resulting in chronic pain, discomfort, and soreness. The soft liner on the tissue side of the dentures provides a cushioning effect to absorb the large dynamic masticatory loads. To achieve such resilient properties, highly plasticized methacrylate or elastomeric materials are used, which are either heat or chemically cured. Heat-cured soft liners are more durable in nature than their counterparts. The clinical situation where long-term soft liners may be indicated: G G G thin, sharp alveolar ridges, excessively resorbed ridges, severe bony undercuts present,
Acrylic denture base materials G G 97 congenital or acquired defects of the palate, and sensitivity due to submucosal exposure of the alveolar bone. An ideal long-term soft liner should G G G G G G G G G G G G G be biocompatible, not degrade easily, be dimensionally stable, have wettability with saliva, exhibit low water solubility or water sorption, exhibit viscoelasticity/compliance/softness, have a good tear and abrasion resistance, exhibit good bond with denture base, be stable and not affected by cleansers or aqueous environment, be easy to handle and manipulate, have good esthetics, be color stable, and not allow colonization of microorganisms and fungi to grow. 5.5.3.1 Methacrylate resin liners (heat-activated) These are supplied as a preformed sheet or in a powder/liquid form. Flexibility of these is imparted by incorporating plasticizers that decrease the glass transition temperature of the resin and act as a lubricant between the polymer chains. The large plasticizer molecules push the polymer chains apart, allowing them to slip on one another thus creating more elastic, flexible, and resilient liners. These polymer resins can bond well with acrylic but with time become stiffer as resilience is lost. The water is absorbed by the liner and leaching of plasticizer occurs, which diminishes the cushioning effect. Clinically, resilient liners are immersed in saliva and dentures are usually stored in water when not in use. These liners are processed in a laboratory, and they show good bond strength to methacrylate base with high tear and abrasion resistance. They can be better polished than silicone. The only drawback is a hard rough surface due to biodegradation of plasticizers promoting calculus formation, food accumulation, and fouling due to microorganisms. 5.5.3.2 Methacrylate resin liners (chemically activated) Soft liners are also available in a chemically activated form having similar composition as heat-activated resins. They are used chair-side and polymerization is usually activated by peroxide tertiary amines and completed in several minutes. They are used as temporary liners as they have a tendency to debond from the denture base and foul within a few weeks. The free monomer also contributes to inferior mechanical properties and poor biocompatibility. 5.5.3.3 Silicone-based liners (heat-activated) They are available as a paste, poly(dimethyl siloxane), and a viscous liquid to which silica is added as a filler, and benzoyl peroxide as an initiator. A crosslinking reaction is catalyzed by heat and peroxide initiator to set the liner. It is
98 Advanced Dental Biomaterials recommended to process liner along with fresh methacrylate dough of the new denture. This has an advantage over the addition of an already polymerized base as a better bond is achieved with the denture base. The laboratory procedures for silicone liners are the same for both heat- and chemical-cured ones. The acceptable thickness of liner is chosen and adequate relief is provided inside the fitting surface of the denture base. Packing, compress molding, and processing are carried out according to manufacturer’s guidelines. The final denture is deflasked, finished, and polished for use. 5.5.3.4 Silicone-based liners (chemically activated)—room temperature vulcanized silicones They are supplied as a two-component system: a paste and a liquid. These liners undergo a condensation cross-linking process which is catalyzed by an organic compound. This reaction is similar in nature to condensation of silicone impression materials. Room temperature vulcanized (RTV) silicones undergo a lesser degree of cross-linking compared to heat-activated silicones, which in turn makes them less favorable as long-term liners and decreases their longevity. Heat-activated silicone rubbers pose better properties than plasticized methacrylates and this enhances their clinical superiority to their counterpart. These soft liners have been advocated to have better qualities providing good cushioning effect with superior flexibility and shock absorption than other soft liners. The major drawback of silicone liner is their tendency to debond from the methacrylate resin base. It has been noted that RTV liners tend to swell and peel off the denture base due to high water sorption and volume changes. They lose the viscoelasticity and become brittle, transferring load on to the bond area. Thus concurrent debonding leads to biofilm adhesion, staining, and decrease in denture strength. Bonding agents and cement have been tried to overcome this problem, but this has not proven to be effective. Both methacrylates and silicones soft liners should not be perceived as permanent liners as both exhibit shortcomings and despite vast clinical benefits they, by far, are not fulfilling the requirements of an ideal material. 5.5.4 Tissue conditioners Short-term resilient liners are also known as tissue conditioners due to their massaging ability to the soft tissues, which can become irritated and inflamed over a period of time. It is a soft material used temporarily to adapt to the fitting surface of the denture. This allows time for the mucosal tissues to regain normal shape and any inflammation to subside. Once tissues regain a better shape and inflammation has been resolved, an impression for the new dentures can be taken. Short-term soft liners vary considerably in their composition depending upon their molecular weight, powder granule size, alcohol content, and plasticizers incorporated. They are supplied as a powder of Polyethylene metacrylate (PEMA), and liquid is not methyl methacrylate but a solvent, such as ethyl alcohol and an aromatic ester such as large molecular-sized plasticizers (e.g., butyl phthalate, butyl glycolate, or
Acrylic denture base materials 99 dibutyl phthalates). They are mixed at chair-side and gelation of tissue conditioner takes place. This is a physical process, devoid of any chemical reaction. Upon mixing the large molecule plasticizers decreases the entanglement process; thereby the softness increases and the Tg decreases. Indications for tissue conditioners/short-term soft liners include the following: G G G G G G G G to treat irritated and inflamed denture-supporting tissues, used for diagnostic purposes to stabilize the vertical dimension of occlusion (VDO), used as an adjunct to correct the occlusion of the old prosthesis in relation to VDO, temporary lining for surgical splints, temporary lining for immediate dentures, relining for obturators, during implant healing as tissue conditioners, and as an adjunct material for functional impressions. Tissue-conditioner requirements differ when used as functional impression materials. For a traumatized tissue, a tissue conditioner should be able to exhibit flow and elastic properties, thus conditioning the tissues with its viscoelastic behavior, which has viscous fluid and elastic solid properties. On the other hand, a functional impression material should exhibit good flow but with no or minimal elastic recovery. This property of the material is of prime importance as under functional stresses it should flow readily to encompass the underlying soft tissues, which are altered under stress conditions. The impression materials must exhibit dimensional stability in regard to solubility, water sorption, and weight change. They should be compliant for many weeks and have a good bond with resin bases. Elastic recovery should be minimal for the success of these as impression materials. To obtain optimum results, functional impressions are removed after 57 days and casts are poured immediately. 5.6 Acrylic artificial teeth Acrylic teeth are supplied as anatomical and nonanatomical zero-degree acrylic resin teeth-sets. There are supplied in different sizes and there are various color shades to select from. Acrylic (methacrylate-based) resin (PMMA) denture teeth are either manufactured by compression dough molding technique or injection molding. The heat-cured acrylic is available in powder and liquid forms for the preparation and to improve the properties and decrease crazing large volumes of cross-linking agents (glycol or other dimethacrylates) are used. Some of the resin-based teeth available in the market are the multiplex polymer matrix resin teeth, resin teeth with Interpenetrating polymer network (IPN) of polymers, and double cross-linked resins. Properties of the acrylic teeth: G G They have adequate biocompatibility. They have good esthetics (different color shades).
100 G G G G G G G G G G G G G G Advanced Dental Biomaterials Their color is stable for good period of time. Crazing can be avoided if cross-linked. Low surface hardness (20 KHN). Abrasive resistance is moderate, tend to wear out by opposing teeth. They have excellent resistance to fracture. They are highly resilient and flexible. They do not produce “click” sound like porcelain teeth. Their bonding to acrylic denture is by chemical means. Their thermal expansion is similar to denture bases (80120 ppm/ C). They are easy to grind and polish. The water sorption causes dimensional changes. The fabrication is a simple and easy procedure. They do not dissolve in solvents. Newer generations of resin teeth are much more stable (wear) and do not get stained. 5.7 Conclusion PMMA has been used as the most popular denture base material for many years due to its unique properties. Efforts have been made to modify or employ novel materials to meet specific needs of the patients. Also, modern technologies have been introduced to the dentistry to reduce the cost and improve the quality and facilitate manufacturing of the dentures. The initial outcome of these advanced technologies is promising; however, further development in technology and material is still required to allow them to be used routinely and replace the conventional techniques and materials. References Abdulkareem, M., Hatim, N., 2015. Evaluation the biological effect of adding aluminum oxide, silver nanoparticles into microwave treated PMMA powder. Int. J. Enhanced Res. Sci. Technol. Eng. 4, 172178. Ahmed, M., El-shennawy, M., Althomali, Y., Omar, A., 2016. Effect of titanium dioxide nano particles incorporation on mechanical and physical properties on two different types of acrylic resin denture base. World J. Nano Sci. Eng. 6, 1111119. Al-bakri, I., Harty, D., Al-omari, W., Swain, M., Chrzanowski, W., Ellakwa, A., 2014. Surface characteristics and microbial adherence ability of modified polymethylmethacrylate by fluoridated glass fillers. Aust. Dent. J. 59, 482489. Alhareb, A., Ahmad, Z., 2011. Effect of Al2O3/ZrO2 reinforcement on the mechanical properties of PMMA denture base. J. Reinf. Plast. Compos. 30, 8693. Alhareb, A., Akil, H., Ahmad, Z., 2015. Mechanical properties of PMMA denture base reinforced by nitrile rubber particles with Al2O3/YSZ fillers. Procedia Manuf. 2, 301306. Aljafery, M., Basima, M., 2015. Effect of addition ZrO2-Al2O3 nanoparticles mixture on some properties and denture base adaptation of heat cured acrylic resin denture base material. J. Baghdad Coll. Dent. 27, 1521.
Acrylic denture base materials 101 Alla, R., 2013. Dental Materials Science. Jaypee Brothers Medical Publishers (Pvt) Ltd, New Delhi. Alnamel, H., Mudhaffer, M., 2014. The effect of silicon dioxide nano-fillers reinforcement on some properties of heat-cure polymethylmethacrylate denture base material. J. Baghdad Coll. Dent. 26, 3236. Alwan, S., Alameer, S., 2015. The effect of the addition of silanized nano titania fillers on some physical and mechanical properties of heat cured acrylic denture base materials. J. Baghdad Coll. Dent. 27, 8691. Anusavice, K., Shen, C., Rawls, H.R., 2012. Phillips’ Science of Dental Materials, twelfth ed. Saunders. Bahrani, F., Safari, A., Vojdani, M., Karampoor, G., 2012. Comparison of hardness and surface roughness of two denture bases polymerized by different methods. World J. Dent. 3, 171175. Bartoloni, J.A., Murchison, D.F., Wofford, D.T., Sarkar, N.K., 2000. Degree of conversion in denture base materials for varied polymerization techniques. J. Oral. Rehabil. 27, 488493. Bian, Y., Zhang, X., Zhu, B., Yu, W., Ruan, D., Min, M., 2007. Influence of different amount of silane coupling agent on the flexural strength of PMMA/nanometer ZrO2 composites. Shanghai Kou Qiang Yi Xue 163, 319323. Bloodworth, K.E., Render, P.J., 1992. Dental acrylic resin radiopacity: literature review and survey of practitioners’ attitudes. J. Prosthet. Dent. 67, 121123. Casemiro, L., Gomes Martins, C., Pires-de-Souza Fde, C., Panzeri, H., 2007. Antimicrobial and mechanical properties of acrylic resins with incorporated silver-zinc zeolite—Part I. Gerodontology 25, 187194. Chaijareenont, P., Takahashi, H., Nishiyama, N., Arksornnukit, M., 2012. Effect of different amounts of 3-methacryloxypropyltrimethoxy silane on the flexural properties and wear resistance of alumina reinforced PMMA. Dent. Mater. J. 3, 623628. Danilewicz-stysiak, Z., 1971. Allergy as a cause of denture sore mouth. J. Prosthet. Dent. 25, 1618. Elshereks, I.N., Mohamed, S., Arifin, A., Mohd Ishak, Z., 2009. Effect of filler incorporation on the fracture toughness properties of denture base polymethylmethacrylate. J. Phys. Sci. 20, 112. Elshereksi, N., Ghazali, M., Muchtar, A., Azhari, C., 2014. Perspectives for titanium-derived fillers usage on denture base composite construction: a review article. Adv. Mater. Sci. Eng. 1, 114. Gad, M., Arrejaie, A., Abdel-Halim, M., Rahoma, A., 2016. The reinforcement effect of nano-zirconia on the transverse strength of repaired acrylic denture base. Int. J. Dent. 2016, 7094056. Ghafari, T., Hamedi-Rad, F., Ezzati, B., 2014. Does addition of silver nanoparticles to denture base resin increase its thermal conductivity? J. Dent. Sch. 32, 139144. Ghaffari, T., Barzegar, A., Hamedi Rad, F., Moslehifard, E., 2016. Effect of nanoclay on thermal conductivity and flexural strength of polymethyl methacrylate acrylic resin. J. Dent. (Shiraz) 17, 121127. Hamedi-Rad, F., Ghaffari, T., Rezaii, F., Ramazani, A., 2014. Effect of nanosilver on thermal and mechanical properties of acrylic base complete dentures. J. Dent. (Tehran) 11, 495505. Jagger, D.C., Harrison, A., 1999. The effect of chopped poly(methyl methacrylate) fibers on some properties of acrylic resin denture base material. Int. J. Prosthodont. 12, 542546.
102 Advanced Dental Biomaterials Jagger, D.C., Jagger, R.G., Allen, S.M., Harrison, A., 2002. An investigation into the transverse and impact strength of “high strength” denture base acrylic resins. J. Oral. Rehabil. 29, 263267. Jasim, B., Ismai, L.I., 2014. The effect of silanized alumina nano-fillers addition on some physical and mechanical properties of heat cured polymethylmethacrylate denture base material. J. Baghdad Coll. Dent. 26, 1823. Jerolimov, V., Brooks, S.C., Huggett, R., Bates, J.F., 1989. Rapid curing of acrylic denturebase materials. Dent. Mater. 5, 1822. John, J., Gangadhar, S.A., Shah, I., 2001. Flexural strength of heat-polymerized polymethyl methacrylate denture resin reinforced with glass, aramid, or nylon fibers. J. Prosthet. Dent. 86, 424427. Jorge, J.H., Giampaolo, E.T., Machado, A.L., Vergani, C.E., 2003. Cytotoxicity of denture base acrylic resins: a literature review. J. Prosthet. Dent. 90, 190193. Kanie, T., Fujii, K., Arikawa, H., Inoue, K., 2000. Flexural properties and impact strength of denture base polymer reinforced with woven glass fibers. Dent. Mater. 16, 150158. Kattadiyil, M.T., Jekki, R., Goodacre, C.J., Baba, N.Z., 2015. Comparison of treatment outcomes in digital and conventional complete removable dental prosthesis fabrications in a predoctoral setting. J. Prosthet. Dent. 114, 818825. Keenan, P.L., Radford, D.R., Clark, R.K., 2003. Dimensional change in complete dentures fabricated by injection molding and microwave processing. J. Prosthet. Dent. 89, 3744. Khindria, S., Mittal, S., Sukhija, U., 2009. Evolution of denture base materials. J. Indian Prosthodont. Soc. 9, 6469. Kim, S.H., Watts, D.C., 2004. The effect of reinforcement with woven e-glass fibers on the impact strength of complete dentures fabricated with high-impact acrylic resin. J. Prosthet. Dent. 91, 274280. Kul, E., Aladağ, L., Yeşildal, R., 2016. Evaluation of thermal conductivity and flexural strength properties of poly(methyl methacrylate denture) base material reinforced with different fillers. J. Prosthet. Dent. 116, 803810. Ladizesky, N.H., Ho, C.F., Chow, T.W., 1992. Reinforcement of complete denture bases with continuous high performance polyethylene fibers. J. Prosthet. Dent. 68, 934939. Lima, J.M., Anami, L.C., Araujo, R.M., Pavanelli, C.A., 2014. Removable partial dentures: use of rapid prototyping. J. Prosthodont. 23, 588591. Lowe, L.G., 2004. Flexible denture flanges for patients exhibiting undercut tuberosities and reduced width of the buccal vestibule: a clinical report. J. Prosthet. Dent. 92, 128131. Mansour, M., Wagner, W., Chu, T., 2013. Effect of mica reinforcement on the flexural strength and microhardness of polymethylmethacrylate denture resin. J. Prosthodont. 22, 179183. Min, K., Silberstein, M., Aluru, N.R., 2014. Crosslinking PMMA: molecular dynamics investigation of the shear response. J. Polym. Sci. B Polym. Phys. 52, 444449. Monteiro, D., Gorup, L., Takamiya, A., de Camargo, E., Filho, A., Barbosa, D., 2012. Silver distribution and release from an antimicrobial denture base resin containing silver colloidal nanoparticles. J. Prosthodont. 21, 715. Munksgaard, E.C., 2004. Leaching of plasticizers from temporary denture soft lining materials. Eur. J. Oral Sci. 112, 101104. Nandal, S., Ghalaut, P., Shekhawat, H., Gulati, M.S., 2013. New era in denture base resins. Dent. J. Adv. Stud. 1, 136143. Negrutiu, M., Sinescu, C., Romanu, M., Pop, D., Lakatos, S., 2005. Thermoplastic resins for flexible framework removable partial dentures. TMJ 55, 295299.
Acrylic denture base materials 103 Nejatian, T., Johnson, A., Van Noort, R., 2006. Reinforcement of denture base resin. Adv. Sci. Technol. 49, 109118. Noort, R.V., Barbour, M.E., 2013. Introduction to Dental Materials. Mosby Elsevier, Edinburgh; New York. Ogle, R.E., Sorensen, S.E., Lewis, E.A., 1986. A new visible light-cured resin system applied to removable prosthodontics. J. Prosthet. Dent. 56, 497506. Pfeiffer, P., Rosenbauer, E., 2004. Residual methyl methacrylate monomer, water sorption, and water solubility of hypoallergenic denture base materials. J. Prosthet. Dent. 92, 7278. Philip, B., 1998. The new acrylic resin with improved thermal diffusivity. J. Prosthet. Dent. 79, 278284. Phoenix, R.D., Mansueto, M.A., Ackerman, N.A., Jones, R.E., 2004. Evaluation of mechanical and thermal properties of commonly used denture base resins. J. Prosthodont. 13, 1727. Rawls, H., Marshall, M., Cardenas, H., Bhagat, H., 1992. Cytotoxicity evaluation of a new radiopaque resin additive-triphenyl bismuth. Dent. Mater. 8, 5459. Rickman, L.J., Padipatvuthikul, P., Satterthwaite, J.D., 2012. Contemporary denture base resins: Part 1. Dent. Update 39, 25. Rimple, Gupta, A., Kamra, M., 2011. An evaluation of the effect of water sorption on dimensional stability of the acrylic resin denture bases. Int. J. Contemp. Dent. 2, 4348. Rodford, R.A., 1990. Further development and evaluation of high impact strength denture base materials. J. Dent. 18, 151157. Safarabadia, M., Khansarib, N., Rezaeic, A., 2014. An experimental investigation of HA/ Al2O3 nanoparticles on mechanical properties of restoration materials. Eng. Solid Mech. 2, 173182. Safi, I., 2014. Evaluation the effect of nano-fillers (TiO2, Al2O3, SiO2) addition on glass transition temperature. E-modulus and coefficient of thermal expansion of acrylic denture base material. J. Baghdad Coll. Dent. 26, 3741. Sakaguchi, R.L., Powers, J.M., 2007. Craig’s Restorative Dental Materials. Elsevier, Mosby, a Division of Reed Elsevier, India Pvt Ltd, New Delhi. Salih, S., Oleiwi, J., Hamad, Q., 2015. Investigation of fatigue and compression strength for the PMMA reinforced by different system for denture applications. Int. J. Biomed. Mater. Res. 3, 513. Salman, T., Khalaf, H., 2015. The influence of adding of modified ZrO2-TiO2 nanoparticles on certain physical and mechanical properties of heat polymerized acrylic resin. J. Baghdad Coll. Dent. 27, 3339. Sawada, T., Sawada, T., Kumasaka, T., Hamada, N., Shibata, T., Nonami, T., et al., 2014. Self-cleaning effects of acrylic resin containing fluoridated apatite-coated titanium dioxide. Gerodontology 31, 6875. Shibata, T., Hamada, N., Kimoto, K., Umemoto, T., Toyoda, M., 2007. Antifungal effect of acrylic resin containing apatite-coated TiO2 photocatalyst. Dent. Mater. J. 26, 437444. Sirirungrojying, S., Srisintorn, S., Akkayanont, P., 1998. Psychometric profiles of temporomandibular disorder patients in southern Thailand. J. Oral. Rehabil. 25, 541544. Soygun, K., Bolayir, G., Boztug, A., 2013. Mechanical and thermal properties of polyamide versus reinforced PMMA denture base materials. J. Adv. Prosthodont. 5, 153160. Takamata, T., Setcos, J.C., 1989. Resin denture bases: review of accuracy and methods of polymerization. Int. J. Prosthodont. 2, 555562. Tandon,, R., Gupta,, S.,, Samarth Kumar, A.,, 2010. Denture base materials: from past to future. Indian J. Dent. Sci. 2, 3339.
104 Advanced Dental Biomaterials Tham, W., Chow, W., Mohd Ishak, Z., 2010. Simulated body fluid and water absorption effects on poly(methyl methacrylate)/hydroxyapatite denture base composites. Expr. Polym. Lett. 4, 517528. Tsutsumi, C., Takakuda, K., Wakabayashi, N., 2016. Reduction of candida biofilm adhesion by incorporation of prereacted glass ionomer filler in denture base resin. J. Dent. 44, 3743. Uzun, G., Hersek, N., 1999. Tinçer effect of five woven fiber reinforcements on the impact & transverse strength of a denture base resin. J. Prosthet. Dent. 81, 616620. Vallittu, P.K., 1996. Comparison of the in vitro fatigue resistance of an acrylic resin removable partial denture reinforced with continuous glass fibers or metal wires. J. Prosthodont. 5, 115121. Vallittu, P.K., 1999. Unpolymerized surface layer of autopolymerizing polymethyl methacrylate resin. J. Oral. Rehabil. 26, 208212. Vallittu, P.K., Lassila, V.P., 1992. Reinforcement of acrylic resin denture base material with metal or fibre strengtheners. J. Oral. Rehabil. 19, 225230. Vallittu, P.K., Lassila, V.P., Lappalainen, R., 1994. Acrylic resin-fiber composite—Part I: The effect of fiber concentration on fracture resistance. J. Prosthet. Dent. 71, 607612. van Noort, R., 2012. The future of dental devices is digital. Dent. Mater. 28, 312. Vojdani, M., Bagheri, R., Khaledi, A., 2012. Effects of aluminum oxide addition on the flexural strength, surface hardness, and roughness of heat-polymerized acrylic resin. J. Dent. Sci. 7, 238244. Yu, S.H., Lee, Y., Oh, S., Cho, H.W., Oda, Y., Bae, J.M., 2012. Reinforcing effects of different fibers on denture base resin based on the fiber type, concentration, and combination. Dent. Mater. J. 31, 10391046. Yu, W., Wang, X., Tang, Q., Guo, M., Zhao, J., 2014. Reinforcement of denture base PMMA with ZrO2 nanotubes. J. Mech. Behav. Biomed. Mater. 32, 192197. Further reading Glossary of Prosthodontic Terms Committee of the Academy of Prosthodontics. The glossary of prosthodontic terms. J. Prosthet. Dent. 117, e1e105.
6 Dental amalgam Nasira Haque1, Safiyya Yousaf2, Touraj Nejatian3, Mansour Youseffi2, Masoud Mozafari4,5 and Farshid Sefat6,7 1 Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom, 2Medical Engineering Department, Faculty of Engineering and Informatics, University of Bradford, Bradford, United Kingdom, 3 Eastman Dental Institute, University College of London, London, United Kingdom, 4 Bioengineering Research Group, Nanotechnology and Advanced Materials Department, Materials and Energy Research Centre (MERC), Tehran, Iran, 5Department of Tissue Engineering & Regenerative Medicine, Faculty of Advanced Technologies in Medicine, Iran University of Medical Sciences (IUMS), Tehran, Iran, 6Interdisciplinary Research Centre in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom, 7Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom Chapter Outline 6.1 Introduction 106 6.2 Dental filling biomaterials 107 6.2.1 Gold fillings 107 6.2.2 Dental composites 108 6.2.3 Amalgam 109 6.3 History of amalgam 111 6.4 Composition of amalgam 112 6.4.1 Low-copper dental amalgam 112 6.4.2 High-copper dental amalgam 112 6.5 Amalgam bonding 113 6.5.1 Nonbonded amalgam restorations 113 6.5.2 Bonded amalgam restorations 114 6.5.3 Nonbonded versus adhesively bonded amalgam restorations 114 6.6 Material properties of amalgam 114 6.6.1 Compressive and tensile strength 114 6.6.2 Creep 115 6.6.3 Tarnish and corrosion 116 6.7 6.8 6.9 6.10 6.11 Dimensional change 117 Hardness 117 Young’s modulus 118 Failure mode 118 Biocompatibility 118 6.11.1 Toxicology of mercury 118 6.12 Conclusion 120 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00006-2 Copyright © 2019 Elsevier Ltd. All rights reserved.
106 Advanced Dental Biomaterials References 121 Further reading 123 6.1 Introduction Dentistry is the branch of medicine which deals with oral health. Services in this field very much depend on the biomaterials, which have developed remarkably over the years. The biomaterials are either directly prepared and placed in a tooth cavity or indirectly made in a laboratory and cemented in the cavity. Various types of biomaterials such as polymer composites, glass ionomer cement, resin cement, gold, and dental amalgams have been used as directed restorative materials. Dental amalgam as a metallic direct restorative material is the primary focus of this chapter. Dental amalgam is a popular biomaterial which has been used successfully in the dental industry for decades. It is composed of a mixture of metal alloy and liquid mercury. Amalgam is mainly used for dental restorations. The metal alloy portion of this biomaterial has a composition made up of silver, tin, copper, and traces of other metals. The unique composition of amalgam gives rise to many useful properties such as excellent durability and strength. These are some of the many factors behind the long service life of amalgam with an average of 810 years (Delta Dental, 2012a,b). The main application of dental amalgam is the restoration of tooth cavities. After removing infected tooth tissue, the cavity is modified to the required shape to receive the amalgam filling. The constituents of amalgam are then mixed to form an amalgam paste, which is then placed in the cavity and left to complete the chemical reaction and become hard (FDA, 2017). Amalgam stays soft for a short period after it is mixed, enough to condense and shape onto the prepared tooth (Gay et al., 1979; Bates, 2006). This biomaterial is relatively cost-effective and widely used. Despite the benefits of amalgam fillings, the safety of them has been a very controversial subject over the years. This is due to the inclusion of elemental mercury within the composition of amalgam. Analytical chemistry techniques were used to investigate dental amalgam, and it has been established that mercury is released continuously (Gay et al., 1979; Bates, 2006). Mechanical friction from the teeth onto the amalgam and mercury being dissolved in saliva, aid the release of mercury. An epidemiological assessment found that there was little to no evidence linking this biomaterial to any effect mortality rates or chronic diseases (Gay et al., 1979; Bates, 2006). A global treaty known as the Minamata Convention on Mercury has been put into place by UNEP (United Nations Environment Programme) to decrease and where possible eliminate the use of mercury. This is to ensure the adverse effects that mercury has on the environment and the human population are decreased. Under this treaty dental amalgam is undergoing a “phase down” of usage. The reason behind this decision is to reduce the adverse effects of using this element, yet to still provide safe healthcare (FDI World Dental Federation, 2017). This book chapter aims to investigate the application of amalgam as a dental biomaterial as well as the composition, biocompatibility, bonding, and material properties of amalgam.
Dental amalgam 6.2 107 Dental filling biomaterials 6.2.1 Gold fillings Gold has been used for over 4000 years as a dental restorative. During the early use of gold in dentistry, a higher importance was given to the aesthetics it had to offer, rather than its contribution to mastication. Gold fillings can be used to restore the function of a tooth affected by dental caries. It is thought that alloyed gold is an excellent choice for fillings, as it has excellent biocompatibility, is easy to manufacture, and lasts over an extended period (Knosp et al., 2003). Pure gold is usually only used for the direct filling of cavities on the tooth surface. Only small cavities may have gold inserted onto them as gold fillings are unable to withstand forces expelled during mastication. Pure gold is soft. Hence it has nearly 50% elongation and a low-stress enduring capacity. These properties allow the pure gold to be cold worked, which is necessary to place it into the cavity easily. Alloyed gold may also be used as a restorative material, and gold is commonly alloyed with base metals (such as indium and copper) and noble metals (Knosp et al., 2003). Cavities can also be filled with an investment cast gold alloy. This material is cemented and used in both dental onlays and inlays (as shown in Fig. 6.1). Alloys with 65%75% gold are commonly used. In addition, gold can be electroformed and covered with porcelain, after which it may be cemented into the designated cavity. Restorative dentistry involves dental bridges and crowns, and involves applying electroformed gold onto porcelain veneers. Electroforming is advantageous because it prevents gold from changing dimensions when it is fired with porcelain, which then minimizes the number of steps required to achieve the desired product. The electroformed gold produced high hardness which is suitable for this application (Knosp et al., 2003). Gold is a noble metal that has a high market value. Gold fillings are costly and therefore unavailable to patients from all socioeconomic backgrounds. Gold casting requires at least two appointments since the cavity must be prepared without undercuts and then the filling must be inserted, which adds to the expense. In comparison to the longevity of a gold filling, the cost is not very significant. The process of insertion is technique sensitive and requires skill from the dentist. In addition, since gold is not tooth colored, it is, therefore, seldom used for anterior tooth restorations (Donovan et al., 2008). The gold casting is created in a laboratory where a replica of the opposing teeth is present, which allows the filling to be shaped, so it aids mastication. It is easier to achieve a smooth finish on a gold restoration, rather than other restorative materials since it is fabricated in the laboratory and is not affected by the oral environment. The smooth finish is advantageous as it is comfortable for the patient and reduces plaque build-up (Donovan et al., 2008). Gold is a “permanent” filling, in contrast to amalgam and resin composites, this is because when it is properly prepared it will not marginally wear or fracture. The gold filling protects the enamel at the edges of the restorations since it can be placed with precision to support the enamel. The coefficient of expansion of gold is
108 Advanced Dental Biomaterials Figure 6.1 Schematic drawing of restorative prostheses: (A) inlay, (B) onlay, (C) overlay, and (D) pinlay. close to a tooth. This property allows both tooth and restoration to undergo contraction and expansion in a complimentary way; this is essential as the temperature within the oral cavity varies (Donovan et al., 2008). 6.2.2 Dental composites Dental composites are used as fillings to restore cavities in a tooth due to decay. Dental composites are usually made of a polymeric resin matrix and a glass filler. The bis-GMA is commonly the main monomer for the resin matrix. However, TEGDMA may be added to make the resin less viscous and, therefore easier to handle. Photopolymerization is used to harden the resin matrix, and the bis-GMA has an essential role in reducing the volumetric shrinkage caused by this process. The size of the glass filler (e.g., fibrillar silicate) can also affect the mechanical properties of the composite. Nanosized fillers aid properties such as Young’s modulus and flexural strength of the composite. The composite filling is desirable for individuals
Dental amalgam 109 since it has a tooth color which is excellent for aesthetic purposes (Fong and Little, 2010). In a study where the capacity of preventing bacterial microleakage was compared between different dental restoratives, the dental composites were one of the best biomaterials which have been used (Murray et al., 2002). Microleakage causes hypersensitivity within dentin due to the irregular fluid movement within its tubules. Uncontrolled microleakage can also cause discomfort to the individual and infection which may lead to pulp inflammation (Cox, 1994). Ormocers are another form of dental composites and are formed via solution and gelation processes (Sivakumar and Valiathan, 2006). They are composed of polysiloxane and have a matrix which has organicinorganic elements. Glass and ceramic make up the inorganic portion of the matrix. Silane molecules connect the organic and inorganic molecules. The properties of ormocers are highly dependent on the proportion of the components that it is made of. Polysiloxanes are influencing factors for the elasticity of this dental biomaterial. The hardness and polarity of an ormocer are determined by the organic polymers, whereas the inorganic polymers impact chemical stability (Zimmerli et al., 2010). This biomaterial was created to minimize volumetric shrinkage, which is caused by the likeness between the thermal expansion coefficient of other composites and the human tooth (Sivakumar and Valiathan, 2006). 6.2.3 Amalgam Dental amalgam is made into a paste and manipulated into the cavity shape onto a carious tooth. This procedure must be carried out before the paste sets and hardens. The mercury inclusion, within the composition of this biomaterial, has been a controversial matter for many years. However, it is not recommended to remove amalgam fillings that have previously been inserted without good reason. Unnecessary loss of a healthy tooth segment and mercury vapor exposure may occur because of the removal of the amalgam filling. Some individuals may be unable to receive amalgam fillings due to sensitivity or allergy to the constituents that make up dental amalgam (FDA, 2017). Dental amalgam will be focused on in greater detail within later sections, as it is the primary focus of this chapter. An amalgam is made up of mercury alloyed to one or more other metals. Mercury does not alloy with certain metals such as iron, platinum, and tungsten. Amalgamation is the process by which an amalgam is processed and is commonly an exothermic reaction (Helmenstine, 2017). There are several examples of amalgams, such as the silver and gold amalgams. In this case, the mercury amalgamates with each of the metals and is used to separate it from its ore. Amalgam extraction is uncommon currently because of its adverse environmental impact. There is a concern with regards to gold mining releasing toxic mercury vapor. Furthermore, another example is the thallium amalgam which has a lower freezing point in comparison to pure mercury. This property makes it useful to use in low-temperature thermometers. Finally, sodium amalgam can act as a reducing agent (Helmenstine, 2017) which is a species that donates electrons to another species.
110 Advanced Dental Biomaterials Dental amalgam is commonly known as “silver fillings” since it has a gray appearance. The main application of dental amalgam is, within dentistry, as a restoration material. The scientific research that has been reviewed by the FDA displays no link between health issues and amalgam fillings. The FDA has deemed dental amalgam fillings safe for use on both adults and children (aged 6 or above). Furthermore, it has been established from limited evidence that this filling does not have an adverse effect on the fetus during pregnancy. Also the levels of mercury in breast milk, due to the mother having amalgam fillings, are acceptable since it is lower than the value considered safe (FDA, 2017). The study that supports this statement (Lygre et al., 2016) included many volunteers and found no substantial evidence to support prenatal amalgam exposure caused health issues for the child. Pereira (2016) conducted a study which compared dental amalgam to glass ionomers, composites, and resin ionomers. When compared to the other direct placement restorations, it was established that amalgam requires a more significant amount of healthy tooth structure removed during cavity preparation. It was seen that the other methods did not require this since they all included an adhesive bonding mechanism. Though this was the case, amalgam is still widely used in clinical conditions since it has a broader tolerance range than the other methods. In an age where aesthetics are becoming more vital, dental amalgam is inadequate since it has a nontooth color. This limitation usually leads to the use of amalgam fillings on mainly posterior teeth. However, in contrast to other direct restoration methods, the cost of amalgam is significantly lower (Pereira, 2016). This is beneficial in poorer countries where people are likely to choose dental restorations, with costs being a more crucial factor in aesthetics. Failure rates of amalgam are low; therefore they require replacement less often than other restoration methods (Pereira, 2016). The durability of amalgam ranges from “good to excellent,” and this factor adds to the longevity of the dental restoration. Mastication expends many forces onto the restorations, and therefore this is an advantageous property. Amalgam also has a superior lifetime of approximately 10 years, which is a longer life span than the glass ionomer, composite, and resin ionomer. Although amalgam has the highest wear resistance in contrast to the other restoration methods, it is also the only one of a brittle nature, since it fragments around the edges (Pereira, 2016). Direct dental restorations must be placed into the cavity by a dental professional. Amalgam filling placement is not technique sensitive. Therefore extensive experience is not be required by the dentist. Amalgam can be easily manipulated, and it is possible to rectify mistakes carried out by the dentist. In addition, as opposed to composites, the placement time of amalgam is short. These factors combined help to make the dentist more comfortable and confident with amalgam usage (Pereira, 2016). To see if there is a trend of amalgam usage worldwide, Burke (2004) reviewed government guidelines on amalgam. Results displayed that there is a decrease in usage worldwide, but the rate of depreciation cannot be identified due to insufficiently available publications. The use of amalgam has declined more slowly in the United Kingdom, in comparison to the United States and Australia. One of the
Dental amalgam 111 main factors influencing this decline in use is the release of mercury. Limited studies have linked the toxic mercury release to Alzheimer’s disease and multiple sclerosis (Bates, 2006). In extremely rare cases, individuals may have an allergic reaction to amalgam fillings. Metal constituents of the filling, such as the elemental mercury, can be the cause of this reaction. Individuals, who have a family history of this issue, are the most susceptible to this condition. Symptoms of this allergic reaction can include skin rashes (Oral Health Centre, 2016b). Amalgam restorations can also lead to amalgam tattoos. These tattoos can occur during the insertion of the amalgam restorations if any of the constituents are accidentally implanted onto neighboring palatal, buccal, gingival, or lingual mucosa. The accumulation of these constituents may leave a gray/black oral lesion behind. Though they are benign, they may resemble oral lesions which are caused by melanoma or Kaposi’s sarcoma. These causes are more severe, but they can be quickly ruled out by biopsy (Dubach and Caversaccio, 2011). 6.3 History of amalgam The exact origin of dental amalgam is unknown, but in 1826 it was recorded that the Frenchman Monsieur Travaux introduced this material. During this period, dental amalgam consisted of mercury combined with finely ground silver coins. In the United States during 1895, a dental amalgam alloy which had a silvertin composition was founded by Dr. G.V. Black. The silvertin alloy particles were combined with mercury. Once the alloy particles and mercury reacted, the initial plastic behavior allowed it to be shaped onto the tooth and then harden into place (Hollenback, 1969; Pereira, 2016). During the early 19th century, amalgam was also used for dental restorations in Europe. During this period there were other metallic restorations available which included hammered lead and gold. In contrast to amalgam, the other restorations had an absence of toxic mercury and had an intricate insertion technique. These factors made the usage of amalgam questionable within the dental field (Greener, 1979; Nicholson, 2002). The controversy over amalgam usage led to the “amalgam wars” in the United States. Dentists during this period had a divided opinion about the safety of amalgam fillings. Some dentists ceased amalgam usage since the mercury had a toxic nature and it was seen to be unethical to put patients at unnecessary risk. The American Society of Dental Surgeons, a professional body at the time, supported the prevention of amalgam usage. They grouped together dentists who did not use amalgam against the dentists who were still using amalgam. Specifications which included thorough testing of amalgam were released in 1929, by the American Dental Association. This aided in making more desirable amalgams, that not only carried out their function but also put patients at a lower risk (Greener, 1979; Nicholson, 2002).
112 Advanced Dental Biomaterials Many figures had an impact on dental amalgam advancements. Notably in 1896, Dr. G.V. Black established the formulae of a balanced dental amalgam composition. This composition has a likeness to the modern-day dental amalgam. The formulae considered the level of contraction and expansion that the amalgam would go through during its lifetime. The cavity preparation techniques which he established made the insertion of amalgam simpler and still act as guidance to dentists today (Singh, 2015). 6.4 Composition of amalgam The alloy segment of dental amalgam can be formed using techniques such as lathe cutting and gas atomization which produces lathe cut or spherical particles, respectively. Various criteria are used for classification of dental amalgams. Among those, the classification based on copper content is probably the most popular one due to the significant effect that copper has on properties of dental amalgam. The traditional amalgam which has a copper particle concentration lower than 6 wt.% is classified as low-copper amalgam, whereas when the alloy portion of the amalgam contains greater than 810 wt.% copper particles, it is classified as high-copper amalgam (Okabe and Cahn, 1990; Hooghan et al., 1996). 6.4.1 Low-copper dental amalgam Low-copper amalgam was initially used for fillings and has a composition of approximately 24.327.6 wt.% tin, 66.772.5 wt.% silver, 1.25.5 wt.% of copper, and occasionally other metals were included (Okabe and Cahn, 1990; Hooghan et al., 1996). High-copper amalgam has now replaced this compound because it has better suited qualities to function as a dental filling (Dental Science, 2016). Hooghan et al. (1996) used transmission electron microscopy to investigate the microstructure of low-copper amalgam. The analysis displayed that the unreacted γ (Ag3Sn) phase was surrounded by the γ1 (Ag2Hg3) and γ 2 (HgSn7) phases. It was also noticed that the reaction layer between γ and γ1 phase included a mixture of β1 (AgHgSn) and some of the Cu6Sn5 (η) phase. The γ1 phase holds the amalgam together and has a similar strength and corrosion resistance to the γ phase, though it has a brittle nature. The γ2 phase is the weakest structure, and if this phase is large it may lead to corrosion (Dental Science, 2016). The amalgamation reaction of low-copper amalgams can be summarized into the following equation (Dental Science, 2016): Excess γ-Ag3 Sn 1 Hg ! Unreacted γ-Ag3 Sn 1 γ1 -Ag2 Hg3 1 γ2 -Sn7 Hg 6.4.2 High-copper dental amalgam High-copper amalgam alloy consists of 17.030.2 wt.% tin, 39.970.1 wt.% silver, 9.529.9 wt.% copper, and may include 0.012.0 wt.% zinc (Okabe and Cahn, 1990;
Dental amalgam 113 Hooghan et al., 1996). Admixed or single-composition alloys are used to prepare this type of amalgam. An admixed composition consists of both lathe cut and spherical particles, whereas single-composition alloys contain the same shape of particles throughout (Dental Science, 2016). The γ2 phase (Sn7Hg) is found in the majority of silvertin alloys (Sarkar and Eyer, 1987; Yap et al., 2004). These alloys were created to reduce the weak γ2 phase (Sn7Hg) by forming more of the (Cu6Sn5) η phase (Beech, 1982). The highcopper amalgam has superior corrosion and creep resistance properties when compared to its predecessor (Sarkar and Eyer, 1987; Yap et al., 2004). Acciari et al. (2005) have suggested that another beneficial quality of the high-copper amalgam is that it decreases toxicity due to mercury vapor emission. During the amalgamation of high-copper dental amalgam, mercury reacts with a powdered amalgam alloy and forms a metallurgical structure. The amalgam alloy is usually made up of copper, silver, tin, and other metals. The solubility of the elements involved is different. Therefore this gives rise to the dissolving of compounds. The mercury defuses into silvertin particles and dissolves them partially. Silvercopper particles, however, are less reactive with mercury (Craig, 1985; Acciari et al., 2005). The γ1 (Ag2Hg3) and γ2 (Sn7Hg) phases are formulated when silvertin particles are partially dissolved in the mercury. The γ2 phase surrounding the silvertin particles and the γ1 phase surrounding the silvercopper particles react to form the η phase (Cu6Sn5). During this process, some silvertin particles (γ phase) are left unreacted (Craig, 1985; Acciari et al., 2005). In other words, high-copper converts the weak and corrosive γ2 phase to a stronger and less corrosive γ1 phase. A structure with up to six phases is produced, and the following equations display the amalgamation process (Craig, 1985; Acciari et al., 2005): γ-Ag3 Sn 1 AgCu 1 Hg ! γ1 -Ag2 Hg3 1 γ2 -Sn7 Hg 1 γ-Ag3 Sn 1 AgCu γ2 -Sn7 Hg 1 AgCu ! η-Cu6 Sn5 1 γ1 -Ag2 Hg3 Although high-copper amalgam is preferred today, the increased content of copper does not always mean an improved clinical performance. When comparing a modified low-copper amalgam to a high-copper amalgam, it is possible that they can show similar clinical performance (Beech, 1982). 6.5 Amalgam bonding 6.5.1 Nonbonded amalgam restorations Dental amalgam does not naturally bond to the tooth surface. The amalgam must be mechanically retained to the tooth via cavity preparation. Retention methods of amalgam may include mechanical devices to aid adherence. Generally, the process of cavity preparation for nonbonded amalgam is less conservative and can lead to loss of healthy tooth structure, which is the main disadvantage of this method.
114 Advanced Dental Biomaterials This leads to the remaining tooth structure becoming weakened; therefore fracture may occur (Dean, 2016). 6.5.2 Bonded amalgam restorations Bonded amalgam restorations are made via an adhesive lining material that is placed under the amalgam. This dental lining has dentin bonding ability and amalgam is placed over it before the adhesive sets. A mechanical bond forms between amalgam and the bonding as the two materials intermix (Dean, 2016). Bonded amalgam restorations reduce the need for cavity preparation to retain the amalgam restoration. This is highly beneficial because the healthy remains of the tooth structure can be salvaged (Bonsor, 2011). Microleakage is the passage of bacteria, molecules, and fluids between the restorative material and tooth (Muliyar et al., 2014). It is beneficial as the restoration will have a low possibility of microleakage and bacterial invasion through the toothmaterial interface. This decreases the risk of sensitivity after restoration and recurrent caries in the long term. Bonding can help with maintaining the integrity of tooth structure which in turn reduces the probability of tooth fracture (Bonsor, 2011). 6.5.3 Nonbonded versus adhesively bonded amalgam restorations A study reviewed articles investigating the bonding of amalgam, and their main findings included an investigation on 31 patients, who in total received 113 amalgam restorations. These restorations were made of either adhesively bonded or nonbonded amalgam. After 2 years there was a follow-up examination, which found no significant difference between the two groups, in marginal change or sensitivity after insertion. Although a small sample size was used and the study was carried out in a university dental clinic, the data may not be representable. This may show that there is limited evidence to show nonbonded amalgam restorations perform worse than adhesively bonded amalgam. The review also suggested dentists may have to consider the additional costs for the adhesively bonded amalgam (Agnihotry et al., 2016). Vanishree et al. (2015) found that bonded amalgam has less probability of microleakage when compared to composite resins and nonbonded amalgam restorations. Also, the study showed that the samples restored by bonded amalgam and composite resins had an inconsistent fracture resistance in comparison to nonbonded amalgam. 6.6 Material properties of amalgam 6.6.1 Compressive and tensile strength An amalgam restoration must be able to withstand biting forces acting during mastication. Therefore it is essential that it has a high compressive strength. The
Dental amalgam 115 compressive strength of the high-copper amalgam is 414 MPa, whereas low-copper amalgam is considerably lower, 380 MPa (Thomas et al., 2013). Jayanthi and Vinod (2013) carried out a study to compare the compressive and flexural strength of dental materials used as core build-up procedures. This study included amalgam (DPI), Fluorocore composite (DENTSPLY, Caulk), Vitremer glass ionomer (3 M), and nanocomposite Filtek (3 M). The flexural strength of amalgam was found to be the lowest of all materials being tested. It had also been established that the compressive strength of amalgam as core build-up material was higher than Vitremer but lower than that of the composites tested. Dental amalgam is unable to withstand high tensile stresses, and therefore to avoid fracture the positioning of it within the cavity must be carefully thought out. This means the tooth cavity should be prepared in a way that the restoration is subjected to the least tensile and shear forces possible. Low-copper amalgam can withstand 60 MPa of a shear force or tension (Manappallil, 2016). During the initial placement of amalgam, since it has a low strength, patients are advised to not bite down hard on it for the first 8 hours. During the setting process when the amalgam begins to solidify, the strength of the amalgam also begins to increase. After these initial 8 hours, the amalgam has nearly reached its complete strength. It is essential for the amalgam to meet the ISO specifications of having a specified compressive strength after 1 hour (100 MPa) and 24 hours (350 MPa). Certain amalgams continue increasing in strength even after 6 months, and this may suggest a continuous reaction between the alloy particles and matrix. The composition is also a factor which affects strength. High-copper single-composition amalgams have the advantage of achieving a high strength (262 MPa) within the first hour of placement (Manappallil, 2016), which will reduce the risk of immature failure of the restoration and patients’ discomfort. Hasheminezhad et al. (2012) investigated whether the compressive strength of amalgam is affected by copper content. They found that high-copper content in amalgam leads to higher compressive strength, due to the elimination of the γ2 phase. There is a positive correlation between strain rate and fracture stress for a copper amalgam. In addition, amalgam alloy particles may change shape from lathe cut into spherical during the processing due to the increase of copper within their composition. 6.6.2 Creep Amalgam suffers from creep when oral forces cause stress and gradually changes its shape. This can occur under constant or intermittent stress (Manappallil, 2016). A high creep rate occurs when there is a gradual failure at the margins of the amalgam filling (Powers and Wataha, 2017). The amalgam begins to flow, which results in parts of it slightly detaching from the cavity and protruding. The protruded edges are unsupported and may fracture under the occlusal forces leaving marginal ditches. The unsupported edges and ditches may also trap food and lead to decay (Thomas et al., 2013). Creep is affected by microstructure in low-copper amalgam. Larger γ1 (Ag2Hg3) phase particles display a lower creep rate, whereas higher creep rates are linked to
116 Advanced Dental Biomaterials the γ2 (Sn7Hg) phase. Low-copper lathe cut amalgam can have up to 6% creep which exceeds the allowable limit. Due to the lack of γ2 phase, single-composition high-copper amalgams, in contrast, have a lower creep rate. Another creep preventative factor in high-copper amalgam is the (η phase) Cu6Sn5 rods stopping the γ1 phase from breaking down. To produce an amalgam with a low creep rate and high strength, it is essential to keep the mercury to alloy ratio low (Manappallil, 2016). The creep values of currently used amalgams are very low (Powers and Wataha, 2017). 6.6.3 Tarnish and corrosion Amalgam restorations that are placed into tooth cavities are susceptible to tarnish and corrosion (Manappallil, 2016). Tarnishing of the amalgam occurs at the surface level; it can be recognized by discoloration from chemical reactions with food. Whereas corrosion can chemically degrade both the surface and the bulk of a tooth, internal corrosion can decrease the strength and affect the shape of the amalgam, which inevitably can lead to failure of the filling (Powers and Wataha, 2017). Lowcopper amalgam restorations are more likely to undergo corrosion since they have a larger γ2 phase. The finding of a study (Amin, 2007) supported low-copper amalgam having a lower corrosion resistance when compared to high-copper amalgam. The study also found that corrosion resistance improved for both types of amalgam when they were left to age in an artificial oral environment. A factor which can increase corrosion levels are patients who are on high sulfur diets, which encourages the formation of black silver sulfide on the surface of the filling. High residual mercury and scratched surface texture of the restoration are the other contributing factors. After the amalgam restoration has been placed into the cavity, polishing the restoration slows down its corrosion. In addition to this, it is crucial to mix the correct proportions of mercury to alloy, since it is possible that excess corrosion may otherwise be caused (Manappallil, 2016). 6.6.3.1 Marginal sealing Upon the initial placement of amalgam, there is a slight gap between the amalgam and cavity walls (Mahler et al., 2009). This gap contains fluid containing molecules, ions, and bacteria (Ben-Amar et al., 1995). Amalgam has a unique ability to seal its marginal gap. This occurs since the margins of amalgam fillings naturally corrode while in service and this leads to corrosive deposits forming. Microleakage is decreased, as the corrosion products build up to fill the gap around the restoration, and seal it. Sealing can reduce chances of bacterial invasion. It is difficult to judge, solely from the appearance of the amalgam’s margins, whether the amalgam filling is sealed and if there is a chance of secondary caries. This is because despite having worn down margins, it is possible to have a well-sealed amalgam filling (Dental Science, 2016). High-copper amalgam has a higher corrosion resistance when compared to lowcopper amalgam. A study investigated whether it would take high-copper amalgam longer to create a seal, due to a slower formation of corrosive deposits. The results
Dental amalgam 117 suggested that it is difficult to anticipate the sealing behavior of amalgam based on corrosion resistance alone. Factors such as the gap size and inclusion of zinc within the amalgams composition, also play a vital role. It was found that it is possible for high-copper amalgam to seal at a similar speed as its low-copper counterpart if the initial gap between the cavity wall and amalgam was small (Mahler et al., 2009). In addition, sealing via corrosive deposits can occur more rapidly if the amalgam is adapted to fit the cavity walls when placed (Ben-Amar et al., 1995). 6.7 Dimensional change Amalgam undergoes dimensional changes during the setting period (Espevik, 1977). Powers and Wataha (2017) defined the overall expansion or contraction of amalgam during its initial setting reaction to be dimensional change. During the initial setting reaction, commonly expansion occurs due to matrix formation, whereas contraction is the result of a reaction between the amalgam alloy particles and mercury. Factors such as the proportion of amalgam alloy to mercury and trituration/ condensation procedures can affect dimensional change. Improper mixing force/ time can adversely affect this factor. Expansion during the setting reaction of amalgam is sought after (Powers and Wataha, 2017). Commonly trituration time of the amalgam alloy can be modified to favor the expansion of amalgam, although it is not advised since other vital properties may become compromised (Espevik, 1977). Once an ideal amalgam is placed into the prepared cavity, it would not expand or contract (Powers and Wataha, 2017). Therefore it is preferred for amalgam to have little dimensional change upon insertion. This is mainly because postoperative sensitivity can be the result of both excessive expansion and contraction. Leakage and decay could also occur if the amalgam underwent excessive contraction (Dental Science, 2016), because this would create a gap between the amalgam and cavity wall. Excessive expansion can also cause the amalgam to move out of the prepared cavity and bulge (Powers and Wataha, 2017). Each type of amalgam behaves differently. However, adherence to the manufacturer’s instructions will positively impact dimensional change (Dental Science, 2016). The requirement for dimensional change, ruled by the American Dental Association (ADA) is 20 μm/cm or less. Most modern amalgams have a dimensional change of zero (Powers and Wataha, 2017). The size of the amalgam restoration can affect the life expectancy. Smaller restorations usually undergo less stress, and therefore their life expectancy increases. Class I amalgam can last between 15 and 18 years, whereas Class II amalgam has a life expectancy of 1215 years. The patient’s oral hygiene and diet largely impact the life expectancy of amalgam restorations (Dental Science, 2016). 6.8 Hardness Optimal hardness in dental amalgam also contributes to the success of amalgam restorations. Although hardness is defined as the resistance of the materials against
118 Advanced Dental Biomaterials scratches, hardness tests actually measure the resistance of a material against indentation from a harder material onto its surface. A study was carried out to see how the addition of zinc oxide (ZnO) and aluminum oxide (Al2O3) nanoparticles affected the hardness of dental amalgam. It was found that when the 40% ZnO filler (the highest proportion) was added to the composition, the hardness increased the greatest, to a value of 0.95 GPa on the Vickers scale. It was observed that different annealing temperatures of ZnO affected the hardness of the dental amalgam. More excellent grain size is created when the ZnO nanoparticles undergo a lower annealing temperature. It was observed that a finer grain size gave better material hardness. In addition, it was observed that adding a more significant proportion of Al2O3 filler produced larger increments in hardness than increasing the ZnO filler (Yahya et al., 2013). 6.9 Young’s modulus Beatty and Pidapartil (1993) established that amalgam has almost 3.5 times greater elastic modulus in tension in comparison to compression. This is greater than the composite resin tested, which had double the elastic modulus in tension in contrast to its compression. This factor was thought to decrease the tensile stress upon the dental amalgam bending. A study carried out by Kumar and Shivrayan (2015) found that high-copper amalgam (named Hi-Aristaloy) possessed the highest value for elastic modulus (17.28 GPa) when compared to other direct core dental biomaterials. In addition, it was found that the γ (Ag3Sn), γ1 (Ag2Hg3), and γ2 (Sn8Hg) phases in amalgam have different values of elastic modulus, where γ was seen to have the highest value (Davies et al., 2010). 6.10 Failure mode Wang and Darvell (2007) conducted a study on the failure mode of dental biomaterials, including amalgam. When the amalgam was subjected to Hertzian indentation, it failed primarily from plastic deformation as the thickness of amalgam became larger or radial cracking originated from its underside. An amalgam of a range of different thicknesses failed. Also, it was reported that a correlation was found between failure load and thickness of amalgam, excluding extremities. 6.11 Biocompatibility 6.11.1 Toxicology of mercury The composition of dental amalgam includes approximately 50% mercury. This inclusion is very controversial since mercury has been linked to many diseases
Dental amalgam 119 (Bates, 2006). Mercury is a naturally occurring element in the environment, and it can also be found in the Earth’s crust. Though amalgam fillings are a source of mercury exposure (Rathore et al., 2012), Clarkson et al. (2003) reported that the other main exposure sources are fish consumption and vaccines. Fish consumption is a source of methylmercury, whereas vaccines can cause exposure to ethyl mercury. These are both organic forms of mercury. Mercury exposure mainly occurs during the removal or insertion of the dental amalgam filling. After the filling has hardened, the mercury vapor release decreases to a value that is below the acceptable threshold. An individual may undergo delayed hypersensitivity because of the insertion of the amalgam filling, the chance of this occurring can be reduced if mercury hygiene procedures are performed (Rathore et al., 2012). The mercury vapor produced is mostly inhaled, and it may also be absorbed into the blood. Dentists are occupationally exposed while they are placing the restorations (Clarkson et al., 2003). Ucar and Brantley (2011) conducted a study to review the literature regarding the toxicology of mercury from dental amalgam. They concluded that the termination of dental amalgam use could not be justified from the literature. Mercury is continuously released from the inserted amalgam filling (Bates, 2006). Ucar and Brantley (2011) reported that factors such as age, amalgam composition, and individual mastication style affect the release of mercury into the oral cavity. Once the mercury has been released, it is dissolved in saliva and passes into the individual. This can occur if intraoral air containing mercury vapor is inhaled and if the filling particles that wear away during mastication are ingested. In addition, swallowing saliva, which has environmental mercury alongside dissolved particles, can pass mercury into an individual. Table 6.1 displays significant forms of mercury exposure, the path the mercury takes within the body, and treatment options. Table 6.1 Major forms of mercury exposure, the path the mercury takes within the body, and treatment options. Variable Mercury vapor Inorganic divalent mercury Methylmercury Ethylmercury Route of exposure Inhalation Oral Oral (from fish consumption) Target organ Central nervous system, peripheral nervous system, kidney Kidney Central nervous system Parenteral (through vaccines) Central nervous system, kidney (Continued)
120 Advanced Dental Biomaterials Table 6.1 (Continued) Variable Mercury vapor Inorganic divalent mercury Methylmercury Ethylmercury Local clinical signs Lungs Gastrointestinal tract Bronchial irritation, pneumonitis ( . 1000 μg/m3 of air) Metallic taste, stomatitis, gingivitis, increased salivation ( . 1000 μg/ m3 of air) Skin Metallic taste, stomatitis, gastroenteritis Urticarial, vesication Systemic clinical signs Kidney Peripheral nervous system Central nervous system Appropriate half-life (whole body) (days) Treatment 6.12 Proteinuria ( . 500 μg/m3 Proteinuria, of air) tubular necrosis Acrodynia Peripheral neuropathy ( . 500 μg/m3 of air) Erethism ( . 500 μg/m3 of air) 60 40 Meso-2dimercaptosuccinic acid Meso-2-3dimercaptosuccinic acid Tubular necrosis Acrodynia Paresthesia, ataxia visual and hearing loss Paresthesia, ataxia visual and hearing loss ( . 200 μg/L of air) 70 20 Chelators not effective Chelators not effective Conclusion Amalgam is a popular biomaterial that has been used successfully for over 150 years, despite the controversial matter of its mercury content. Dental amalgam is mainly used as a direct restorative material. Placement of amalgam restorations is less technique sensitive than resin composites. Therefore dentists find it easy to use with a more predictable outcome. Also, its ability to gain marginal seal after corrosion and to achieve reasonably high compressive strength contributed to its good clinical performance. Despite the risk from mercury exposure, it has kept its popularity in most countries due to its low cost.
Dental amalgam 121 References Acciari, H.A., Guastaldi, A.C., Brett, C.M.A., 2005. Corrosion of the component phases presents in high copper dental amalgams. Application of electrochemical impedance spectroscopy and electrochemical noise analysis. Corros. Sci. 47 (3), 635647. Agnihotry, A., Fedorowicz, Z., Nasser, M., 2016. Adhesively bonded versus non-bonded amalgam restorations for dental caries. Cochrane Database Syst. Rev. (3), CD007517. Amin, W.M., 2007. Comparative electrochemical investigation of the effect of aging on corrosion of dental amalgam. Quintessence Int. 38 (7), e417e424. Bates, M.N., 2006. Mercury amalgam dental fillings: an epidemiologic assessment. Int. J. Hyg. Environ. Health 209 (4), 309316. Beatty, M.W., Pidapartil, R.M.V., 1993. Elastic and fracture properties of dental direct filling materials. Biomaterials 14 (13). Available from: https://www.sciencedirect.com/science/ article/pii/0142961293901925 (accessed 9.03.18.). Beech, D.R., 1982. High copper alloys for dental amalgam. Int. Dent. J. 32 (3), 240251. Ben-Amar, A., Cardash, H.S., Judes, H., 1995. The sealing of the tooth/amalgam interface by corrosion products. J. Oral Rehabil. 22 (2), 101104. Bonsor, S.J., 2011. Bonded amalgams and their use in clinical practice. Dent. Update 38 (4), 222224. 226-8, 230. Burke, F.J.T., 2004. Amalgam to tooth-coloured materials—implications for clinical practice and dental education: governmental restrictions and amalgam-usage survey results. J. Dent. 32 (5), 343350. Available from: https://www-sciencedirect-com.brad.idm.oclc. org/science/article/pii/S0300571204000430?via%3Dihub. Clarkson, T.W., Magos, L., Myers, G.J., 2003. The toxicology of mercury—current exposures and clinical manifestations. N. Engl. J. Med. 349 (18), 17311737. Cox, C.F., 1994. Evaluation and treatment of bacterial microleakage. Am. J. Dent. 7 (5), 293295. Available from: https://www.ncbi.nlm.nih.gov/pubmed/7986455. Craig, R.G., 1985. Restorative Dental Materials, seventh ed. Mosby. Davies, R.A., Ardalan, S., Mu, W., Tian, K., Farsaikiya, F., Darvell, B.W., et al., 2010. Geometric, electronic and elastic properties of dental silver amalgam γ-(Ag3Sn), γ1(Ag2Hg3), γ2-(Sn8Hg) phases, comparison of experiment and theory. Intermetallics 18 (5), 756760. Available from: https://doi.org/10.1016/j.intermet.2009.12.004 (accessed 10.03.18.). Dean, J.A., 2016. McDonald and Avery’s Dentistry for the Child and Adolescent, 10th ed. Elsevier. Delta Dental, 2012a. Dental amalgam or resin composite fillings?. ,https://www.deltadentalins.com/oral_health/amalgam.html.. Delta Dental, 2012b. What is plaque?. ,https://www.deltadentalins.com/oral_health/plaque. html.. Dental Science, 2016. Amalgam: properties, uses, toxicity and advantages. ,http://www.dental-science.com/amalgam/.. Donovan, T., Simonsen, R.J., Guertin, G., Tucker, R.V., 2008. Why gold castings are excellent restorations. J. Oper. Dent. 33 (2), 113115. Available from: https://www.jopdent. com/journal/editorial/openFile.php?takeFile 5 Volume_33__Issue_2.txt. Dubach, P., Caversaccio, M., 2011. Amalgam tattoo. N. Engl. J. Med. 364, e29. Available from: https://doi.org/10.1056/NEJMicm1011669. Espevik, S., 1977. Effect of trituration on dimensional changes of dental amalgam. Acta Odontol. Scand. 35 (5), 251255.
122 Advanced Dental Biomaterials FDA, 2017. About dental amalgam fillings. ,https://www.fda.gov/MedicalDevices/ ProductsandMedicalProcedures/DentalProducts/DentalAmalgam/ucm171094.htm#1.. FDI World Dental Federation, 2017. Minamata convention on mercury comes into force in august affecting use of dental amalgam. ,https://www.fdiworlddental.org/news/ 20170606/minamata-convention-on-mercury-comes-into-force-in-august-affecting-useof-dental.. Fong, H., Little, M.J., 2010. Dental Composites with Nano-Scaled Fillers. Nova Science Publishers, Inc. Available from: https://ebookcentral-proquest-com.brad.idm.oclc.org/lib/ bradford/detail.action?docID 5 3018712. Gay, D.D., Cox, R.D., Reinhardt, J.W., 1979. Chewing releases mercury from fillings. Lancet 1 (8123), 985986. Greener, E.H., 1979. Amalgam—yesterday, today, and tomorrow. Oper. Dent. 4 (1), 2435. Hasheminezhad, A., Zebarjad, S.M., Sajjadi, S.A., Rahanjam, L., 2012. Effect of copper content on compressive strength and microstructure of dental amalgams. Sci. Res. 4, 155159. Available from: https://doi.org/10.4236/eng.2012.43020. Helmenstine, A.M., 2017. Amalgam Definition and Uses. ThoughtCo. Available from: https://www.thoughtco.com/amalgam-definition-4142083. Hollenback, G.M., 1969. A consideration of amalgam. J. Calif. Dent. Assoc. 45, 1518. Hooghan, T.K., Pinizzotto, R.F., Watkins, J.H., Okabe, T., 1996. A study of a low copper dental amalgam by analytical transmission electron microscopy. J. Mater. Res. 11 (10), 24742485. Jayanthi, N., Vinod, V., 2013. Comparative evaluation of compressive strength and flexural strength of conventional core materials with nanohybrid composite resin core material an in vitro study. J. Indian Prosthodontic Soc. 13 (3), 281289. Knosp, H., Holliday, R.J., Corti, C.W., 2003. Gold in dentistry: alloys, uses and performance. Gold Bull. 36 (3), 93102. Available from: https://link.springer.com/content/pdf/ 10.1007/bf03215496.pdf (accessed 23.02.18.). Lygre, G.B., Haug, K., Skjærven, R., Björkman, L., 2016. Prenatal exposure to dental amalgam and pregnancy outcome. Community Dent. Oral Epidemiol. 44 (5), 442449. Available from: https://doi.org/10.1111/cdoe.12233. Mahler, D.B., Pham, B.V., Adey, J.D., 2009. Corrosion sealing of amalgam restorations in vitro. Oper. Dent. 34 (3), 312320. Available from: https://doi.org/10.2341/08-94. Manappallil, J.J., 2016. Basic Dental Materials., fourth ed. Jaypee Brothers Medical Publishers (P) Ltd.. Muliyar, S., Shameem, K.A., Thankachan, R.P., Francis, P.G., Jayapalan, C.S., Hafiz, K.A. A., 2014. Microleakage in endodontics. J. Int. Oral Health 6 (6), 99104. Murray, P.E., Hafez, A.A., Smith, A.J., Cox, C.F., 2002. Bacterial microleakage and pulp inflammation associated with various restorative materials. Dent. Mater. 18 (6), 470478. Available from: https://www.ncbi.nlm.nih.gov/pubmed/12098576. Nicholson, J.W., 2002. The Chemistry of Medical and Dental Materials, Royal Society of Chemistry.. Royal Society of Chemistry. Available from: https://ebookcentral-proquestcom.brad.idm.oclc.org/lib/bradford/reader.action?docID 5 1185355&query. Okabe, T., Cahn, R.W., 1990. Encyclopedia of Materials Science and Engineering. Pergamon Press, Oxford. Oral Health Centre, 2016b. Problems After Dental Fillings. Boots Web MD.. Available from: https://www.webmd.boots.com/oral-health/guide/problems-dental-fillings. Pereira, T., 2016. Silver amalgam: a clinician’s perspective. J. Dent. Res. 4 (2), 2530. Available from: http://www.jresdent.org/article.asp?issn 5 2321-4619;year 5 2016; volume 5 4;issue 5 2;spage 5 25;epage 5 30;aulast 5 Pereira.
Dental amalgam 123 Powers, J.M., Wataha, J.C., 2017. Dental Materials, Foundations and Applications., 11th ed. Elsevier. Rathore, M., Singh, A., Pant, V.A., 2012. The dental amalgam toxicity fear: a myth or actuality. Toxicol. Int. 19 (2), 8188. Available from: https://doi.org/10.4103/09716580.97191. Sarkar, N.K., Eyer, C.S., 1987. The microstructural basis of creep of c1 in dental amalgam. J. Oral Rehabil. 14 (27). Shivrayan, A., Kumar, G., 2015. Comparative study of mechanical properties of direct core build-up materials. Contemp. Clin. Dent. 6 (1), 16. Singh, H., 2015. Remembering Sir G.V. Black. Indian J. Dent. 6 (3), 147148. Available from: https://doi.org/10.4103/0975-962X.163047. Sivakumar, A., Valiathan, A., 2006. Dental ceramics and ormocer technology—navigating the future. Trends Biomater. Artif. Organs 20 (1), 4043. Thomas, G.P., 2013. Amalgam-chemical composition, mechanical properties and common applications. AZO Materials. Available from: https://www.azom.com/article.aspx? ArticleID 5 8081. (accessed 19.01.18.). Ucar, Y., Brantley, W.A., 2011. Review article, biocompatibility of dental amalgams. Int. J. Dent. 2011, 17. Available from: https://doi.org/10.1155/2011/981595. Vanishree, H.S., Shanthala, B.M., Bobby, W., 2015. The comparative evaluation of fracture resistance and microleakage in bonded amalgam, amalgam, and composite resins in primary molars. Indian J. Dent. Res. 26 (5), 446450. Wang, Y., Darvell, B., 2007. Failure mode of dental restorative materials under Hertzian indentation. Dent. Mater. 23 (10), 12361244. Yahya, N., Puspitasari, P., Latiff, N.R.A., 2013. Hardness improvement of dental amalgam using zinc oxide and aluminium oxide nanoparticles. In: Ochsner, A., da Silva, L., Altenbach, H. (Eds.), Characterization and Development of Biosystems and Biomaterials. Advanced Structured Materials, vol. 29. Springer, pp. 932. Yap, A.U.J., Ng, B.L., Blackwood, D.J., 2004. Corrosion behaviour of high copper dental amalgams. J. Oral Rehabil. 31, 595599. Zimmerli, B., Strub, M., Jeger, F., Stadler, O., Lussi, A., 2010. Composite materials: composition, properties and clinical applications, a literature review. Schweiz Monatsschr Zahnmed 120, 972979. Further reading Ajlan, S.A., Aldahmash, A.M., Ashri, N.Y., 2015. Dental pulp stem cells, biology and use for periodontal tissue engineering. Saudi Med. J. 36 (12), 13911399. Amalgam-chemical composition, mechanical properties and common applications. AZO Mater. ,https://www.azom.com/article.aspx?ArticleID 5 8081.. American Dental Association, 2012a. Primary tooth development [Image]. ,http://www. mouthhealthy.org/en/az-topics/e/eruption-charts.. American Dental Association, 2012b. Permanent tooth development [Image]. ,http://www. mouthhealthy.org/en/az-topics/e/eruption-charts.. Beek, G.C.V., 1983. Dental Morphology, an Illustrated Guide, second ed. Billing & sons Ltd, Worcester, pp. 4547. Berkovitz, B.K.B., Holland, G.R., Moxham, B.J., 2018. Oral Anatomy, Histology and Embryology., fifth ed. Elsevier.
124 Advanced Dental Biomaterials British Orthodontic Society, 2014. Patient information leaflet, tooth transplants. https://www. bos.org.uk/Portals/0/Public/docs/PILs/toothtransplantsmarch14.pdf. Chano, L., Clokie, C.M.L., Yau, D.M., 2001. Autogenous tooth transplantation: an alternative to dental implant placement. J. Can. Dent. Assoc. 67 (2), 9296. Cohen, A.S., Pogrel, M.A., Shen, T.C., 1995. Transplanting teeth successfully: autografts and allografts that work. J. Am. Dent. Assoc. 126 (4), 481485. Davenport, T., 2017. What are the most common dental problems?. ,https://www.verywell. com/top-common-dental-problems-1059461.. Dentagama, 2016. Image of the restorative prothesis: inlay, onlay, overlay, pinlay. ,https:// dentagama.com/news/what-is-the-difference-between-inlay-onlay-overlay-and-pinlay. (accessed 25.02.18.). D’Aguino, R., Graziano, A., Laino, G., Papaccio, G., 2008. Dental pulp stem cells: a promising tool for bone regeneration. Stem Cells Rev. 4 (1), 2126. Harris, E.F., 2016. Odontogenesis. In: Irish, J.D., Scott, G.R. (Eds.), A Companion to Dental Anthropology. John Wiley & Sons, pp. 145149. Healthline, 2015. Incisors. ,https://www.healthline.com/human-body-maps/incisor.. Hoffman, M., 2009. Picture of the Teeth. WebMD [Image] . Available from: https://www. webmd.com/oral-health/picture-of-the-teeth#1. Hu, J.C.-C., Simmer, J.P., 2001. Dental enamel formation and its impact on clinical dentistry. J. Dent. Educ. 65 (9), 896905. Huang, G.T.J., 2009. Pulp and dentin tissue engineering and regeneration: current progress. Regener. Med. 4 (5), 697707. Available from: https://doi.org/10.2217/rme.09.45. Jethmalani, Y.D., Potdar, P.D., 2015. Human dental pulp stem cells: applications in future regenerative medicine. World J. Stem Cells 7 (5), 839851. Keels and The American Academy of Pediatrics Section on Oral Health, 2014. Management of dental trauma in a primary care setting, the section on oral health. Pediatrics 133 (2), e468. Available from: http://pediatrics.aappublications.org/content/133/2/e466. Kim, N., Cho, S.G., 2013. Clinical applications of mesenchymal stem cells. Korean J. Intern. Med. 28 (4), 387402. Available from: https://www.ncbi.nlm.nih.gov/pmc/articles/ PMC3712145/. Konigsberg, S., Northway, W.M., 1980. Autogenic tooth transplantation: the “state of the art”. Am. J. Orthod. Dentofacial Orthop. 77 (2), 146162. MacConaill, M.A., 2017. Joint. Encyclopaedia Britannica, Inc. Available from: https://www. britannica.com/science/joint-skeleton/Fibrous-joints#ref470986. Marrelli, M., Tatullo, M., Shakesheff, K.,M., White, L.J., 2015. Dental pulp stem cells: function, isolation and applications in regenerative medicine. J. Tissue Eng. Regener. Med. 9 (11), 12051216. Matalova, E., Antonarakis, G.S., Sharpe, P.T., Tucker, A.S., 2005. Cell lineage of primary and secondary enamel knots. Dev. Dyn. 233 (3), 754759. Mayo Clinic, 2017a. TMJ Disorders. Mayo Clinic. Available from: https://www.mayoclinic. org/diseases-conditions/tmj/symptoms-causes/syc-20350941. Mayo Clinic, 2017b. Mouth cancer. ,https://www.mayoclinic.org/diseases-conditions/ mouth-cancer/symptoms-causes/syc-20350997.. Mayo Clinic, 2017c. Periodontitis. ,https://www.mayoclinic.org/diseases-conditions/periodontitis/symptoms-causes/syc-20354473.. NHS Choices, 2016. Tooth decay. ,https://www.nhs.uk/conditions/tooth-decay/.. National Institute of Dental and Craniofacial Research, 2015. Oral cancer: causes and symptoms & the oral cancer exam. ,https://www.nidcr.nih.gov/OralHealth/Topics/ OralCancer/AfricanAmericanMen/CausesSymptoms.htm..
Dental amalgam 125 Newman, T., 2016. Bad breath (halitosis): causes, diagnosis and treatment. Medical News Today . Available from: https://www.medicalnewstoday.com/articles/166636.php. Norris, J., 2010. Tooth Enamel: Nature’s Crowning Achievement. University of California San Francisco. Available from: https://www.ucsf.edu/news/2010/08/6000/tooth-enamelresearch-combines-science-and-engineering. Open Stax, 2016. Anatomy & physiology. Open Stax CNX. Available from: http://cnx.org/ contents/14fb4ad7-39a1-4eee-ab6e-3ef2482e3e22@8.24. Oral Health Centre, 2016a. Tooth Enamel Erosion and Restoration. Boots Web MD. Available from: https://www.webmd.boots.com/oral-health/guide/tooth-enamel-erosionrestoration. Oral Health Centre, 2017. Temporomandibular Joint Disorders (TMJ and TMD). Boots Web MD. Available from: https://www.webmd.boots.com/oral-health/guide/temporomandibular-joint-disorders-tmj-tmd?page 5 3. Science Learning Hub, 2011. Rate of digestion. ,https://www.sciencelearn.org.nz/resources/ 1834-rate-of-digestion.. Skaria, S., 2016. Fibrous Joint—Gomphosis [Image]. ,https://www.slideshare.net/ SajanScaria/artrology.. Slootweg, P.J., 2007. Dental Pathology: A Practical Introduction.. Springer. Stay, F., 2014. How your teeth affect your digestive system. ,http://www.totalhealthmagazine.com/Dental-Health/How-Your-Teeth-Affect-Your-Digestive-System.html.. Swindler, D.R., 2002. Primate Dentition: An Introduction to the Teeth of Non-human Primates.. Cambridge University Press, Cambridge. The editors of Encyclopaedia Britannica, 2007. Cementum. Encyclopaedia Britannica, Inc. Available from: https://www.britannica.com/science/cementum. The editors of Encyclopaedia Britannica, 2014. Dentin. Encyclopaedia Britannica, Inc. Available from: https://www.britannica.com/science/dentin. Watson, S., 2017. What is the Function of Tooth Pulp. Very Well. Available from: https:// www.verywell.com/tooth-pulp-dental-terms-1059180. Yau, E.C.K., 2009. Tooth autotransplantation as a treatment option. Dent. Bull. 14 (6), 2124. Available from: http://www.fmshk.org/database/articles/03db05_2.pdf. Zimmermann, K.A., 2016. Digestive System: Facts, Function & Diseases. Live Science. Available from: https://www.livescience.com/22367-digestive-system.html.
Resin-based dental composites for tooth filling 7 Ahmed El-Banna1, Dalia Sherief1 and Amr S. Fawzy2 1 Dental Biomaterials Department, Faculty of Dentistry, Ain-Shams University, Cairo, Egypt, 2UWA Dental School, University of Western Australia, Nedlands, WA, Australia Chapter Outline 7.1 Introduction 128 7.2 General composition 7.2.1 7.2.2 7.2.3 7.2.4 7.2.5 128 The resin matrix 128 Fillers 130 Silane coupling agent 134 Initiatoraccelerator system 134 Pigments and other components 136 7.3 Classification of resin composites 7.3.1 7.3.2 7.3.3 7.3.4 136 According to the fillers size and distribution 136 According to the composite consistency 137 According to the packing (placement) technique 138 According to the curing techniques 138 7.4 Clinical indications of resin composites 7.5 Properties and limitations 139 7.5.1 7.5.2 7.5.3 7.5.4 7.5.5 7.5.6 7.5.7 7.5.8 138 Degree of conversion 139 Polymerization shrinkage and polymerization shrinkage stresses Physical properties 144 Esthetic properties 144 Mechanical properties 145 Biocompatibility 145 Degradation 147 Clinical durability 147 140 7.6 Attempts for resin composite improvement 148 7.6.1 Regarding material formulation 148 7.6.2 Regarding manipulation 154 7.6.3 Regarding both material formulation and manipulation 156 7.7 Guidelines and recommendations for future laboratory and clinical researches 165 7.7.1 Guidelines for laboratory evaluation of resin composite (mechanical behavior and technique sensitivity) 166 7.7.2 Recommendations for future clinical studies 166 References 171 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00007-4 Copyright © 2019 Elsevier Ltd. All rights reserved.
128 7.1 Advanced Dental Biomaterials Introduction In the mid-20th century acrylic resins were launched in the field of esthetic restorative dentistry to overcome the limitations of silicate cement. In 1955 Buonocore developed the concept of using phosphoric acid as an attempt to enhance the bonding of acrylic resin restorations to enamel. Phosphoric acid induced selective dissolution within the enamel creating microporosities into which resin flows and polymerizes creating micromechanical interlocking. In 1962 Bowen developed the high molecular-weight bisphenol-aglycidyldimethacrylate (Bis-GMA) monomer as a trial to upgrade the physical properties of acrylic resins, as their monomers only allowed linear chain polymers to be formed. Resin composites were developed in 1962 by combining dimethacrylates (epoxy resin and methacrylic acid) with silanized quartz powder. Early chemically cured composites required the mixing of both base and catalyst pastes together leading to inaccurate proportioning and mixing thus affecting the properties. From 1970, composite materials cured by electromagnetic radiation emerged, eliminating the need for mixing and its drawbacks. Ultraviolet light sources were used first to provide the required polymerization energy. The limited polymerization depth and induced side effects of ultraviolet radiation led to its replacement by visible light (427491 nm), which is currently in use and undergoing further development (Fig. 7.1) (Leprince et al., 2013; Ferracane, 2011). During the past decade, resin composites have become the most widely used restorative materials, gradually replacing amalgam restorations, which have been the standard restorative material for more than 100 years. Such transition was not only due to the favorable esthetics of resin composites but also due to the highly conservative cavity preparation compared to that of amalgam. Since amalgam toxicity is still a point of debate, there is a trend toward reducing the use of mercury or even banning mercury-containing products in some countries, thus dental resin composites remain a main interest of dental material research (Rasines Alcaraz et al., 2014). 7.2 General composition Resin composites consist of three main components which are the resin matrix (organic part), the inorganic fillers (inorganic part), and the coupling agent. The resin matrix includes the monomers, initiators, inhibitors, pigments, and modifiers. 7.2.1 The resin matrix The monomer systems are considered as the primary foundation of the composite resin formulation. Bis-GMA remains to be the widely used monomer for constructing today’s composites; whether alone or in combination with urethane dimethacrylate (UDMA). Such monomers are slightly viscous and have to be mixed with low molecular-weight diluent monomers so that a clinically workable consistency may
Figure 7.1 Evolution of dental resin composite throughout the last 65 years.
130 Advanced Dental Biomaterials be produced upon addition of the fillers. Low molecular-weight compounds with difunctional carbon double bonds, for example, triethylene glycol dimethacrylate (TEGDMA), are added by the manufacturer to reduce and control the viscosity of the composite. In addition to the conventional dimethacrylates, composite resin matrix may contain acidic carboxylic groups. Such composites are known as compomers. The name is derived from the terms composite and glass ionomer. Self-adhesive composite resin cements is another group of acidic monomers containing composites. Resin cements are formed by a combination of usually dimethacrylate monomers and acid-functionalized monomers, such as methacryloyloxydecyl dihydrogen phosphate (MDP), glycerol phosphate-dimethacrylate (GPDM), or dipentaerythritol Penta acrylate phosphate (PENTA). Another category of resin composites is those containing anionically ring-opening polymerizable cycloaliphatic epoxides. The term “silorane” was introduced to represent such composites whose monomer systems include both siloxane and oxirane structural units. The main advantage of these resins is the reduced polymerization shrinkage. Attempts to modify and improve the properties of resin composite matrix (Fig. 7.2) will be discussed later in Section 7.6 (Baroudi and Rodrigues, 2015). 7.2.2 Fillers Inorganic fillers represent the dispersed phase of resin-based composites whose nature plays a significant role in determining their physical and mechanical properties. Fillers vary according to their chemistry, shape, and size; such variations are reflected in the properties of composites. Inorganic fillers are generally added to enhance the strength and modulus of the organic resin matrix. However fillers also reduce the thermal expansion coefficient, water sorption, and polymerization shrinkage, as well as imparting radiopacity and enhancing the esthetic properties. Silicon dioxide was the first filler used in dental resin composites. It is termed silica, or quartz when present in its crystalline form. It is also considered the basis for other types of glass fillers that are formed of different silicate compounds. Silica has a refractive index which is lower than that of the traditional BisGMA/TEGDMA resin matrix, making it more opaque than other fillers used in such a resin system. Alkaline glass fillers are also used in many of the currently available commercial composites such as TetricEvoCeram (IvoclarVivadent, Liechtenstein), Grandio (Voco, Germany), Esthet-X (Dentsply Caulk, United States), and Herculite XRV (Kerr, United States). They are composed mainly of silion dioxide, however they contain a fraction of alkaline oxides such as barium oxide (BaO) and strontium oxide (SrO), which incorporate into the silica network causing interruption of the structure. The main advantage of this material, however, is the inclusion of heavier elements into the filler imparting radiopacity of the material without needing to add separate radiopacity agents, such as ytterbium or yttrium fluoride. Another important advantage of such particles is superior optical properties since they have a refractive index higher than silica, making it closer to that of traditional resin matrix, producing more transparent composites. Many forms of
Resin-based dental composites for tooth filling 131 Figure 7.2 Chemical structure of different resin systems used in dental resin composite. alkaline glasses have been used, including barium borosilicates, barium aluminosilicates, barium aluminum borosilicates, and strontium silicates (Habib et al., 2015). Other glasses fillers containing calcium, sodium, or phosphorus oxides have also been tested as fillers in resin composites. Such glasses are known as bioactive glass. The aim of using such fillers is producing what is known as smart composites that help in the remineralization of the tooth structure. Smart composites and other composite fillers modification will be discussed in Section 7.6.1.2. 7.2.2.1 Filler size and filler loading Increasing the filler loading enhances the properties related to them, such as high mechanical properties, low polymerization shrinkage, and better esthetics. Filler sizes distribution is used to maximize the filler loading, where smaller particles fill the spaces between the larger fillers. This reduces the void space between the larger fillers and yields higher filler loading.
132 Advanced Dental Biomaterials Using small particles also improve esthetics and polishability of the resin composite restoration. Filler sizes larger than the visible light wavelength were found to cause light scattering. Such scattering enhances opacity and produces a rough surface when the fillers are exposed at the resin composite surface. Surface roughness of the resin composite restoration causes accumulation of food debris, stains, and dental plaque. It should be considered that the smaller the filler particle size the higher the surface/volume ratio present to form hydrogen bonding with the resin matrix. This increases the viscosity of the resin composite and in turn its resistance to mixing and manipulation. Thus the smaller the fillers size the less the filler fraction that can be added. 7.2.2.2 Prepolymerized filler particles These specialized fillers are produced by adding very high concentrations of inorganic micro- or nanofillers to a resin monomer under high heat and pressure. Once polymerized, the resultant blocks are ground into macrosized filler particles, with size ranging from 30 to 65 μm. The resin composite material is made up by adding a combination of these larger macrosized filler particles and individual microfillers in a quantity of uncured monomer. This dramatically minimizes the amount of uncured monomer, thus decreasing polymerization shrinkage (Fig. 7.3). 7.2.2.3 Nanofilled resin composite Nanofilled composites have been introduced to the dental market as a product of nanofiller technology. They are formulated with nanomer and nanocluster filler particles. Nanomers are distinct nonagglomerated particles of 2075 nm in size, while nanoclusters are slightly bound agglomerates of nano-sized particles (Fig. 7.4). The mixture of nano-sized particles and nanoclusters decreases the interstitial spacing of the filler particles, thus allowing higher filler loading, better physical properties, and high surface polish retained for a considerable period of time. The first commercial nanofilled resin composite (Filtek Supreme, 3M ESPE) was launched in 2002. As for nanohybrid composites, they contain large-sized fillers (0.45 μm) together with nanometer-sized fillers. They are considered hybrid composites rather than true nanofilled composites (Ferracane, 2011). Figure 7.3 Diagrammatic representation of prepolymerized fillers fabrication.
Resin-based dental composites for tooth filling 133 Figure 7.4 Schematic diagram for nanofillers and nanoclusters. Source: Courtesy 3M ESPE. There are several aims behind incorporating nanofillers in dental composites. First, the size of nanofillers is less than the wavelength of visible light (400800 nm), which provides the ability to develop highly translucent resin composites. Using nanoparticles alone to form highly filled composites will result in poor rheological properties (high viscosity). In order to overcome such problem nanoclusters were designed. The nanoclusters are made of lightly sintered nanomers. The clusters allow having a wide size distribution varying from 100 nm to submicron. Nanofilled composites have mechanical strength similar to that of microhybrid composites, and at the same time retain a high surface finish like microfilled composites. In the nanofilled composite, the nanoclusters shear at a similar rate as the resin matrix during wear process, thus maintaining a long-term polished surface (Ferracane, 2011). 7.2.2.4 Fiber-reinforced composites The wide use of fiber-reinforced polymers in many industrial fields encouraged the investigation of such materials in dental composites. Glass fibers are the most commonly used fibers in dental composites due to their low cost as well as their chemical composition and refractive index that are nearly similar to that of the silica fillers. Glass fibers also showed good adhesion to the resin matrix when treated with silane coupling agents. Electrical or E-glass fibers are those most commonly used in dental composites; they are characterized by their high strength and high electrical resistivity. Different types of polymeric fibers were also investigated to improve the overall properties of the composites, such as polyethylene fibers and aramid fibers. Aramid fibers significantly increased the strength properties of resin-based composites; however, the yellow color of these fibers affected the esthetic appearance, limiting their use in resin-based composites. The incorporation of polyethylene fibers generally enhanced the mechanical properties of the resin composites; however, their low surface energy and poor adhesion of the resin matrix limited their reinforcing effect. Plasma etching was an attempt to improve the wetting of these fibers and their mechanical interlocking with the resin matrix. Various kinds of nanotubes have been also investigated as reinforcements for dental resin composites. Carbon nanotubes were the most commonly used. The
134 Advanced Dental Biomaterials addition of small percentages of these carbon nanotubes to a commercial composite increased the flexural strength by 23%. Double-walled halloysite nanotubes were also explored to reinforce unfilled and glass-filled dental resin composites. The addition 12.5 wt.% of halloysite nanotubes showed a twofold increase in the fracture resistance of these composites; however, it produced a slight increase in the flexural strength. More rigid whiskers were also attempted. These whiskers included ceramics, carbides, zinc oxide, zirconia/silica, silicon nitride, and silicon carbide. Whiskers fused with silica nanoparticles were also tried as resin composite fillers. These silica-fused whiskers showed 33% decrease in wear and a twofold increase in flexural strength compared to the control (Nayar et al., 2015; Khan et al., 2015). 7.2.3 Silane coupling agent Surface treatment of the resin composites’ filler particles with silane coupling agents is essential to maintain a strong and durable bond between the polymer matrix and the filler particles. Silanized fillers were shown to enhance tensile strength and tensile bond strength of resin composites compared to nonsilanized filler particles. Filler particles showed improved wettability and dispersion within the resin matrix thus improving the physical and mechanical properties of resin composites. The water uptake by silanized filler particles was also found to be significantly lower than nonsilanized filler particles. The chemical formulation of organofunctional silanes is as below: XUUUSiUUUðORÞ3 where X is an organofunctional (methoxy) group and R is a hydrolyzable alkyl group. The alkyl group is hydrolyzed to silanol to react with the fillers’ mineral surface, on the other hand the organofunctional group copolymerizes with the resin matrix. Dental silane products are based on 3-methacryloxypropyltrimethoxy silane, an organofunctional trialkoxysilane which reacts via its alkoxysilane groups with the filler and itself, and with the resin matrix via its methacrylate functional group. An important factor affecting the efficiency of the silanation process is the thickness of the silane film. The closest monolayer to the filler surface chemically bonds to this surface. As the silane thickness increases, its layers become more disorganized and physicosorbed on the surface of the fillers increasing the risk of weak bond formation. A critical thickness of the silane layers is thus required, which is controlled by the quantity of silane used. 7.2.4 Initiatoraccelerator system The function of the initiatoraccelerator system is to polymerize and cross-link the composite resin into a hardened form. The polymerization reaction can be triggered by light (light-cured), chemicals (self-cured, auto-cured), or both chemical and light-curing (dual-cured).
Resin-based dental composites for tooth filling 135 7.2.4.1 Self-cured composites They are formed of a catalyst paste containing benzoylperoxide (BPO) initiator and a base paste containing an aromatic tertiary amine, commonly N,N-dimethyl-p-toluidine. Such tertiary amine (TA) is required to induce the decomposition of BPO at an ambient temperature. 7.2.4.2 Light-cured composites They are supplied as a single paste whose setting is triggered by visible light. Such resin composites materials must be protected from premature curing when stored in the dental office. They are packaged in opaque, plastic syringes or unit-dose capsules, known as compules. Photoinitiators The most widely used photoinitiator in light-cured resin composites is a combination of camphorquinone (CQ) with TAs. When CQ absorbs the light, it interacts with the TAs forming a photoexcited complex with free radicals on both CQ and TAs. The free radicals attack the CQC bonds of monomers, forming new radicals with a longer polymer chain (propagation step). Such propagation of the polymer chain continues until the reaction process terminates. The peak sensitivity of CQ is near 470 nm in the blue wavelength range (Fig. 7.5). Although CQ initiator systems have good acceptance, the yellow-colored CQ may adversely affect the composite color. Another major problem of CQ, is that its peak absorption is within the visible range, causing its photopolymerization under ambient light (fluorescent lamps and dental lamps) and giving a short working time. Other initiators are now incorporated in resin composites, as an attempt to solve the problems of CQ/TA, such as phenyl-propanedione (PPD) and diphenyl (2,4,6-trimethylbenzoyl) phosphine oxide (TPO). Figure 7.5 Free radical formation by light activation of camphorquinone.
136 Advanced Dental Biomaterials Lucirin—TPO (2,4,6-trimethylbenzoyldiphenylphosphine oxide) and phenylpropanedione photoinitiator systems TPO is an acyl phosphine oxide, that when cured with halogen or polywave light control units (LCUs), TPO results in higher degree of conversion (DC) than CQ/TA. TPO is especially useful for the extra white shades of resin composites needed in bleached teeth since it gets rid of the unwanted yellow effect of CQ/TA. Its absorption spectrum ranges from 380 to nearly 425 nm. PPD is also a photoinitiator that tends to reduce the yellowing effect caused by CQ, its absorption spectrum is from below 350 to nearly 490 nm. 7.2.4.3 Dual-cured composite They are formed of two pastes (base paste and catalyst paste) similar to that of selfcured composites, however their base paste contains the CQ/TA photoinitiation system required for visible light activation. An aromatic TA is also present in the base paste for chemical polymerization. When the two pastes are mixed together, they set up very slowly via the chemical-cured mechanism. The curing rate then speeds up on “command” via light-curing. The main advantage of this system is the guarantee of complete polymerization, even if the light-cure is insufficient. The major drawback is the porosity caused by the mixing of the two pastes; however, this problem has been solved to a great extent via the use of mixing syringes. Most resin-based luting cements or core materials are dual-cured, combining both lightand chemical-cure mechanisms (Santini et al., 2013; Kwon et al., 2012). 7.2.5 Pigments and other components Inorganic oxides are added in small percentages to provide shades that match the majority of tooth shades. Iron oxide represents the most common pigment. Numerous shades are supplied at present extending from very light shades to yellow and gray. Fluorescent agents are also provided to simulate the appearance of natural teeth. These are dyes or pigments that absorb light in the ultraviolet and violet region (usually 340370 nm) and re-emit it in the blue region (typically 420470 nm). 7.3 Classification of resin composites 7.3.1 According to the fillers size and distribution Resin-based composites can be divided into fiber-reinforced composites (FRC) and particulate filler composites. FRCs can be further categorized into long-fiber and short-fiber composites based on the fiber length. Fiber orientation is another classification into which resin composites can be further classified with continuous unidirectional and bidirectional fibers. Regarding the particulate filler composites, a beneficial method of classifying dental resin composites is by the particle size and the particle-size distribution (Fig. 7.6 and Table 7.1).
Resin-based dental composites for tooth filling 137 Figure 7.6 Classification of resin composites based on filler size, shape, and distribution. Table 7.1 Classification of resin composites based on particle size and particle-size distribution. Type of composite Filler size Conventional (large fillers) Hybrid (large fillers) Small/fine filled Midfilled Hybrid (midfilled) Minifilled Hybrid (minifilled) Homogeneous microfilled Heterogeneous microfilled Nanofilled composites Nanohybrid 150-μm silica or glass fillers A combination of 120 μm glass fillers and 40 nm silica fillers 0.110 μm 110 μm A combination of 0.110 μm glass fillers and 40 nm silica fillers 0.11 μm A combination of 0.12 μm glass fillers and 40 nm silica fillers 40-nm silica A combination of 40-nm silica particles together with prepolymerized resin particles containing 40-nm silica ,100-nm silica or zirconia nanoparticles together with nanoclusters A combination of 0.12-μm glass or prepolymerized resin microparticles together with # 100-nm nanoparticles 7.3.2 According to the composite consistency Composite materials could also be classified according to their consistency into flowable and packable composites. Flowable composites are designed to be dispensed from very fine bore syringes into limited spaces providing better adaptation. Reducing the filler content and/or addition of other modifying agents, such as surfactants, are the main methods used to reduce consistency of flowable composites. Large reduction in filler content should be avoided as it would significantly reduce
138 Advanced Dental Biomaterials the mechanical properties and increase the polymerization shrinkage of flowable composites. Packable composites are designed to provide adequate resistance to condensation in order to avoid slumping and allow the formation of tight interproximal contacts. Packable composites develop their thicker consistency through modification in filler size distributions or the addition of other types of fillers, such as fibers, but not through increasing the filler loading (Lutz et al., 1983). 7.3.3 According to the packing (placement) technique Based on the packing technique resin composites may be also classified into incremental and bulk-packed composite. Polymerization shrinkage, as well as the limited curing depth makes it generally recommended to place composite resin in 2 mm increments or less. Incremental placement however may be accompanied with porosity and voids especially within large posterior composite restorations. In order to simplify and speed up the placement of large posterior RBCs, manufacturers have produced resin composite materials that can be placed as single thick increments (4 mm or more). Such resin composites are known as bulk-fill RBCs. 7.3.4 According to the curing techniques Resin composites may be further classified according to the curing techniques into light-cured, chemical-cured, and dual-cured composites (previously mentioned in Section 7.2.4). Resin composites can be categorized into direct and indirect resin composites. Recent direct composite resins show superior optical and mechanical properties, however their use in larger posterior restorations is still a challenge since polymerization shrinkage exists as a problem in cavities with high C-factor. Indirect composites are fabricated on a die and additionally cured extraorally using combinations of light, pressure, and heat to increase the degree of polymerization, mechanical properties, wear resistance, and also lessen the polymerization shrinkage. Such indirect restorations are bonded using resin cements (Kwon et al., 2012). 7.4 Clinical indications of resin composites Resin-based composite restorations are widely used in the dental field. The clinical indications of each type of resin composite are dictated by the amount of stresses applied on the restoration as well as the esthetic requirements. Regarding high stress bearing areas requiring adequate polishability (classes I, II, III, IV), either hybrid large fillers or hybrid midifil resin composites are used. As for moderate stress bearing areas requiring optimal polishability (classes III, IV), hybrid minifilled and nanohybrid composites are indicated. For low stress and subgingival areas that require high luster and polishability, microfilled resin composite
Resin-based dental composites for tooth filling 139 is the choice. Nanofilled composites are used as anterior restorations that require high translucency and surface finish Packable composites are indicated in situations where good condensability is required (classes I, II). On the other hand flowable composites are required where improved flow is needed in areas of limited accessibility (class II) (Hervas-Garcia et al., 2006). 7.5 Properties and limitations 7.5.1 Degree of conversion Properly cured resin is considered the principle factor for a successful resin-based restoration. This is accomplished mainly by having an adequate DC, which is presented by the percentage of carboncarbon double bonds (CQC) that are transformed to carboncarbon single bonds (CaC). DC for dimethacrylate polymers ranges between 43% and 75%. A high DC is essential to improve the physical, mechanical, and biological properties of the resin-based composites. Several factors affect the degree of polymerization of resin composites (Braden, 1974). 7.5.1.1 Effect of resin shade Various shades and translucencies of resin composites are available to allow better matching with adjacent tooth structures, thus improving the esthetic appearance of the restoration. Resin composites with darker and more opaque shades prevent adequate light transmission and thus will have a lesser DC. 7.5.1.2 Effect of resin increment thickness Placement of resin composites in more than 2 mm increment thickness causes significant reduction in DC. This is due to the limited curing depth of light-curing units since they do not exceed 2 mm. Bulk-fill composites however permits resin composite placement in increments up to 4 mm while preserving an adequate DC. This is due to higher translucency of bulk-fill RBC that allows more light transmission compared to conventional resin composites. 7.5.1.3 Effect of light-curing system Quartz Tungsten Halogen (QTH) is formed of a quartz bulb containing halogen, iodine, or bromine gas, and a tungsten filament. Light is produced when electric current passes through the tungsten filament. The light emitted is a white light with a wavelength of broad spectrum. Light is filtered by a filter present inside the LCU to eliminate the ineffective wavelengths. Filtered light has a wavelength of 400500 nm, which is compatible with CQ, the most commonly used photoinitiator. A major disadvantage of the
140 Advanced Dental Biomaterials QTH LCU is that the useless light produced by the bulb results in an elevated operating temperature, reducing the lifetime of the bulb (to 100 hours). Fans are essential to reduce such elevated temperature. The light-emitting diode (LED) is the latest light-curing unit submitted to the dental market. LED technology uses junctions of doped semiconductors rather than a bulb to produce light. The light emitted has a narrower spectrum with less heat generation than a QTH LCU, but is within the range of CQ. Filters and fans are therefore, not required in LED LCUs. A LED light unit has a life span of about 10,000 hours. Newly introduced LED LCUs emit lights with two or more wavelength ranges (polywave LED LCUs). They produce both a shorter violet wavelength and a longer blue wavelength. Violet light activates photoinitiators sensitive to light within the range of 350420 nm wavelength while blue light activates photoinitiators (mainly CQ), with maximum absorbance near 468 nm. These polywave LED LCUs are the method of choice to activate a broader range of photoinitiators (Santini, 2010). 7.5.1.4 Effect of light-curing tip distance from RBC surface A decrease in radiant exposure of the light occurs on increasing the curing distance between the composite surface and the light cure. This reduction is due to cusp height and cavity depth, thus causing the curing tip to be at a distance from the cured resin composite. The distance from the light-curing tip and the gingival floor of a proximal box in deep Class II RBC restorations may reach 6 mm or more; a 50% reduction of light irradiance was reported by increasing the curing distance from 0 to 6 mm. 7.5.1.5 Effect of cavity location The location of some resin-based restorations, such as on the buccal or lingual surfaces of the second molar, makes it impossible to position the LCU tip directly and at 90 over the RBC restoration surface. This can affect the DC of the cured RBC increment. 7.5.1.6 Effect of light-curing duration For a 2-mm resin composite increment to have adequate polymerization, it should receive a radiant exposure within the range of 1624 J/cm2. This energy (E) is determined by multiplying the irradiance level (I) coming from the LCU (mW/cm2) by its duration (T). A longer curing time is thus needed when the irradiance level is lower. 7.5.2 Polymerization shrinkage and polymerization shrinkage stresses Polymerization shrinkage is one of the main drawbacks of resin-based composite restorations. During polymerization the conversion of monomer molecules into a closely packed polymer network results in volumetric shrinkage. Polymerization of dimethacrylate-based composites is associated with approximately 2%6% volumetric shrinkage (Braga et al., 2005).
Resin-based dental composites for tooth filling 141 Figure 7.7 Consequences of polymerization shrinkage stress. 7.5.2.1 Polymerization shrinkage stress Resin composite strain is constrained by the bonding of the material to the tooth structure, as a result shrinkage expresses itself as stress. The resulting built-up stresses may lead to marginal leakage, cuspal deflection, and cracking of enamel. This causes restoration fracture and secondary caries leading to clinical failure of the restoration (Fig. 7.7). 7.5.2.2 Factors affecting polymerization shrinkage stresses in dental composites Volumetric shrinkage The amount of composite volumetric shrinkage is determined by its filler content, DC, and monomer composition of the resin matrix. This could be explained by the fact that the amount of polymer matrix decreases by increasing the filler fraction and hence reducing the amount of volumetric shrinkage. Monomer chemistry may obviously affect the developed volumetric shrinkage. Higher TEGDMA/Bis-GMA ratios result in higher stress values due to increased volumetric shrinkage, as a result of enhanced conversion. Diluents monomers such as TEGDMA have lower molecular weight than Bis-GMA, thus they increase the amount of polymerizable carbon double bonds leading to more shrinkage.
142 Advanced Dental Biomaterials Several modifications have been made in the monomer systems of resin composite materials to decrease the polymerization shrinkage. Various low-shrinkage dental composites have been commercially marketed (discussed in Section 7.6.1.1). Viscoelastic behavior and polymerization kinetics Resin composites are solids with a unique viscoelastic behavior. They are characterized by their ability to flow at early polymerization stages followed by an increased elastic modulus on further polymerization. At the start of the polymerization reaction, composite resin is in aplastic viscous form allowing it to flow. This stage is known as the pre-gel stage, during which the resin monomers can still move or “slip” within the organic matrix. Larger molecules are formed as the polymerization process proceeds and the composite hardens into a solid form. At a certain point no further monomer movement is possible; such point is known as the gel point. It denotes the transition of resin composites from the pre-gel to post-gel stage. Within the post-gel stage resin composites experience and increase in the elastic modulus, but are still contracting due to further polymerization. Such polymerization shrinkage with an increased elastic modulus and minimal monomer slippage causes stresses to develop. The most rigid material (the material having the highest elastic modulus) will result in the highest stress. C-factor and substrate compliance Since composite shrinkage strain is limited by its bonding to the cavity walls, two factors must be considered. The first factor is the amount of confinement exploited on the material, which is calculated by the percentage of composite surface bonded to the substrate in relation to the total surface area. This ratio is known as the “cavity configuration factor” (C-factor). It is defined as the ratio between bonded and unbonded surfaces of the composite sample. The second factor is the bonding substrate’s compliance. The substrate’s compliance refers to the degree of stiffness of the substrate walls. Using a near-zero compliance system during testing and higher C-factors are usually related to higher stress values. Water sorption Water sorption of resin composites and their resultant expansion might compensate for the resin composite shrinkage. However relief caused by water sorption expansion occurs slowly in comparison to the rapid polymerization shrinkage stresses. Although water sorption lessens polymerization shrinkage stresses, it causes a series of negative consequences such as deterioration of mechanical properties and color stability of resin composites. 7.5.2.3 Management of polymerization shrinkage stress Light-curing modes Initial light exposure at lower irradiance values was found to reduce the reaction rate and stress development. This is due to the extended pre-gel stage that increases the chance for resin flow before the vitrification stage.
Resin-based dental composites for tooth filling 143 The “soft-start” protocol involves an initial light exposure with lower irradiance for a certain period of time, followed by full irradiance. Another protocol involves a few seconds initial exposure with reduced light irradiance followed a waiting period without irradiance then full light irradiance later. Such protocol is known as “pulse-delay.” Such alternative light-curing protocols may not profoundly influence the shrinkage stresses and the final properties of the hardened composite resin. This may be due to that: (1) most shrinkage stress is developed during the post-gel stage; (2) soft-start irradiation procedures may result in reduced DC levels, which in turn lead to decreased cross-link density, thus adversely affecting the properties of resin composites in terms of solubility and elastic modulus (Santini, 2010). Incremental layering technique Incremental layering of resin composites has been highly recommended as it reduces the effects of shrinkage stresses caused by high C-factor, thus enhancing the bond strength and reducing interfacial microleakage, cuspal deflection, and postoperative pain. Despite the importance of incremental placement of light-cured resin composites to achieve adequate light penetration and maximum depth of cure, it may lead to void entrapment between layers, as well as increased chair time. Low elastic modulus liners Flowable composites have an elastic modulus 20%30% less than conventional hybrid composites. According to the “elastic cavity wall concept,” the shrinkage stress generated by a higher modulus resin composite can be absorbed by an elastic intermediate layer of flowable composite. This is claimed to help in lessening the stress at the toothrestoration interface displayed clinically as a reduction in cuspal deflection. Lower shrinkage stress monomer chemistry The evolution of low-volumetric shrinkage monomers such as Silorane or DX-511 has not necessarily led to a similar decrease in shrinkage stresses. This has provoked the development of “lower contraction stress” monomers such as SureFil SDR flow and another novel low-shrinkage stress monomer TCD-DI-HEA (bis (acryloyloxymethyl)tricyclo[5.2.1.02,6]decane), which is found, for example, in Venus Diamond Nanohybrid Composite (Heraeus Kulzer). This will be discussed later in detail (Section 7.6.1.1). Preheating Preheating of resin composites has been presented as a technique used to increase composite flow, marginal adaptation, and monomer conversion. Higher resin composite temperature reduces system viscosity and enhances radical mobility, resulting in additional polymerization and higher conversion before the vitrification point, decreasing the magnitude of stress. This technique will be discussed in detail in Section 7.6.2.2.
144 Advanced Dental Biomaterials 7.5.3 Physical properties 7.5.3.1 Thermal properties In general, the linear coefficient of the thermal expansion (α) values for composites are less than the mean of its components (fillers and matrix) summed together; however, the values exceed those for dentin and enamel. The α of fine composites is less than that of microfilled composites. Higher values for microfilled composites are mainly due to the greater amount of polymer matrices. Thermal stresses place an extra strain on the interfacial bonding between the restoration and tooth structure, which adds to the damaging effect of the polymerization shrinkage. During cyclic application of either hot or cold stimuli, the high difference between the thermal coefficient of expansion of composites and teeth may cause percolation of oral fluids and early bond failure. The thermal conductivity of fine composites is greater than that of microfilled composites because of the higher filler loading of fine composites. Inorganic fillers have higher thermal conductivity compared to the polymer matrix. 7.5.3.2 Solubility Inadequate light-curing intensity and duration may lead to insufficient polymerization, especially within the deeper increments. Improperly polymerized composites have greater water sorption and solubility, displayed clinically as early discoloration. Water storage of resin composites may lead to the breakdown of interfacial bonding between the fillers and resin matrix accompanied by the leaching of inorganic ions. During the first 30 days of storage in water, silicon leaches in the highest quantity (1517 μg/mL) but decreases later with time. Barium, boron, and strontium, which are present in glass fillers, also leach out to various degrees. Wear and abrasion of composites are mainly caused by the breakdown of the fillers/ matrix bonding (Braden, 1974). 7.5.4 Esthetic properties Esthetic restoration of anterior teeth is quite challenging as it mandates dealing with optically complex natural tooth structure with variation in optical behavior among enamel, dentine, and pulp. Enamel is translucent with approximately 70% light transmission, whereas dentine is opaque with 52% light transmission (Franco et al., 2007). An impeccable esthetic outcome with resin composite depends on different factors related to both technique and materials. Regarding techniques, esthetic restoration is influenced by shade mapping, meticulous isolation, design of teeth preparation, proper bonding protocol, as well as anatomical shaping and contouring to mimic natural teeth. The execution of perfect esthetic restoration is dependent on various materialrelated factors, such as available shades, fluorescence, translucency, chameleon effect, polishability, wear resistance, and gloss retention (Nahsan et al., 2012). Resin composite translucency and color are related to compositional structure and formulation, including pigments and opacifiers, filler size and loading, the
Resin-based dental composites for tooth filling 145 mismatch in refractive index between fillers and matrix, and photoinitiator system, as well as resin type and its curing quality (Mikhail and Johnston, 2014). Early brands of resin composite offered only single body translucency with different Vita shades, whereas later on manufacturers produced composite kits with variable translucencies (enamel, body, dentine) in addition to characterization stains and special effects (transparent incisal, opalescent halo, white spots). It should be usually considered that understanding the optical properties is more important than having several commercial kits in order to select the appropriate composite resin (Nahsan et al., 2012). Surface smoothness and high gloss improve the natural appearance of the final restoration. Surface smoothness is influenced by the filler size and finishing and polishing protocol. It is believed that resin with smaller fillers (microfilled and nanofilled) produces a smoother surface and greater luster compared to large particles (microhybrid). Recently, systematic review of influence of filler size on the surface finish of composite revealed that there is no in vitro evidence for superiority of submicron or nanofilled composites over microhybrid with regard to surface smoothness and gloss retention (Kaizer et al., 2014). Regarding the esthetics of a posterior composite, it is mainly dependent on proper bonding and adaptation with the cavity margin with proper anatomical reconstruction. 7.5.5 Mechanical properties Advancements in dental resin composites over the years have focused on the enhancement of the mechanical properties. Generally, current dental resin composites show satisfactory mechanical behavior to be used in most of the clinical restorative procedures. Concern regarding placement of resin composite in high stress situations still exists. This concern is justified as one of the main reasons for composite failures is fracture. Composite mechanical properties rely mainly on the filler loading (percentage, distribution, size and type of fillers) and bonding between fillers and matrix. Generally, the higher the filler loading the higher the strength, stiffness, and toughness. There is considerable data in literature showing great variability in mechanical properties of dental resin composite commercial products even within the same category (Ferracane, 2011; Ilie and Hickel, 2009). So it is difficult to predict mechanical performance of a dental composite from its category alone. But generally the in vitro mechanical evaluation is essential as a primary screening and prediction of clinical performance and durability. Guidelines for mechanical evaluation and its correlation to clinical outcomes will be discussed later in this chapter (Table 7.2). 7.5.6 Biocompatibility Biocompatibility considerations for resin composite include cytotoxicity and allergic reactions for both patient and dentist. Several factors affect the biocompatibility of composite restoration such as composition (type of resin, fillers,
146 Advanced Dental Biomaterials Table 7.2 List of mechanical properties for various resin composite categories. Material type Flexural strength (MPa) Flexural modulus (GPa) Diametral tensile strength (MPa) Compressive strength (MPa) Hybrid composites Packable composites Ormocer-based composites Nanohybrid composites Compomers Flowable compomers Flowable composites Microfilled composites 116.6 ( 6 23.9) 7.3 ( 6 2.3) 32.5 ( 6 8.0) 79.0 ( 6 3.2) 105.9 ( 6 28.1) 8.4 ( 6 2.5) 34.3 ( 6 8.8) 80.3 ( 6 3.5) 104.3 ( 6 21.6) 7.5 ( 6 2.4) 35.2 ( 6 9.5) 77.0 (0) 103.1 ( 6 19) 5.0 ( 6 0.8) 40.5 ( 6 8.3) 78.4 ( 6 2.5) 94.7 ( 6 34) 89.1 ( 6 25.2) 9.2 ( 6 1.9) 4.2 ( 6 1.1) 33.9 ( 6 9.4) 33.5 ( 6 8.3) 74.9 ( 6 5.6) 64.4 ( 6 4.1) 99.8 ( 6 27.4) 4.4 ( 6 1.6) 38.3 ( 6 10) 66.1 ( 6 4.3) 73.5 ( 6 10.9) 3.8 ( 6 0.8) 24.2 ( 6 3.8) 60.7 ( 6 4.0) Source: Adapted from Ilie, N., Hickel, R., 2009. Investigations on mechanical behaviour of dental composites. Clin. Oral Investig. 13, 427438. bonding between matrix and fillers, and initiator system), method and quality of polymerization, as well as proper manipulation (cavity management, proper restorative technique, finishing and polishing). As biocompatibility concerns regarding composite are highly dependent on how much the composite leaches, so the compositional factor and high quality polymerization can be considered the key factors for reducing any harmful effects from composite restorations (Mallineni et al., 2013). Laboratory evaluation of freshly set methacrylate resins shows moderate cytotoxicity on directly exposed cultured cells over 2472 hours, however recent types of resins cause minimal cytotoxicity with higher quality polymerization and less leachants. The cytotoxicity is significantly reduced within 2448 hours after polymerization and significantly diminished by the presence of a dentine barrier. Generally, properly light-cured resins are less cytotoxic compared to chemically cured resins. DC of light-cured resins was found to range from 55%75% and can be increased up to 80% with further postcuring (Mallineni et al., 2013). In vivo usage tests revealed low to moderate pulpal inflammation after 3 days with about 0.5 remaining dentine thickness. Within 58 weeks, the reaction diminishes and reparative dentine formation was observed. Pulpal inflammatory response was minimal with use of protective liner or stable bonding agent. Polished composites show significantly less cytotoxic effect in laboratory testing. Also, the extent of the restoration especially its surface area is considered one of the factors that can affect the leaching procedure clinically (Mallineni et al., 2013).
Resin-based dental composites for tooth filling 147 In addition to short-term leaching, resin composite is susceptible to the combined effect of mechanical, hydrolytic, and enzymatic degradation leading to long-term elution of degradation products with varying extent. This long-term elution is considered of more biological significance, as it theoretically occurs throughout the service life of the restorations (Bakopoulou et al., 2009). Controversy about estrogen-like response (xenoestrogenic effect) of residual BisGMA resin has been raised. It was found that Bis-GMA can cause xenoestrogenic effect in vitro. However, there is no clinical evidence for an estrogen-like effect of any current commercial products (Mallineni et al., 2013). 7.5.7 Degradation As mentioned previously, degradation over time can caused elution of different products from resin composite. Resin degradation may be caused by different stimuli, such as thermal, photocatalytic, and/or chemical effects. For example, it has been found that biologically derived enzymes can degrade the monomer components of composite resins, which may then result in the liberation of methacrylic substances. Various environmental challenges face composite restorations during service causing significant changes due to aging and degradation. Aging-derived changes can be due to chemical breakdown by hydrolysis and enzymatic degradation by cholesterol esterase and pseudo-cholinesterase. Degradation also can be influenced by the applied stresses associated with swelling of resin leading to formation of voids and microcracks and premature failure (Drummond, 2008). The nature and actual extent of these changes are not clear. However, degradation effects should be considered during materials evaluation and characterization as it is not appropriate to evaluate only the initial properties. Current evidence supports the leading roles of the organic phase and the matrix/filler interface in the degradation of strength and toughness (Lohbauer et al., 2013). Accordingly, recent research directions have focused on development of degradation-resistant restorative materials. The recent modification and enhancement trials to develop fatigue resistant materials will be discussed later in this chapter. 7.5.8 Clinical durability Success and longevity of dental restorations are dependent factors related to the material, patient (oral hygiene, masticatory forces, and abnormal habits), dentist (skills, isolation, techniques), and cavity size and complexity. Since the introduction of resin composite, the improvement of its clinical performance and durability has been highly challenging, especially in stress bearing posterior restorations. Recently, the marked improvement in composite resin materials have significantly enhanced its clinical performance in different clinical situations. A systematic review of different clinical trials over the last two decades has revealed overall satisfactory success rates of posterior composite restorations. The overall success rate between 1995 and 2005 was 89.41%,while for 200616 it was
148 Advanced Dental Biomaterials 86.87% for all studied restorations. The reported overall success rates for long-term studies improved in the 200616 decade ranging from 64% to 96.9% compared with the 50% to 83% for the 19952005 decade. The annual failure rate of posterior composite ranges from 1.5% to 2.4%. This trend indicates that the clinical performance and durability of posterior composite have improved (Alvanforoush et al., 2017). The main reasons for failure in the first 4 years of a composite restoration’s service life were fracture, marginal defects, and secondary caries. In contrast, for restorations older than 5 years, secondary caries and fracture were the predominant failures. There is an important shift in failure reasons from recurrent caries, wear, and postoperative sensitivity toward mechanical failure and endodontic treatment. This shift may be related to the increased frequency of placing composite restoration in more complex cavities (Alvanforoush et al., 2017). According to Cochrane clinical reviews, there is low-quality evidence to suggest that amalgam is more durable than resin composite restorations. There is insufficient evidence to support or refute any adverse effects amalgam or composite restorations may have on patients (Rasines Alcaraz et al., 2014). Regarding anterior composite restoration, there is proven good clinical performance with long-term follow-up periods with annual failure rates ranging from 0% to 4.1%. Failure behavior in anterior composite restorations differs from posterior teeth, with more restoration replacement due to esthetic reasons and less secondary caries. However there is still the need for long-term well-designed clinical trials to evaluate the durability and clinical performance of resin composite restorations, especially with the rapid development of investigation techniques and the improvement of restorative and adhesive technologies (Demarco et al., 2015). 7.6 Attempts for resin composite improvement 7.6.1 Regarding material formulation 7.6.1.1 Low-shrinkage composites Different strategies have been evolved to reduce polymerization shrinkage of resin composites which include: G G Reduction of resin matrix through addition of more fillers Nanofilled, hybrid filler loading, prepolymerized fillers This approach is discussed in detail earlier in Section 7.2.2 Synthesis of low-shrinking resins (Fig. 7.2) Ring-opening polymerization (Silorane, 3M ESPE) High molecular-weight resin (DX-511, Kalore, GC) Low-shrinkage TCD resin (Venus diamond, HeraeusKulzer) Dimer acidbased dimethacrylate resin (N’Durance, Septodont). Ring-opening epoxy siloxane (Silorane) In 2007 an interesting ring-opening monomer was introduced called Silorane (3M ESPE), which is named for the two chemical subunits that make up the
Resin-based dental composites for tooth filling 149 molecule: siloxanes and oxiranes. Siloxanes provide an inorganic backbone of alternating silicon and oxygen atoms (SiOSiO) that has either an hydrogen or an organic side group attached to the Si atoms. Siloxanes can be branched, unbranched, or cyclic and are best known for imparting hydrophobic properties. In Silorane, the siloxane backbone is cyclic and the organic side groups are oxiranes. Oxiranes are three-member rings consisting of two carbons and one oxygen, which is the simplest epoxide, and are members of the epoxy resin group. Polymers formed from oxiranes are known for low-shrinkage and excellent physical properties (Pitel, 2013). Silorane polymerizes via cationic ring-opening of the oxirane groups, leading to very minimal overall shrinkage. According to the manufacturer, the volumetric shrinkage obtainable with the silorane-based composite is approximately 0.9% when evaluated by the bonded disk method, also known as the Watt’s method. This was significant because it was the first time a commercially available direct composite has reported a shrinkage factor ,1. However, when using other generally accepted test methods, the actual polymerization shrinkage stresses may not be significantly lower than that possible with more conventional resin-based composites (Pitel, 2013). Although it was a fascinating idea to have a low-shrinkage Silorane, it has several serious limitations. It was not compatible with other resins, it was extremely hydrophobic, and it required special proprietary bonding agents. In addition, the claimed lower shrinkage did not lead to either reduced stresses in the cavity or clinically significant superiority over conventional resins. High molecular-weight resin (DX-511, Kalore, GC) Limitations of the Silorane monomer have led to research focusing on developing another strategy. Since volumetric shrinkage has been linked to molecular sites where there is a conversion of an electrostatic van der Waals interaction to a covalent bond, it stands to reason that reducing the total number of such conversions should effectively reduce the total volumetric shrinkage. This can be accomplished by utilizing monomer molecules with higher molecular weights, which also typically occupy a larger volume of space. With fewer total monomer molecules to occupy the same potential space compared to smaller monomers, there is a reduction in the number of covalent bonds formed per unit space. This will ultimately reduce the total polymerization shrinkage and shrinkage stress. An example of this strategy is found in Kalore (GC America, Inc.), which was introduced in 2009. Kalore is 82% filled by weight and has a reported volumetric shrinkage of approximately 1.7% (by mercury dilatometer). It replaces the BisGMA monomer with DX-511, a new lower shrinking monomer chemically related to UDMA that has been developed by DuPont. Thus Kalore is able to retain full methacrylate compatibility. DX-511 has a much higher molecular weight (895 g/mol) than Bis-GMA (513 g/mol) and has both stiff and flexible segments in the monomer, the latter of which reduces the monomer viscosity, permitting lower amounts of diluent monomers to be used in the formulation (Pitel, 2013).
150 Advanced Dental Biomaterials Dimer acidbased dimethacrylate monomer (N’Durance, Septodont) Septodont introduced phase separating low-shrinkage dimerbased dimethacrylate resin in N’Durance composite. The core structure of the dimer acidbased monomer is composed of both linear and cyclic aliphatic structures (Fig. 7.2). “Dimer acid” means any of the class of cycloaliphatic carboxylic acids that are highmolecular-weight dibasic acids which are viscous liquid and can be polymerized directly with alcohols and polyols to form polyesters. The manufacturer claims both low-volumetric shrinkage, and high DC. The high conversion degree limits the elution of any residual monomer and thus enhances the biocompatibility of the material. The low shrinkage combined with high DC can be explained by the high molecular weight with bulky nature of the core structure of the dimer acidbased dimethacrylates. Also, it is claimed that phase separation occurs during polymerization leading to reduced shrinkage stresses. Low-shrinkage TCD urethanebased monomer (Venus Diamond, Heraeus Kulzer) Another novel low-shrinkage/stress monomer is TCD urethanebased resin, patented by Heraeus Kulzer in 2008 and introduced as Venus Diamond resin composite. The TCD monomer applies several strategies to reduce volumetric shrinkage and shrinkage stress. First, it has a backbone similar to Bis-GMA, which helps reduce molecular vibrations during polymerization. In addition, it lacks the polar hydroxyl groups found in Bis-GMA. Both of these properties help reduce the intermolecular distances between monomer molecules in the uncured state. Therefore the change between the uncured and cured states is minimized, which would help impart lower volumetric shrinkage. TCD resin also has an elastic urethane side chain allowing stress absorption and compensation of the volumetric shrinkage. 7.6.1.2 Smart composites Smart materials have been promoted widely as a key technology that will underpin all manner of novel products with unique capabilities. Indeed, in 2013, the World Economic Forum’s Global Agenda Council on Emerging technologies identified a class of smart materials, as one of the top 10 most promising technological trends. Smart materials can respond by changing some properties with external stimuli such as stress, pH, and temperature. As secondary caries is considered one of the most common causes of resin composite failure, recent conservative emerging trends focus on the development of smart bioactive remineralizing antibacterial resin composites. The therapeutic conservative treatment aims to improve resin durability with more tooth structure conservation. Different approaches to develop such material will be reviewed and clarified (Maas et al., 2017) (Fig. 7.8). Bioactive remineralizing composites Composites containing calcium phosphate (CaP) particles are considered “smart materials” as its ion release increases with more acidic conditions. Different CaP phases have been inserted in dental composite as bioactive fillers such as tetracalcium phosphate, amorphous calcium phosphate dicalcium phosphate
Resin-based dental composites for tooth filling 151 Figure 7.8 Diagram representing different types of smart resin composite. anhydrous, and dicalcium phosphate dehydrate. It is hypothesized that the released calcium and phosphate ions will supersaturate the surrounding environment favoring deposition and formation of hydroxyapatite. There are several factors that influence the ion release pattern in resin composite. The higher the surface area, volume fraction, and resin matrix hydrophilicity, the higher the ion release. Unfortunately, the increased amount of hydrophilic CaP fillers may lead to significant weakening of the resin composite material. This weakening can be attributed to the lack of strong bonding between bioactive fillers and resin matrix. This could be improved with surface functionalization using a silane coupling agent but it reduced the ion release and remineralization tendency due to silane hydrophobicity (Maas et al., 2017). Ion release is a multifactorial process and does not occur indefinitely, so all factors should be finely tuned for tailoring a potent long-lasting remineralizing composite without deterioration of physical and mechanical properties.
152 Advanced Dental Biomaterials Table 7.3 Different antimicrobial agents added to resin composite. Filler modification Nonreleased antibacterial agents G G Released antibacterial agents G G G G Silver-supported fillers MDPB SrF2 Silver ions Silversilica glass ZnO Matrix modifications G G G G G G G G G Triclosan QPEI Acrylic-amine-HF salts Methacryloyl acidfluoride Acrylic-amine-BF3 Chlorhexidine Benzalkonium chloride Cetylpyridinium chloride Chitosan HF, Hydrogen fluoride; MDPB, 12-Methacryloyloxydodecylpyridinium bromide; QPEI, quaternary ammonium polyethyleneimine; SrF2, strontium fluoride; ZnO, zinc oxide. Antibacterial composites Resin composites have more tendency for plaque accumulation compared to other restorative materials and intact tooth structure. This bacterial affinity is believed to increase frequency of secondary caries around composite restorations (Ferracane, 2017). Antibacterial composite formulations rely on modification of either matrix and/ or fillers with an antibacterial component. The antimicrobial strategies can be divided into stationary nonreleased additives or soluble released agent (Beyth et al., 2014) (Table 7.3). Fluoride anticariogenic potential through its antibacterial effect and formation of acid-resistant fluoroapatite is well documented. Accordingly, there have been different attempts to add these anticariogenic qualities to resin composite. Fluoride can be added in the form of fluoridated silicate glass particles, such as barium-aluminofluoride glass, or addition of inorganic fluoride compounds, such as ytterbium trifluoride (YbF3), sodium fluoride (NaF), and strontium fluoride (SrF2), to the filler system. Another approach to add a fluoride-releasing property to a composite is incorporation of prereacted glass ionomer particles, this led to the production of a commercial giomer called Beautifil from Shofu (Nedeljkovic et al., 2015). Fluoride release from fillers occurs through water diffusion into the resin matrix and exchange reaction. It is a challenging problem to achieve adequate fluoride release over time without compromising the physical properties of composite. Different promising approaches have been tried to overcome this challenge through incorporation of calcium fluoride nanoparticles to filler system. Also, using organic fluoride compounds, such as methacryloyl fluoride methyl acrylate and acrylic the amine-HF salt in the resin matrix which have revealed promising results. However, the leached fluoride form composite is still much lower and less effective compared to resin-modified glass ionomer.
Resin-based dental composites for tooth filling 153 Silver is another well-known antimicrobial agent that has been used either as a releasing or nonreleasing additive in resin composite. Silver can be added in different forms such as a glass filler releasing silver ions, silver nanoparticles, silver-supported apatite, and silver-supported zeolite. Two other silver carriers incorporated in composites, silver-supported silica gel and silver-supported zirconium phosphate, showed potent contact antibacterial effect without silver ion release. This showed potent in vitro antimicrobial effect without compromising mechanical properties. Zinc oxide nanoparticles have been added to composite, but they exhibited only a short-term contact inhibition of bacterial growth, without the release of zinc ions. Another simple approach to develop antimicrobial composites is modification of the resin matrix with a soluble disinfectant, such as triclosan, chlorhexidine, quaternary ammonium compounds (benzalconium chloride and cetylpyridinium chloride), biopolymer chitosan, antibiotics (vancomycin and metronidazole), and green tea extract (epigallocatechin-3-gallate). Unfortunately, all soluble antibacterial additives in composites showed “burst effect,” which means an initial large amount of leaching for a few days followed by a significant decrease in the remaining antimicrobial. Moreover, the antimicrobial release led to a porous structure with significant compromise of the material’s mechanical properties. To overcome these drawbacks, the synthesis and copolymerization of an antibacterial resin with the methacrylate resin was extensively investigated. This would exhibit a long-term contact antibacterial effect without leaching or compromising the mechanical properties. Methacryloyloxydodecyl pyridinium bromide is the most extensively studied antibacterial resin with highly promising results. It was incorporated as a prepolymerized resin filler. Quaternary ammonium polyethyleneimine (QPEI) is another antibacterial monomer that has been incorporated as nanoparticles to provide a large surface area for a potent antibacterial effect. QPEI includes for example, quaternary ammonium dimethacrylate and methacryloxylethyl-cetyldimethyl ammonium chloride (DMAE-CB). Macromolecular resins also have the advantage of being chemically stable, nonvolatile, and sustaining long-term antibacterial potential (Cheng et al., 2017). To develop solid evidence regarding new generations of therapeutic, bioactive, and antimicrobial composite resins, more clinical studies are highly needed as most of the current studies are in vitro. Nonetheless, the multidisciplinary research concerning these material categories is expected to offer tremendous benefits to oral and dental health. Polyacid modified composite (compomer) In the early 1990s, compomer was introduced as a new material class trying to combine the esthetics of resin composite with adhesion and fluoride release of glass ionomer. It is composed of methacrylate resin (UDMA) as a traditional composite modified with polymerizable dicarboxylic acid. The fluoroaluminum silicate glass of glass ionomer is added in the form of glass filler embedded in the resin matrix. There is not sufficient water for a glass ionomer acidbase reaction to occur and compomer setting is dependent only on free radical polymerization as per
154 Advanced Dental Biomaterials conventional composites. Surface pretreatment of tooth structure with adhesive resins is essential for bonding as the compomer was not able to bond chemically with the tooth structure as a glass ionomer. Fluoride release and recharge of compomer was found to be much lower than glass ionomer cement. This may be attributed to the hydrophobic nature of the resin matrix compared to hydrophilic glass ionomer. The clinical use of compomer is limited to pediatric restoration and low stress bearing areas in adults as its mechanical performance is much lower than conventional resin composites (Zimmerli et al., 2010). Self-healing composites Self-healing materials are another example of smart material that responds to mechanical stimulus and extends its survival by restoring mechanical integrity after damage. Self-healing composite relies on an encapsulated polymerizable solution with reactants which are released when the capsule is ruptured by crack propagation. Then et al. (2011) described the first experimental self-healing dental resin composite using the microcapsule technique. They made it using a synthesized dicyclopentadiene (DCPD) capsule filled with urea formaldehyde (UF) dispersed in Bis-GMA/UDMA/TEGDMA resin with 55 wt.% of silanated silica fillers and 2 wt. % of catalyst. It should be noted that the presence of the capsules within the resin may provide some benefit in terms of mechanical properties, regardless of the selfhealing effect. It revealed a significant recovery of fracture toughness without compromising biocompatibility. Concerns regarding the biological safety of DCPD led to the development of different experimental self-healing systems in dental composite. Microcapsules of poly urea formaldehyde filled with TEGDMA resin and TA activator were used. This self-healing system utilizes BPO in resin matrix to initiate polymerization of the healing resin (Fugolin and Pfeifer, 2017). Several recent studies using microcapsule self-healing mechanisms with various healing resins revealed a minimum of 65% recovery of the fracture toughness with minimal cytotoxicity (Fugolin and Pfeifer, 2017; Maas et al., 2017). 7.6.2 Regarding manipulation 7.6.2.1 Layering technique Traditionally, direct composite restoration is recommended to be laid in 2 mm increments for proper polymerization and reduction of shrinkage stresses. Incremental application of composite can be placed in either horizontal or oblique successive layers. Oblique increments are placed in such a way to try to avoid bonding of opposite walls. This technique is claimed to result in less cuspal deflection and reduced shrinkage stresses as a result of the controlled C-factor (ratio of bonded to unbonded restoration surfaces) by bonding to single cusp with each increment application. However, the subsequent increments will eventually bond to opposite walls producing significant stresses. Also the C-factor concept is too simple to explain shrinkage stresses which are a multifactorial complex phenomenon.
Resin-based dental composites for tooth filling 155 Finite element stress analysis compared oblique and horizontal incremental application of a premolar restoration and revealed that oblique increments produced less stresses along the compositeenamel interface, and higher stresses along the compositedentin interface. Two millimeters increment seems to produce adequate mechanical properties with low-shrinkage stresses. It is extremely important to note that an increased number of increments give more chance of voids and porosity. Although there is no conclusive evidence in the literature concerning the incremental layering advantages regarding polymerization shrinkage, it is recommended as the standard technique for composite placement, claiming dissipation of shrinkage stresses and with proper polymerization. 7.6.2.2 Composite preheating Composite preheating before insertion in the cavity offers several advantages. It reduces composite viscosity and allows injection molding of highly filled viscous composite into the cavity with better adaptation, polymerization quality, and fewer voids. This advantage allows handling of the composite like a flowable restorative material while maintaining the superior properties of packable composites. The warm composite shows enhanced molecular mobility leading to improved mechanical properties after curing. There was some concern regarding injecting warm composite inside the cavity with fears of pulp thermal injury. However, composite cools immediately after placement and results in only slight temperature elevation compared to intraoral temperature. Although composite preheating requires additional equipment (Fig. 7.9) and some training for mastering placement technique, the cost is relatively low and the learning curve is shallow with high benefits and very limited adverse events (da Costa et al., 2011). The most popular heating devices present on the market are Calset and Thermaflo. Some simpler methods have been used for composite warming instead of these devices such as hot air blower, hot water bath, and microwave oven. Modified glass bead sterilizer and wax warmer are also promising suggested innovations especially where economic condition and availability are the major concerns to implement these techniques. 7.6.2.3 Use of lining material As mentioned previously, polymerization shrinkage stresses are multifactorial and elastic modulus of the composite material is one of the factors that can affect it. It is believed that low elastic modulus materials generate less shrinkage stresses compared to rigid materials. Based on this assumption, the concept of “elastic cavity wall” and “sandwich technique” have been evolved by placing less rigid material as a cavity liner under composite restoration. Several studies have investigated different resilient lining materials such as thick adhesive layer, glass ionomer, resinmodified glass ionomer, or flowable composite. This approach is claimed to reduce the polymerization shrinkage stresses with its deleterious consequences and increase the durability and success rate of composite restoration (Soares et al., 2017).
156 Advanced Dental Biomaterials Figure 7.9 Calset composite heating device. Nevertheless, there is inconsistent evidence concerning the claimed advantages of the use of lining material under composite. There is no evidence of enhanced performance or durability when comparing restorations with and without elastic cavity liners. Accordingly, there is no need to add cost and/or unnecessary clinical step by placing lining material underneath composite restorations (Schenkel et al., 2016). 7.6.3 Regarding both material formulation and manipulation Much effort has been concentrated on developing materials that are simpler to use by virtue of requiring fewer application steps, such as bulk-fill and self-adhesive composites. 7.6.3.1 Self-adhesive composites Self-adhesive flowable composites were introduced in 2009. Currently, there are three commercially available products (Table 7.4). These materials are currently recommended for liners and small restorations but with simplified application without separate bonding protocol pretreatment. Self-adhesive composite technology is relevant to self-etch adhesive systems as it relies on the presence of acidic functional monomers such as trimelitic acid methacrylate (4-MET), GPDMs, or
Resin-based dental composites for tooth filling 157 Table 7.4 Different products, manufacturers, and compositions of the commercially available self-adhesive composites. Product Manufacturer Composition Fusio Liquid Dentin Pentron Clinical, Orange, CA, United States Vertise Flow Kerr, Orange, CA, United States Embrace Wetbond Pulpdent, Watertown, MA, United States Resin: UEDMA, TEGDMA, HEMA, 4MET, catalyst Fillers: SiO2 (65 wt.%), silanated barium glass, NaF Resin: GPDMA, HEMA, Bis-GMA, catalysts Fillers: Prepolymerized filler, silanated barium glass, nano-sized colloidal SiO2, YF3 (70 wt.%) Resin: UEDMA, BMEP, TMPTMA, HEMA, water, catalysts Fillers: SiO2, NaF (37 wt.%) 4-MET, Trimelitic acid methacrylate; Bis-GMA, bisphenol glycidil dimethacrylate; BMEP, bis-methacryloyloxy ethyl phosphate; GPDMA, glycerolphosphoric acid dimethacrylate; HEMA, hydroxyethyl methacrylate; TEGDMA, triethylene glycol dimethacrylate; TMPTMA, trimethyloyl propane trimethacrylate; UEDMA, aliphatic diurethane dimethacrylate. bis-methacryloyloxy ethyl phosphate (BMEP), in addition to hydroxyethyl methacrylate (HEMA). These monomers vary in acidity and are responsible for partially etching the tooth structure and penetrating through the smear layer, forming a hybrid layer. HEMA is added to increase the wettability and allow better hybridization with the moist tooth structure. These composite formulas were designed with the aim of getting the benefits of the advancing technology of micromechanical and chemical bonding of the recent simplified self-etch adhesive systems (Maas et al., 2017) (Fig. 7.10). Under the limitations of the present studies, there is overall agreement that selfadhesive flowable composite are less efficient than conventional restorative systems with few exceptions. This is attributed to the relative high viscosity limiting the diffusion into the tooth substrate in addition to the limited extent of decalcification. Some researchers found that application of phosphoric acid pre-etch and/or selfetch universal adhesives increased the bond strength significantly. Unfortunately, the different surface pretreatment deprived the self-adhesive composites of its main advantage. Accordingly, routine application of self-adhesive flowable composites should be carefully considered. 7.6.3.2 Bulk-fill composites Bulk-fill resin composites have been introduced to the market in both flowable and conventional/sculptable viscosities, with the premise of simplified application (Fig. 7.11). Different concerns were raised with this concept such as ensuring proper polymerization, depth of cure, adaptation of the material to cavity walls and margins, in
158 Advanced Dental Biomaterials Figure 7.10 Chemical structure of glycerol phosphate-dimethacrylate (GPDM) resin with acidic side chain for self-adhesive composite. Figure 7.11 Different packing techniques comparing conventional incremental packing to bulk-fill composite either single 4 mm increment or 4 mm base with capping composite. addition to polymerization shrinkage and shrinkage stresses. These concerns should be considered while tailoring the material composition and evaluating its laboratory and clinical performance. Regarding proper polymerization and depth of cure, it has been achieved through different approaches, which include modified initiators, translucent fillers, and modification of resin chemistry. Different modifications have been invented to reduce polymerization shrinkage and its stresses. Among these approaches are high molecular-weight resins, stress decreasing modulators, and addition fragmentation polymer, in addition to stress absorbing fillers. Flowable bulk-fill materials generally have lower filler loading than nonflowable, sculptable materials and require that the occlusal layer be filled with a “cap” of a more highly filled composite that is expected to be stronger and more wear resistant under occlusal loading (Chesterman et al., 2017). The mechanical properties, esthetic outcomes, and placement techniques vary significantly between the available bulk-fill composite materials. Bulk-fill materials are highly beneficial in the restoration of posterior teeth with difficult patient control and limited time available at present, for example, in children, anxious patients, and large posterior cavities. Overall there are early promising in vitro studies supporting the use of bulk-fill composites which confirm the manufacturer’s claims. However, there is very little clinical research on the long-term outcomes of these materials. Such evidence is
Resin-based dental composites for tooth filling 159 highly necessary before the real effectiveness of these materials can be confirmed. Until further data are available, it is recommended that clinicians carefully select materials and strictly follow the current recommendations in application (Table 7.5). Indirect composite There are three main restorative techniques available, namely direct, indirect, and semidirect. Indirect technique allows better control and easier restoration of occlusal anatomy and proximal contours. In addition, it is expected to provide better margins due to the limitation of the polymerization shrinkage stresses into the confined thin layer of the resin cement. Moreover, the highly controlled polymerization and sometimes combination of different polymerization protocols produces restorations with better polymerization quality and hence improved physical and mechanical properties. However, conventional indirect technique requires two visits, additional clinical steps as recording impression, provisional restoration fabrication, and laboratory contribution, leading to increased time and cost (Alharbi et al., 2014). Indirect composites can be differentiated according to: 1. 2. 3. 4. manufacturing process (manual layering vs machinable CAD/CAM), microstructure (dispersed filler vs polymer infiltrated ceramic network “PICN”), polymerization (light, high temperature, or high temperature with high pressure), and composition (matrix and filler technology). Different types of indirect composites are summarized in Table 7.6. Manual indirect composites have similar structures, compositions, and layering techniques to the direct composite. The key difference is that it is built up extraorally on a model in the lab and subjected to additional highly controlled polymerization, improving the DC. Digital computerized techniques have recently revolutionized the world of dentistry, introducing digital designing and computerized milling of different industrial high-performance materials blocks. CAD/CAM industrially processed blocks offer homogenous materials with less flaws and extremely high filler loading up to levels that couldn’t be possible with direct composite. Recent CAD/CAM blocks are BisGMA-free, and they are no longer photopolymerized. They are processed using industrial grade high temperature and sometimes high pressure. CAD/CAM blocks’ microstructure can be either formed of a high amount of dispersed fillers or a PICN (Mainjot et al., 2016). PICN blocks are inspired by glass-infiltrated ceramics produced by Vita in the 1980s. In 1997, R. Giordano patented the idea of the infiltration of presintered glassceramic scaffold with resin to produce PICN. PICNs differ significantly from composite materials with dispersed fillers classically incorporated by mixing. In contrast to dispersed fillers, the ceramic network forms a three-dimensional scaffold of interconnected particles forming a real skeleton which helps in a more effective stress distribution in all directions, thus resisting mechanical breakdown. In 2012, Vita produced the first commercial PICN blocks called Vita Enamic. The high volume fraction of ceramic has led to introduction of the confusing
Table 7.5 Summary of commercially available bulk-fill resin composite with the manufacturers, claimed depth of cure, application technique, composition, and claimed technology. Product Manufacturer Type Available shades Claimed depth of cure (mm) Composition Claimed technology Surefil SDR Dentsply Flowable base Universal 4 Resin: Modified UDMA, TEGDMA, EBPADMA Fillers: 68 wt.%, 44 vol.%; barium and strontium fluoroaluminosilicate glass Initiator: CQ, BHT photoaccelerator Resin: Bis-EMA, UDMA, TEGDMA, TMPTMA Fillers: 86 wt.%, 66 vol.% silanated strontium aluminum sodium fluoride phosphate silicate glass Initiator: CQ Resin: Bis-GMA, Bis-EMA, UDMA Fillers: 81 wt.% 61 vol.% barium glass, ytterbium trifluoride, mixed oxide, prepolymer Initiators: CQ, TPO, Ivocerin Resin: Bis-GMA, Bis-EMA, UDMA Fillers: 68 wt.%,46 vol.% barium glass, ytterbium trifluoride, mixed oxide, prepolymer Initiator: CQ, TPO, Ivocerin G Quixfil Tetric EvoCeram Bulk-fill Tetric EvoFlow Bulk-fill Dentsply Ivoclar Vivadent Ivoclar Vivadent Viscous full depth filling Viscous full depth filling nanohybrid Flowable base Universal IVA IVB IVW IVA IVB IVW 4 4 4 G G G G G G Lower filler content Novel high molecular UDMA-based monomer (849 g/mol) with stress decreasing modulator Increased filler loading reducing shrinkage Increased translucency and amount of photoinitiator increasing the depth of cure Photoinitiator system containing Ivocerin makes it more efficient in promoting polymerization depth Prepolymerized fillers reducing shrinkage and relive stresses The main difference between these different products in filler loading which controls viscosity of the material either flowable or viscous
Tetric NCeram Bulk-fill Ivoclar Vivadent Viscous full depth filling nanohybrid IVA IVB IVW 4 Tetric N-Flow Bulk-fill Ivoclar Vivadent Flowable base IVA IVB IVW 4 SonicFill Kerr Sonic activated full depth filling A1, A2, A3, B1 5 Resin: Bis-GMA, Bis-EMA, UDMA Filler: 78 wt.%, barium glass, ytterbium trifluoride, mixed oxide, prepolymer Initiators: CQ, TPO, Ivocerin Resin: Bis-GMA, Bis-EMA, UDMA Filler: 78 wt.%,barium glass, ytterbium trifluoride, mixed oxide, prepolymer Initiators: CQ, TPO, Ivocerin Resin: Bis-GMA, Bis-EMA, TEGDMA Fillers: 83.5 wt.% 69 vol.% Silanated barium boron aluminum silicate glass, silica G G G Filtek Bulk-fill Filtek Bulk-fill flowable 3M ESPE 3M ESPE Viscous full depth filling Flowable base A1, A2, A3, B1, C3 U, A1, A2, A3 4 (Class I) 5(ClassII) 4 Resin: Bis-GMA, AUDMA, UDMA, and 1, 12dodecane-DMA Fillers: 76.5 wt.%, 58.4 vol.% Nano silica, nano zirconia, zirconia/silica nanocluster, ytterbium trifluoride Initiators: CQ, EDMAB Resin: Bis-GMA, Bis-EMA (6), procrylate, and UDMA Fillers: 64.5 wt.%, 42.5 vol.% zirconia/silica fillers, ytterbium trifluoride Initiators: CQ, EDMAB G G G G Sonic activation of the composite dramatically lowers the viscosity during placement High translucency and increased amount of photoinitiator for deeper polymerization High amount of filler loading High molecular-weight AUDMA reduce polymerization shrinkage Addition fragmentation resin reducing shrinkage stresses by photocleavage during polymerization Nanofiller technology reducing shrinkage Enhanced photoinitiators for deeper curing (Continued)
Table 7.5 (Continued) Product Manufacturer Type Available shades Claimed depth of cure (mm) Composition Claimed technology X-tra fil VOCO Viscous full depth filling Universal 4 Resin: Bis-GMA, UDMA, TEGDMA Fillers: 86 wt.%, 70 vol.% barium boron aluminum silicate glass Resin: Bis-GMA, UDMA, TEGDMA Fillers: 75 wt.%, barium boron aluminum silicate glass Resin: UDMA, Bis-EMA Fillers: 65 wt.% 38 vol% barium glass filler, ytterbium fluoride and silica Bis-GMA, TEGDMA, PMMA Hybrid filler, fractions, and Eglass fibers G X-tra base VOCO Flowable base Universal, A2 4 Venus Bulk-fill Heraeus Kulzer, Flowable base Universal 4 Gc Fiberreinforced bulk-fill base Everx posterior G High filler loading to reduce shrinkage High translucency for deeper light penetration G 4 The fibers may conduct and scatter the light over longer distances, explaining the relatively deep and wide polymerization observed BHT, Butylated hydroxyl toluene; Bis-EMA, ethoxylated bisphenol A dimethacrylate; Bis-GMA, bisphenol glycidil dimethacrylate; CQ, camphorquinone; TEGDMA, triethylene glycol dimethacrylate; TMPTMA, trimethyloyl propane trimethacrylate; TPO, diphenyl (2,4,6-trimethylbenzoyl) phosphine oxide; UDMA, urethane dimethacrylate.
Table 7.6 Different types of indirect composite materials, microstructure, manufacturers, polymerization modes, and compositions. Manufacturing process Manual layering Microstructure Dispersed fillers Polymerization mode Light Light 1 temperature complement Material Manufacturer Composition Matrix Fillers Ceramage and Ceramage up Gradia Shofu UDMA, HEMA Silica-based glass GC Corp UDMA, other DMA Signum Heraeus Kulzer Sinfony 3M ESPE UDMA, TEGDMA, other DMA UDMA, other DMA Solidex Shofu UDMA SR Nexeco Ivoclar Vivadent UDMA, other DMA Silica powder, silicate glass powder, prepolymerized filler (75 wt.%) Silica-based glass fillers (74 wt.%) Strontium aluminum borosilicate glass, silica, and 5% glass ionomer fillers 50 wt.% More than 53% ceramic fillers Silica (1050 nm), Zirconia, silica-based glass (7080 wt.%) VITAVM LC Estenia C&B VITA Zahnfabrick Kuraray SR adoro Ivoclar Vivadent Twiny Yamamoto, Precious Metal Co BPA, TEGDMA, other DMA Unknown DMA (1Bis-GMA in opaque paste) UDMA 1 other DMA UDMA 1 TEGDMA Silica-based glass, alumina (2 μm and 2 nm) (92 wt.%) Silica-based composite Silica (20100 nm), zirconia, alumina, silicaparticles (200600 nm), zirconia aluminasilicaclusters (16 μm) (Continued)
Table 7.6 (Continued) Manufacturing process Machinable blocks Microstructure Dispersed fillers Polymerization mode Light High temperature PICN High temperature and high pressure Material Manufacturer 3M ESPE Composition Matrix Fillers Silica (0.6 μm), zirconia (0.6 μm) (85 wt.%) Silica-based glass, silica (20 and 300 nm) (71 wt.%) Silica (20 nm) zirconia (411 nm), zirconiasilica clusters (0.610 μm) (79 wt.%) Silica-based glass, silica (61 wt.%) Glassceramic sintered network (86 wt.%) Paradigm MZ 100 block Cerasmart GC America Lava Ultimate 3M ESPE Bis-GMA, TEGDMA UDMA 1 other DMA UDMA Shofu block HC Shofu UDMA, TEGDMA VITA Enamic VITA Zahnfabrik UDMA, TEGDMA AUDMA, Aromatic Urethane dimethacrylate; BPA, Bisphenol A; Bis-GMA, Bisphenol glycidil dimethacrylate; DMA, Dimethacrylate; EBPADMA, ethoxylatedbisphenol Adimethacrylate; EDMAB, ethyl 4-dimethyl aminobenzoate; HEMA, hydroxyethyl methacrylate; PICN, polymer infiltrated ceramic network TEGDMA, triethylene glycol dimethacrylate; UDMA, urethane dimethacrylate.
Resin-based dental composites for tooth filling 165 commercial name “hybrid ceramic” to describe this material. “Hybrid ceramic” is a common confusing commercial name that does not conform to materials science and does not reflect the actual properties of the material. Different descriptive names are present for PICNs such as double-network materials, interpenetrating phase ceramicresin composites, or ceramic-based interpenetrating phase materials. Among CAD/CAM composites, PICN materials constitute an innovative and promising class of materials with benefits in term of high flexural strength, hardness value between enamel and dentin (thus reducing tooth wear), elastic modulus intermediate between enamel and dentin, and minimal toxicity and resin leaching. Recent advances with CAD/CAM composites are promising according to the current in vitro studies. Unfortunately, there is limited clinical evidence regarding this category of materials and consequently no evidence-based recommendations can be supported for clinical practice at the moment (Mainjot et al., 2016; Angeletaki et al., 2016). Regarding selection of the suitable restorative protocol in clinical practice according to the current evidence, there is no difference in terms of clinical durability between direct and indirect resin composite restorations in different clinical situations. Given the current level of evidence, it appears reasonable to suggest that the less costly and less time-consuming technique of placing direct resin restorations should be the preference when guiding patient decision-making (Farsai, 2017; Radford, 2017). 7.7 Guidelines and recommendations for future laboratory and clinical researches There are always confusing gaps and difficulty in correlating the standardized in vitro evaluations with clinical performance of the restorative materials. Different trials to make such correlations are often criticized for lack of evidence for significant correlation with clinical performance. Although most of the laboratory methods are standardized and validated, there is interinstitute variability and different values can be presented for the same product even with ISO testing. Clinical studies show greater variability and a lack of standardization with the risk of different types of bias. Reasons for this variation may be accounted for by factors related to the patient, materials, skills, treated teeth, procedure complexity, and/or type of outcome and its assessment method. Correlations between laboratory and clinical outcomes were found to be moderately positive with a few significant results. Restoration fracture was found to be mostly correlated to the material’s fracture toughness, while clinical wear was correlated with flexural strength. No clinical correlations were found for compressive strength and flexural modulus. Unfortunately, these conclusions should be carefully considered as they are based on a huge amount of variable studies with inadequate descriptions or studies with a high risk of bias (Heintze et al., 2017).
166 Advanced Dental Biomaterials For better standardized and comparable future outcomes, eminent dental researchers with the Academy of Dental Materials have set guidelines for resin composite laboratory testing and recommendations for future randomized clinical trials. 7.7.1 Guidelines for laboratory evaluation of resin composite (mechanical behavior and technique sensitivity) In 2017 the Academy of Dental Materials set an initiative to review and critically appraise the different laboratory testing methodologies to evaluate resin composite mechanical performance and technique sensitivity, in an attempt to develop guidelines for in vitro composite evaluation. Test methods have been ranked on the priority of the specific property being tested, as well as on the specific test methods for evaluating that property. Focus was placed on the methods that are considered to be of the highest priority in terms of being the most applicable, useful, supported by the literature, and clinical correlation. The properties for evaluating resin composites were ranked in the priority of measurement regarding mechanical performance and technique sensitivity and are listed in Table 7.7. These guidelines were meant to help the researchers in choosing the proper method to assess the key properties of dental resin composites with regard to their mechanical behavior and technique sensitivity (Ilie et al., 2017; Ferracane et al., 2017) (Table 7.8). 7.7.2 Recommendations for future clinical studies Clinical research in restorative dentistry is confounded by various problems of study designs, trials length, and collected information in addition to high cost. Opdam et al. reviewed and discussed the pros and cons of different study designs and outcomes for the evaluation of dental restorations and set recommendations for future clinical trials. They recommended that clinical trials should have control groups, include confounders, such as patient risk factors, in the data, and analysis with assessment of patient-centered outcomes with relevance to the profession and patients. For further development of restorative dentistry, there is plenty of room for different research questions to be studied with well-designed and executed trials. Studies comparing new materials to their predecessors are needed before launching new products but it is not of prime importance. Randomized clinical trials comparing different treatment protocols are much more valuable. For example, comparison of direct versus indirect restorations, tooth replacement versus tooth retention, restoration repair versus replacement, bridge versus implant, all need to be assessed. Comparing materials with each other should be done against the gold standard. For material testing trials, clarifying any conflict of interest or sponsorship is mandatory. The primary outcome should be valuable and of relevance to patients, dentists, and stakeholders. The use of globally agreed outcomes for future trials would be
Table 7.7 Summary of the available methods for mechanical evaluation of resin composite. Clinical issue/requirement Properties Property rank Method Test rank Fracture and deformation resistance Strength 1 Toughness 2 Fracture toughness 1 Edge strength— chipping Fatigue 2 3-Point bending (ISO 4049) 4-Point bending Biaxial flexural (ASTM F394-78) Compression (ASTM D695) Tensile (ASTM D638) Diametral tensile (ANSI/ADA Spec 27) Impact (ISO 179/1961) Transverse impact (DIN 53 453) Shear Shear punch (ASTM D732-46) Calculated from strength test Impact (ISO 179/1961) Transverse impact (DIN 53 453) Double torsion Indentation—Vickers Chevron notched specimens (ASTM E1304-97) Single-edge notched beam (ASTM E399-12) Compact tension (ASTM E399-12) CK10 instrument (with acoustic emission) 1 2 2 3 1 4 2 2 4 2 1 2 2 2 3 2 1 2 1 1 Elastic modulus 1 Fatigue strength—Staircase Fatigue resistance—uniaxial loading (ASTM E606/E606M-12) 3-Point bending (ISO 4049) 4-Point bending Biaxial flexural strength (ASTM F394-78) Compression (ASTM D695) Tensile Diametral tensile (ANSI/ADA Spec 27) 2 2 2 1 3 3 1 4 Indentation hardness 1 Indentation (ISO/FDIS 14577-1) Martens (universal) (E DIN 50359) Vickers (ISO 6507-1) Knoop (ISO 6507-1) Rockwell (ISO 2039-2) Brinell (ISO 6506-1:200) 1 1 3 3 3 3 (Continued)
Table 7.7 (Continued) Clinical issue/requirement Properties Property rank Method Test rank Wear resistance Wear—abrasion/three body 1 Wear—attrition/ contact/two body 1 Wear—toothbrush 3 OHSU abrasion (ISO TS No. 14569-2) Alabama generalized (ISO TS No. 14569-2) ACTA (ISO TS No. 14569-2) OHSU attrition (ISO TS No. 14569-2) Alabama localized (ISO TS No. 14569-2) Ivoclar-Willytec simulator Munich-Willytec simulator Toothbrush/toothpaste(ISO TS No. 14569-2) 1 2 3 2 3 2 2 3 Source: Adapted from Academy of Dental Materials’ guidance.
Table 7.8 Summary of available tests for evaluation of resin composite handling, polymerization, and dimensional stability. Clinical issue/requirement Properties Property rank Method Test rank Technique sensitivity: handling —placement Stickiness 2 Slump resistance 2 Viscosity 2 Porosity 1 Radiopacity Sensitivity to ambient light Degree of conversion 1 1 1 Polymerization kinetics—rate 1 Depth of cure 1 Unplugging force, unplugging work Measure length/area of withdrawn composite Extrude and cure Imprint and cure (SRI) Rheology (with Viscometer) Pressing under standard load and measure size Section—assess under magnification 3D microtomography Archimedes method, ASTM D3171-11 X-ray ISO 4049 (vs aluminum step wedge) and ISO13116 ISO 4049 (Xenon light box) FT-IR spectroscopy Near IR FT-Raman DSC NMR FT-IR, NIR, Raman DSC Shrinkage/dilatometry Optical—interferometry, fluorescent probes Elastic modulus—DMA, rheometer Acoustic DEA ISO 4049—scraping Microhardness vs depth Chemical vs depth (FT-IR, NIR, Raman) Penetrometer Solvent dissolution Transition in translucency 1 2 1 1 1 2 1 2 2 2 1 1 1 1 2 2 4 1 1 2 2 3 2 2 2 1 1 2 3 3 Polymerization (Continued)
Table 7.8 (Continued) Clinical issue/requirement Properties Property rank Method Test rank Dimensional stability Polymerization shrinkage 1 Polymerization shrinkage rate/ kinetics and gel point Polymerization stress 1 Bonded disk SSA Linometer Dilatometer (e.g., Hg) Pyknometer Archimedes method—ISO (17304) Accuvol camera imaging Strain gage Digital image correlation E.g., from bonded disk or SSA 1 1 2 2 2 1 2 2 2 1 1 Hygro (swelling) expansion 1 Via UTM Bioman method Tensilometer (cantilever beam) SSA Photoelastic method Thin ring Indentation crack analysis Laser scanning of disk—stored for 3 months or more in water or other solvent. Calipers or dial gage Measuring microscope Displacement measured over a temperature range 1 1 1 1 2 2 3 1 Thermal expansion 3 2 2 2 DEA, Dielectric analysis; DSC, Differential scanning calorimetry; FTIR, Fourier transform infrared; NIR, Fourier transform near-infrared spectroscopy; NMR, Nuclear magnetic resonance; SRI, slump resistance index; SSA, stressstrain analyzer; UTM, universal testing machine. Source: Adapted from Academy of Dental Materials’ guidance.
Resin-based dental composites for tooth filling 171 desirable also to reduce the risk of selective reporting. Different patient-related risks such as caries risk and bruxism should be recorded and considered in analysis. The time and causes for any interventions should also be recorded to allow differentiation between success, failure, and survival. Different trial designs should be considered when planning to compare restorative treatments and materials: while randomized clinical trials reduce selection bias, their observation times are usually short and sample sizes limited. Retrospective and prospective practice-based studies circumvent these limitations, but have considerable risk of indication bias and confounding which need to be dealt with appropriately. Perhaps the answer lies in coordinated, multicenter studies of different design, no one design being capable of yielding all the required outcomes. Whatever the way forward, funders and sponsors of clinical research in restorative dentistry should encourage investigators to come forward with new innovative approaches to answer priority questions (Opdam et al., 2017). References Alharbi, A., Rocca, G.T., Dietschi, D., Krejci, I., 2014. Semidirect composite onlay with cavity sealing: a review of clinical procedures. J. Esthet. Restor. Dent. 26, 97106. Alvanforoush, N., Palamara, J., Wong, R.H., Burrow, M.F., 2017. Comparison between published clinical success of direct resin composite restorations in vital posterior teeth in 19952005 and 20062016 periods. Aust. Dent. J. 62, 132145. Angeletaki, F., Gkogkos, A., Papazoglou, E., Kloukos, D., 2016. Direct versus indirect inlay/ onlay composite restorations in posterior teeth. A systematic review and meta-analysis. J. Dentistry 53, 1221. Bakopoulou, A., Papadopoulos, T., Garefis, P., 2009. Molecular toxicology of substances released from resin-based dental restorative materials. Int. J. Mol. Sci. 10, 38613899. Baroudi, K., Rodrigues, J.C., 2015. Flowable resin composites: a systematic review and clinical considerations. J. Clin. Diagn. Res. 9, Ze18Ze24. Beyth, N., Farah, S., Domb, A.J., Weiss, E.I., 2014. Antibacterial dental resin composites. React. Funct. Polym. 75, 8188. Braden, M., 1974. Selection and properties of some new dental materials. Dent. Update 1, 489501. Braga, R.R., Ballester, R.Y., Ferracane, J.L., 2005. Factors involved in the development of polymerization shrinkage stress in resin-composites: a systematic review. Dent. Mater. 21, 962970. Cheng, L., Zhang, K., Zhang, N., Melo, M.A.S., Weir, M.D., Zhou, X.D., et al., 2017. Developing a new generation of antimicrobial and bioactive dental resins. J. Dent. Res. 96, 855863. Chesterman, J., Jowett, A., Gallacher, A., Nixon, P., 2017. Bulk-fill resin-based composite restorative materials: a review. Br. Dent. J. 222, 337344. da Costa, J.B., Hilton, T.J., Swift Jr., E.J., 2011. Critical appraisal: preheating composites. J. Esthet. Restor. Dent. 23, 269275. Demarco, F.F., Collares, K., Coelho-De-Souza, F.H., Correa, M.B., Cenci, M.S., Moraes, R. R., et al., 2015. Anterior composite restorations: a systematic review on long-term survival and reasons for failure. Dent. Mater. 31, 12141224.
172 Advanced Dental Biomaterials Drummond, J.L., 2008. Degradation, fatigue, and failure of resin dental composite materials. J. Dent. Res. 87, 710719. Farsai, P.S., 2017. Longevity of direct and indirect resin composite restorations in permanent posterior teeth: a systematic review and meta-analysis. J. Evid.-Based Dent. Pract. 17, 110112. Ferracane, J.L., 2011. Resin composite—state of the art. Dent. Mater. 27, 2938. Ferracane, J.L., 2017. Models of caries formation around dental composite restorations. J. Dent. Res. 96, 364371. Ferracane, J.L., Hilton, T.J., Stansbury, J.W., Watts, D.C., Silikas, N., Ilie, N., et al., 2017. Academy of dental materials guidance-resin composites: Part II-Technique sensitivity (handling, polymerization, dimensional changes). Dent. Mater. 33, 11711191. Franco, E.B., Francischone, C.E., Medina-Valdivia, J.R., Baseggio, W., 2007. Reproducing the natural aspects of dental tissues with resin composites in proximoincisal restorations. Quintessence Int. 38, 505510. Fugolin, A.P.P., Pfeifer, C.S., 2017. New resins for dental composites. J. Dent. Res. 96, 10851091. Habib, E., Wang, R., Wang, Y., Zhu, M., Zhu, X., 2015. Inorganic fillers for dental resin composites: present and future. ACS Biomater. Sci. Eng. 2, 111. Heintze, S.D., Ilie, N., Hickel, R., Reis, A., Loguercio, A., Rousson, V., 2017. Laboratory mechanical parameters of composite resins and their relation to fractures and wear in clinical trials—a systematic review Siegward. Dent. Mater. 33, E101E114. Hervas-Garcia, A., Martinez-Lozano, M.A., Cabanes-Vila, J., Barjau-Escribano, A., FosGalve, P., 2006. Composite resins. A review of the materials and clinical indications. Med. Oral Patol. Oral Cir. Bucal 11, E215E220. Ilie, N., Hickel, R., 2009. Investigations on mechanical behaviour of dental composites. Clin. Oral Investig. 13, 427438. Ilie, N., Hilton, T.J., Heintze, S.D., Hickel, R., Watts, D.C., Silikas, N., et al., 2017. Academy of dental materials guidance-resin composites: Part I—Mechanical properties. Dent. Mater. 33, 880894. Kaizer, M.R., DE Oliveira-Ogliari, A., Cenci, M.S., Opdam, N.J.M., Moraes, R.R., 2014. Do nanofill or submicron composites show improved smoothness and gloss? A systematic review of in vitro studies. Dent. Mater. 30, e41e78. Khan, A.S., Azam, M.T., Khan, M., Mian, S.A., Ur rehman, I., 2015. An update on glass fiber dental restorative composites: a systematic review. Mater. Sci. Eng. C, Mater. Biol. Appl. 47, 2639. Kwon, T.Y., Bagheri, R., Kim, Y.K., Kim, K.H., Burrow, M.F., 2012. Cure mechanisms in materials for use in esthetic dentistry. J. Investig. Clin. Dent. 3, 316. Leprince, J.G., Palin, W.M., Hadis, M.A., Devaux, J., Leloup, G., 2013. Progress in dimethacrylate-based dental composite technology and curing efficiency. Dent. Mater. 29, 139156. Lohbauer, U., Belli, R., Ferracane, J.L., 2013. Factors involved in mechanical fatigue degradation of dental resin composites. J. Dent. Res. 92, 584591. Lutz, F., Setcos, J., Phillips, R., Roulet, J., 1983. Dental restorative resins. Types and characteristics. Dent. Clin. N. Am. 27, 697712. Maas, M.S., Alania, Y., Natale, L.C., Rodrigues, M.C., Watts, D.C., Braga, R.R., 2017. Trends in restorative composites research: what is in the future? Braz. Oral Res. 31, e55. Mainjot, A.K., Dupont, N.M., Oudkerk, J.C., Dewael, T.Y., Sadoun, M.J., 2016. From artisanal to CAD-CAM blocks: state of the art of indirect composites. J. Dent. Res. 95, 487495.
Resin-based dental composites for tooth filling 173 Mallineni, S.K., Nuvvula, S., Matinlinna, J.P., Yiu, C.K., King, N.M., 2013. Biocompatibility of various dental materials in contemporary dentistry: a narrative insight. J. Investig. Clin. Dent. 4, 919. Mikhail, S.S., Johnston, W.M., 2014. Confirmation of theoretical colour predictions for layering dental composite materials. J. Dent. 42, 419424. Nahsan, F.P.S., Mondelli, R.F.L., Franco, E.B., Naufel, F.S., Ueda, J.K., Schmitt, V.L., et al., 2012. Clinical strategies for esthetic excellence in anterior tooth restorations: understanding color and composite resin selection. J. Appl. Oral Sci. 20, 151156. Nayar, S., Ganesh, R., Santhosh, S., 2015. Fiber reinforced composites in prosthodontics—a systematic review. J. Pharm. Bioallied Sci. 7, S220. Nedeljkovic, I., Teughels, W., De munck, J., Van meerbeek, B., Van landuyt, K.L., 2015. Is secondary caries with composites a material-based problem?. Dent. Mater. 31, e247e277. Opdam, N.J.M., Collares, K., Hickel, R., Bayne, S.C., Loomans, B.A., Cenci, M.S., et al., 2017. Clinical studies in restorative dentistry: new directions and new demands. Dent. Mater. Pitel, M.L., 2013. Low-shrink composite resins: a review of their history, strategies for managing shrinkage, and clinical significance. Compend. Contin. Educ. Dent. 34, 578590. Radford, J.R., 2017. Longevity of direct and indirect resin composite restorations in permanent posterior teeth: a systematic review and meta-analysis. BDJ 222, 30. Rasines Alcaraz, M.G., Veitz-Keenan, A., Sahrmann, P., Schmidlin, P.R., Davis, D., IheozorEjiofor, Z., 2014. Direct composite resin fillings versus amalgam fillings for permanent or adult posterior teeth. Cochrane Database Syst. Rev. 31, Cd005620. Santini, A., 2010. Current status of visible light activation units and the curing of lightactivated resin-based composite materials. Dent. Update 37, 214216. 218-20, 223-7. Santini, A., Gallegos, I.T., Felix, C.M., 2013. Photoinitiators in dentistry: a review. Prim. Dent. J. 2, 3033. Schenkel, A.B., Peltz, I., Veitz-Keenan, A., 2016. Dental cavity liners for class I and class II resin-based composite restorations. Cochrane Database Syst. Rev. 10, Cd010526. Soares, C.J., Faria, E.S.A.L., Rodrigues, M.P., Vilela, A.B.F., Pfeifer, C.S., Tantbirojn, D., et al., 2017. Polymerization shrinkage stress of composite resins and resin cements— what do we need to know? Braz. Oral Res. 31, e62. Then, S., Neon, G.S., Abu kasim, N.H., 2011. Performance of melamine modified ureaformaldehyde microcapsules in a dental host material. J. Appl. Polym. Sci. 122, 25572562. Zimmerli, B., Strub, M., Jeger, F., Stadler, O., Lussi, A., 2010. Composite materials: composition, properties and clinical applications. A literature review. Schweiz Monatsschr Zahnmed 120, 972986.
8 Glass-ionomer cement: chemistry and its applications in dentistry Saroash Shahid1 and Tomas Duminis2 1 Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom, 2Centre for Oral Bioengineering, Institute of Dentistry, Barts and the London School of Medicine and Dentistry, London, United Kingdom Chapter Outline 8.1 Introduction 175 8.2 Development of glass-ionomer cements 176 8.3 Components of glass-ionomer cements 178 8.3.1 Composition and nature of the glass component 178 8.3.2 Composition and nature of the acid component 180 8.3.3 Water: the reaction medium 181 8.4 Chemistry of the setting reaction 181 8.4.1 Decomposition of the glass powder 182 8.4.2 Gelation phase 182 8.4.3 Maturation phase 183 8.5 Fluoride release from glass-ionomer cements 183 8.5.1 Source of fluoride 183 8.5.2 Mechanism of fluoride release 184 8.5.3 Factors effecting fluoride release 184 8.6 Mechanical properties 185 8.6.1 Compressive strength 185 8.6.2 Flexural strength 186 8.7 Esthetics 188 8.8 Chemical adhesion with tooth 188 8.9 Moisture sensitivity of glass-ionomer cements 189 8.10 Use of glass-ionomer cements in alternative restorative technique 8.11 Nanoapatite-filled glass ionomers 189 8.12 Thermo-cured glass ionomers 190 8.13 Resin-modified glass-ionomer cements 190 8.14 Glass ionomer as a “nondental” cement 191 References 191 Further reading 194 8.1 189 Introduction It has been almost 50 years since the first glass-ionomer cements (GICs), now officially termed as “glass polyalkenoate cements,” were introduced into the dental Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00008-6 Copyright © 2019 Elsevier Ltd. All rights reserved.
176 Advanced Dental Biomaterials world. In recent years this material has undergone further development resulting in the availability of a range of materials which undergo setting or “curing” via different proportions of acidbase and free-radical reactions. In the wake of these developments it seems essential to define what GICs are. To serve this purpose, the following definition seems an adequate description of the chemistry and properties of the material: The term glass-ionomer cement is reserved exclusively for a material consisting of acid-decomposable glass and a water-soluble acid that sets by neutralisation reaction which takes place within a clinically acceptable time. McLean (1994) Ever since its advent, GIC has received a mixed response from clinicians. It is acknowledged for its chemical adhesion with the tooth and continuing fluoride release (Wilson, 1989; Smith, 1998), but also it is subjected to criticism due to its lack of physical strength and translucency (Crisp et al., 1974; Mount and Makinson, 1982). Despite this criticism, the material has found its place in a broad spectrum of applications such as luting/lining cement; base, or dentine substitute under composite resin; sealant over an active carious lesion; and a restoration in its own right. The driving force in all these aforementioned applications is indeed the chemical adhesion with tooth and the longterm fluoride release (Anusavice, 2003; Mount and Hume, 1998; Mount, 1998). 8.2 Development of glass-ionomer cements The inspiration for the development of GIC can be traced back to the era of clinical dentistry which was governed by the use of amalgam, gold, and porcelain as restorative materials. The limitations of these materials drove the need for new materials, which could be used for luting/lining purposes and which also had a better esthetic appeal. This paved the way for the development of the zinc phosphate cement by Pierce in 1879 which was later further modified by Ames and Fleck (Ames, 1892; Fleck, 1902) to give the modern-day zinc-phosphate cement. Coinciding with Pierce’s discovery was Foster’s zinc oxideeugenol cement which soon gained popularity due to its obtundant effect on the pulp (Smith, 1998). Despite all these developments, the first breakthrough on the esthetics front was the silicate cement. Developed originally in 1873 by Thomas Fletcher, the silicate became famous in 1904, only after it had been modified by Steenbock of Berlin. Due to these ongoing developments, the end of the first quarter of the 20th century saw three basic types of dental cements, namely, zinc phosphate cement, zinc oxideeugenol cement, and silicate cement. Over the next 50 years, these cements underwent considerable technical improvement, but generally, the fundamentals of their setting chemistry remained unchanged. In 1963 the idea of polyelectrolyte cement was developed, which involved the reaction of metal oxide with reactive water-soluble polymers. Based on this concept, Smith produced the first zinc polyacrylate cement using zinc oxide and poly
Glass-ionomer cement: chemistry and its applications in dentistry 177 (acrylic acid). By 1964, this cement had been developed into a usable system. It was capable of chemically bonding to the tooth structure due to the use of poly (acrylic acid) which has the ability to complex calcium ions and form hydrogen bonds with collagen. After undergoing technological improvements and clinical trials, the material became commercially available in 1968. In the next decade the material underwent considerable development and materials containing fluorides and using copolymers became commercially available. In spite of the novel properties of these polyacrylate cements, the metal oxide component rendered them opaque and hence not esthetically acceptable. This led to the development of formulations using zinc-containing glass ceramics and silicate cement powders, but these had problems such as the prolonged setting time when mixed with poly(acrylic acid). Wilson and Kent made a major development in this direction by taking poly(acrylic acid) and mixing it with an ion-leachable glass (Wilson and Kent, 1972). The resultant product was given the name ASPA that was an acronym for aluminosilicate polyacrylate. ASPA (Fig. 8.1) was a hybrid of the dental silicate cement and the carboxylate cement. The development of more reactive glasses with high fluorine content and the discovery of the effect of tartaric acid in improving the setting properties resulted in the first practical GIC in 1972. These earlier materials, however, had drawbacks mainly related to their sluggish setting. They showed prolonged sensitivity to moisture and when set they were slightly opaque. However, since these early materials appeared various modifications had been carried out which have led to materials with better characteristics. These modifications included the following: 1. Use of alternative polymers such as acrylic/maleic acid as the polyacid component (Nicholson, 1998). 2. The use of vacuum-dried polymer powders blended with the glass and activated by the addition of water (Prosser et al., 1984). 3. The development of cermet containing cements in which the filler consists of a cermet which is silver sintered to the surface of calcium fluoroaluminosilicate glass. These Figure 8.1 Aluminosilicate polyacrylate (ASPA), the first commercial glass-ionomer cement (GIC).
178 Advanced Dental Biomaterials materials have their use in situations where radiopacity is required and for core buildup under crowns (McLean and Gasser, 1985). 4. Metal-reinforced cements in which a metal such as a silver tin alloy or stainless steel is added as an inert filler to conventional glass-ionomer in an attempt to reinforce the set cement (Williams et al., 1992). 5. Resin-modified cements in which the conventional acidbase reaction of GIC are complemented by photochemical polymerization of added monomers and initiators. 8.3 Components of glass-ionomer cements Conventional GICs have the following constituents: G G G Glass powder: A specially formulated acid-degradable fluoroaluminosilicate glass. Acid: The acid is a homopolymer of acrylic acid or a related copolymer. (1)Tartaric acid is added to improve the working and sharpen the setting time. Water: The reaction medium. 8.3.1 Composition and nature of the glass component The glasses used in GIC are fluoroaluminosilicates. These are different from glasses we see around us, for example in windows, which are mainly soda lime silica glasses. Window glasses possess highly cross-linked OSiO linkages and are chemically stable in nature and do not react with conventional acids. Such glasses would serve no purpose in GIC since in order to produce set cement it is essential that the glass is reactive in nature and capable of releasing metal cations. However, the addition of network modifying cations to these glasses disrupts the OSiO linkages to produce nonbridging oxygens (BOs) (NBOs). The resulting glass then represents an ionic polymer where the positive charge of the cation balances the negative charge on the network (produced by NBO). For glasses used in GIC, alumina acts as the intermediate oxide, that is, it can take part in glass structure but cannot form glass on its own. The addition of alumina produces negative sites such as the AlO2 4 tetrahedra makes these glasses “acid decomposable,” hence allowing them to be used in cement formation. Fluoroaluminosilicates are produced by the fusion of silica, alumina, calcium fluoride, and calcium oxide with aluminum fluoride, aluminum phosphate, and cryolite (Na3AlF6) added in minor quantities. Structurally, these glasses are based on AlO4 and SiO4 tetrahedra with a threedimensional glass structure (Fig. 8.2). O is at the vertex and Si and Al are at the center of the tetrahedron. Alkaline ions or alkaline earth ions such as Na1, Ca21, and Sr21 exist near the 31 Al ion in order to charge balance AlO2 4 tetrahedra. These modifying ions can also result in the formation of NBO. It was originally believed that fluoride and phosphate ions included in the glass structure do not form a part of the skeletal structure of the silicate network (Davidson and Mjör, 1999); however, some studies
Glass-ionomer cement: chemistry and its applications in dentistry 179 Figure 8.2 Skeletal structure of fluoroaluminosilicate glass (a tetrahedron) (Davidson and Mjör, 1999). (Note: The structure is 3D but is drawn in 2D for simplicity.) Table 8.1 Components of fluoroaluminosilicate glass and their effects. Component Percentage Effects SiO2 29.0 Al2O3 16.6 CaF2 34.2 These are three essential components of the glass which fuse to form the calcium fluoroaluminosilicate glass. Glasses higher in SiO2 (more than 40%) are more translucent, whereas those high in CaF2 or Al2O3 are more opaque. Complements the fluxing action of CaF2, i.e., reduce fusion temperature Improves translucency and adds body to the cement paste Used to replace calcium fully or partially to give radiopacity to the glass Na3AlF6 AlPO4 Sr, Ba, La, salts 5.0 9.9 Source: Compiled from Mount G.J., Hume W.R., 1998. Preservation and Restoration of Tooth Structure. Mosby, London; Wilson A.D., McLean J.W., 1988. Glass-Ionomer Cement. Quintessence, Chicago, IL. have shown the existence of phosphate as a network former, and there is strong evidence that PO41 tetrahedra locally charge balance AlO2 4 tetrahedra. The glasses are commercially prepared by fusing the mixture of its ingredients (Table 8.1) at 1200 C1550 C. In the next step the melt glass is shock cooled resulting in the formation of coarse granules which are then ground by dry milling in a ball mill to produce a particle size depending upon its application. Generally, a particle size of ,50 µm is preferred for restorative cements, whereas ,20 µm is produced for the luting type of material (Wilson and McLean, 1988). In either case ball milling is often followed by acid-washing of the glasses with 5% aqueous acetic acid. This is done in order to reduce the reactivity of the glass. Early work on simple glasses suggests that the Al2O3/SiO2 ratio is crucial in determining the glass reactivity (Wilson et al., 1979); however, more recently
180 Advanced Dental Biomaterials Griffin and Hill (1999) have provided evidence that this ratio has no significant influence on the properties of GICs derived from fluoro-alumino-phosphosilicate glasses. Fluoride fluxes (CaF2 and Na3AlF6) apart from lowering the fusion temperature of the glass, impart the unique property of fluoride release from the cement. Griffin and Hill (2000) reported that the glass transition temperature, Tg, for glasses prepared in their experiments falls by approximately 200 C when the amount of fluorine is increased. They explained that the reduction in the temperature is caused by the replacement of BOs by fluorines to form NBOs in the glass. Thus it disrupts the network and allows the relaxation phenomena to take place at a lower temperature. In addition, fluorine converts Ca21 to F-Ca(n) reducing the number of NBOs attached to Si. The presence of fluorine in the glass structure also affects the rheological and setting properties of GIC. Fluorine-containing glass will be more “network disrupted” and hence more susceptible to acid attack. Crisp and Wilson (1974a,b) found that increasing the fluoride content aids in the extraction of ions from the glass. De Barra and Hill (2000) also found that increasing the content of fluorine in a glass reduces the working and setting times of the cement. In most commercial formulations, calcium is wholly or partially replaced by strontium, barium, or lanthanum to impart radiopacity to the set cement. Strontium seems to be an ideal candidate since it has a similar ionic radius to that of calcium and hence can replace it without disrupting the glass and produce any loss of translucency (Deb and Nicholson, 1999). 8.3.2 Composition and nature of the acid component Acids used in the conventional glass-ionomer system are polyelectrolytes. These, as the name denotes, are both polymers and electrolytes and hence are soluble in water. Polyelectrolytes used in GIC are the poly(alkenoics). These polyacids include the homopolymers and copolymers of unsaturated mono-, di-, and tricarboxylic acids. Of these, the most important used to date have been poly(acrylic acid)s, copolymers of acrylic and itaconic acids (Crisp et al., 1980), and copolymers of acrylic and maleic acids. The polyacids are generally prepared by free-radical polymerization of the appropriate monomers in aqueous solution in the presence of a chain transfer agent and an initiator such as ammonium persulfate. These polyacids consist of linear chains with no cross-linking but have pendant acid groups (Wilson and McLean, 1988). To produce a cement of adequate mechanical properties it is essential that the polyacids used have a molecular weight in the range of 40,000 to 60,000—the higher the molecular weight the better will be the mechanical properties—however, in practice molecular weight is limited by viscosity, and some balance has to be achieved between concentration, viscosity, and molecular weight (Wilson et al., 1989). The polyacid used in the original glass ionomer was 50% by weight aqueous solution of poly(acrylic acid) (Crisp et al., 1975). However, poly(acrylic acid) molecules being flexible and mobile are capable of intermolecular hydrogen bonding. This results in gelation of poly(acrylic acid) solutions on storage prepared at a concentration greater than 45% by weight. It was found that methylation of the poly
Glass-ionomer cement: chemistry and its applications in dentistry 181 (acrylic acid) inhibits the gelation process (Crisp et al., 1975); however, cements prepared from such a liquid stained in the mouth. Crisp and Wilson (1974a,b) hypothesized that the copolymers of acrylic acid with other saturated carboxylic acids would be less susceptible to gelation due to reduced stereoregularity. Based on this assumption, they synthesized a copolymer of acrylic and itaconic acid which proved to be indefinitely stable to gelation at a concentration of 50% m/m. Due to a greater number of carboxylic acid groups, cements produced from copolymers with higher acid functionality were found to be harder than those of poly(acrylic acid) (Mount and Makinson, 1982). However, it was also observed that cements based on acrylic acid copolymers show less adhesion with tooth structure (Aboush and Jenkins, 1986) and are less resistant to acid attack (Setchell et al., 1985; Wilson et al., 1986) than cements of poly(acrylic acid). Tartaric acid (1) is an essential additive in the glass-ionomer system. It is added at a level of 5% or 10% in order to improve the handling properties, which it does by extending the working time and sharpening the setting time (Crisp et al., 1975). Using fourier transform infrared spectrsocopy (FTIR) analysis, Nicholson (1998) suggested that this is because (1) tartaric acid delays the formation of calcium carboxylate. Raman analysis also confirms this finding. An increase in compressive strength (CS) has also been reported with the use of tartaric acid (Crisp et al., 1975). 8.3.3 Water: the reaction medium GICs are water-based materials, and hence water plays an important role in their setting and structure. It serves as a solvent and a medium for transporting the ions. Apart from this, water also hydrates the siliceous hydrogel and the metal poly(acrylate) salts formed (Wilson and Mclean, 1988). Water is present in the set cement in at least two different states depending on whether they can be removed or not by desiccation over a suitable desiccant. These states have been classified as “evaporable” and “nonevaporable” or more commonly referred to as “loosely bound” and “tightly bound” water, respectively (Wilson et al., 1979). Loss of loosely bound water during initial setting stages retards the setting and produces surface crazing and cracks, whereas moisture contamination at this stage results in loss of soluble ions producing weak cements (Wilson et al., 1979; Causton, 1981). As the cement matures, the ratio of tightly bound water to loosely bound water increases, decreasing its susceptibility to desiccation accompanied by an increase in strength and modulus and a decrease in plasticity (Paddon and Wilson, 1976; Wilson and Kent, 1972). 8.4 Chemistry of the setting reaction The setting reaction is an acidbase reaction with the ion-leachable glass acting as a base and the acid being poly(acrylic acid) or its copolymer (Crisp and Wilson, 1974a,b).
182 Advanced Dental Biomaterials The setting process is in three overlapping stages (Crisp and Wilson, 1974a,b): Stage 1: The decomposition of the glass powder; Stage 2: Gelationprecipitation of cations and anions; Stage 3: Maturation phase. 8.4.1 Decomposition of the glass powder This phase may be best described as the ion leaching or extraction phase. During this stage, ionization of the carboxylic acid liberates protons (H1) from the carboxyl (COOH) group. These protons attack the surface of the glass liberating Al31, Ca21, Na1, F2 (if present), and H2PO24 ions into the aqueous phase (Wilson and Prosser, 1982). Barry et al. (1979) found that this attack was not uniform but seemed to take place more at calcium-rich sites in the glass since these areas are more basic. Whatever the nature of the attack, it results in the formation of silicic acid which later condenses to form silica gel (Wasson and Nicholson, 1990, 1993). Crisp and Wilson (1974a,b) predicted that as a result of this attack, 20%30% of the glass particles are degraded. However, Billington et al. (2006) argue that approximately 7% of the glass is consumed to produce the cement matrix. This degradation of the glass surface and cross-linking of polyacids is accompanied by a rise in pH (Crisp and Wilson, 1974a,b) and the viscosity of the cement paste (Wilson and Mclean, 1988). 8.4.2 Gelation phase As the reaction continues, ions released from the glass accumulate and are removed by precipitation as insoluble polyacrylates. Al31, Ca21, and F2 form metallic salt bridges with free (COO2) groups resulting in the cross-linking of the polycarboxylate chains and ultimately leading to the setting. This process continues until all the ions are insoluble. It is important to note that not all COOH groups lose their hydrogen ions to form COO2 ions. This is because when most of the carboxylic groups have ionized, the negative charge on the polymer chain increases. This results in the positively charged ions becoming more strongly bound to the polymer chains, hence making it difficult for the metal ions to replace them. Furthermore, as the density of the cross-links increases, it hinders the movement of metal ions toward carboxyl sites. Due to these reasons, the neutralization reaction does not quite go to completion. It has been shown that calcium polyacrylate is largely responsible for the initial gelation and setting of the cement in phase-separated glasses producing a clinically hard material within 410 minutes of mixing. This is followed by a slower formation of aluminum salts which may take up to 48 hours for completion (Crisp et al., 1974a,b; Barry et al., 1979). However, FTIR analysis has shown that, finally, both salts are present in equal quantities. The fluoride and phosphate ions form insoluble salts and complexes. Na1 contributes to the formation of orthosilicic acid on the surface of the glass particles which converts to silica gel as the pH rises. The silica gel assists in binding the powder to the matrix (Mount and Hume, 1998).
Glass-ionomer cement: chemistry and its applications in dentistry 183 8.4.3 Maturation phase Soon after the gelation phase, the material undergoes what is commonly known as the “maturation phase.” During this stage, the material has been shown to increase in CS. According to Wilson and Prosser (1982), this is due to the less mobile aluminum cross-linking the remaining the COOH acids or by replacing the already cross-linked calcium ions. More recently, Pires et al. (2004) have pointed at the maturation-related release of the more covalently bound “network forming” aluminum ions to complement the cross-linking phenomenon. Interestingly, Boyd and Towler (2005) have observed a maturation phenomenon in a GIC formulated from aluminum-free calcium-zinc-silicate glasses. Since the glass did not contain any “less mobile” aluminum ions, it suggests the existence of other mechanisms to be associated with improvements in mechanical properties as a result of maturation phenomenon. Wasson and Nicholson (1990) using inductively coupled plasma—optical emission spectrometry (ICP-OES) found that silica is released in large amounts during the decomposition of the glass and hypothesized its role in cement formation. Such a release has also been observed by De Maeyer et al. (2002) using FTIR analysis. Wasson and Nicholson (1993) demonstrated the role of silica in the maturation process and associated it with the increase in CS of the cements. Although a possible candidate, the authors did not look at the presence of factors other than silica. Using IR and NMR analysis, Matsuya et al. (1996) have also related the increase in CS with the silica phase in the matrix, but they too ignored the presence of any complimentary mechanisms. Apart from the above factors, hydration has also been associated with the maturation phenomenon. Wilson et al. (1979) found an increase in the ratio of bound to unbound water with maturation and have correlated it with the increase in CS. The presence of free and bound water has also been observed in silica desiccants. Since the posthardened silica network formed in GIC has been found to be similar to that of amorphous silica (Matsuya et al., 1996), it may give a possible explanation for this phenomenon of free and bound water in GIC. 8.5 Fluoride release from glass-ionomer cements GICs are known to release clinically beneficial amounts of fluoride (Guida et al, 2002; Khouw-Liu et al., 1999; Jones et al., 2003). The irony, however, is that leachable fluoride had not been intentionally included to make the cement suitable for dental use, but it occurred as a constituent of several chemicals used in the manufacture of the glass (Williams et al., 2002). 8.5.1 Source of fluoride It is not clear as to how and from where the fluoride ion arrives at the surface of the cement. However, it is hypothesized that all intrinsic fluoride must initially come from the glass (Williams et al., 2002). Fluoride in the glass is released into
184 Advanced Dental Biomaterials the matrix during the setting reaction. An important point to note here is that the fluoride is released from the cement along with sodium and perhaps calcium ions in order to maintain the electrical neutrality of the cement, and due to this dependency on other ions the amount of fluoride released from the cement cannot be determined solely by the quantity of fluoride in the initial glass (Thevadass et al., 1996). 8.5.2 Mechanism of fluoride release During the setting of GIC, the fluoride ions arise initially from the surface of the glass particles which are then held in the siliceous gel of the matrix. Since fluoride ions do not form part of the cement matrix, thus a continuous exchange of fluoride can occur from the cement depending on the gradient of fluoride in the oral environment at any given time. However, the mechanism behind the release of fluoride from GICs has not been understood completely, but all researchers seem to agree on the kinetics associated with the fluoride release according to which at least two processes are occurring simultaneously, that is, a short-term fluoride burst release and a long-term fluoride release. This can be demonstrated by the equation below which adequately and accurately describes the cumulative fluoride [F]c release from a restorative glass ionomer in distilled water (De Moor and Verbeeck, 1998). pffi ð½FI tÞ 1β t ½Fc 5 t 1 t1=2 where [F]I 5 maximum value of fluoride released during short-term reaction; t 5 time; t1/2 5 “half-life” of short-term release, that is, the time required for the fluoride released by short-term reaction to reach half of its maximum value; β 5 a constant which is measure for the driving force of the long-term release. A large amount of the fluoride released from GICs is during the “short-term release” phase which occurs within a few days of placement in the oral cavity. This amount then declines sharply during the first week and stabilizes after 23 months. The long-term release of fluoride is substantially lower, but it appears to be sufficient to prevent caries (Mount and Hume, 1998). The process of diffusion seems to be associated with the long-term release of fluoride from glass ionomers. This is evident by Forsten’s (1990) study which showed the existence of a linear relationship between the amounts of fluoride released and the square root of time, thus following the theory of simple diffusion. Another study, by Williams et al. (1999), also pointed toward a relationship between the surface area and the amount of fluoride released, which again shows the existence of a diffusion mechanism. However, an erosive mechanism also seems to complement the fluoride release since studies have shown an increase in fluoride release in an acidic environment (Thevadass et al., 1996). This may account for the morphological changes which are frequently observed. 8.5.3 Factors effecting fluoride release Fluoride release has been shown to be dependent upon variables such as the composition of the glass and polyalkenoic acid, the relative proportions of the constituents
Glass-ionomer cement: chemistry and its applications in dentistry 185 in the cement mix, the mixing process, and the elution medium (De Witte et al., 2000). The relationship between the composition of aluminosilicate glasses and its effect on fluoride release can be observed from Williams et al.’s (2002) study where a sodium-free glass LG26 released lower fluoride than a sodium-containing glass G338. This may be attributed to the high solubility of sodium. However, in glasses where calcium is substituted with strontium for the purpose of radiopacity, a substantial increase in fluoride has not been observed. The only slight increase by strontium has been attributed to its increased density which affects the powder/liquid mixing ratio (Guida et al., 2002). The amount of fluoride release in different elution media is an interesting phenomenon to observe since the oral cavity provides an environment of varying pH and ionic concentration. Various researchers have concurred that there is an increase in fluoride release in an acidic environment, due to surface degradation (De Moor and Verbeeck, 1998; Czarnecka et al., 2002). This increase in fluoride release at a lower pH enhances the anticariogenic properties of GIC since the caries process also decreases the pH of the environment surrounding the tooth. Furthermore, their study also showed the release of more fluoride in an acidic environment. This increase in fluoride release was also previously observed by De Moor and Verbeeck (1998). Hadley et al. (2000, 2001) have also observed the phenomenon of surface degradation on exposure of GICs to neutral NaF solution. Since the solution used in this case was a neutral one, the presence of ions in the elution medium may be a possible explanation for the surface roughness. Among the various elution media used for fluoride release, artificial saliva is an important one since it provides a close emulation of the oral environment. Mallakh and Sarkar (1990) found that glass ionomers released more fluoride in deionized water than in artificial saliva. This pattern of fluoride release was also observed by Williams et al. (1997, 2001) using the same artificial saliva. Although the mechanism behind the reduction has not been understood, but it has been proposed to be either due to the higher ionic strength of the artificial saliva or probably the calcium in the artificial saliva forms an insoluble calcium fluoride layer on the surface of the cement (Williams et al., 2001). The findings of Williams et al. (1997, 2001) and Mallakh and Sarkar highlights the difference in fluoride release in artificial saliva compared with that in deionized water. This not only questions the clinical correlation of studies in deionized water, but it also suggests that further work is required in this area in order to understand the pattern of fluoride release in the oral environment. 8.6 Mechanical properties 8.6.1 Compressive strength Compressive strength (CS) is the maximum stress a material can withstand under a constant load over a given area. CS can be expressed by the following equation: σ5 F A
186 Advanced Dental Biomaterials where F is the applied load in unit newton and A is the area in meters squared. CS is the only strength test specified by the International Organization for Standardization standard for water-based cements for use in dentistry (ISO 99172:2017). The ISO standard defines CS, relative to the mode of measurement, as: c5 4F πd2 where F is the maximum compressive force applied in unit newton, and d is the sample diameter in millimeters. Compared to the alternative mechanical strength tests, the CS test is relatively simple and has a fast turnaround. However, the use of this test in the assessment of GIC mechanical failure has been criticized since the 1990s (McCabe et al., 1990). McCabe et al. (1990) found that the CS test is unreliable because of the unacceptable variation between the test results. McCabe et al. (1990) proposed that “standards should place less emphasis on the mean value of relatively small number of test specimens and the use of a simple form of probability theory in which, say, 80 per cent of specimens are required to achieve a certain pass level” proposing that some standards may need to increase the number of test specimens to achieve meaningful results. Based on the works of Kendall (1978), Darvell (1990), Fennell and Hill (2001a), and Baig and Fleming (2015) argue that stress at failure calculation does not take into account the failure mechanism operating during the test, whereby cylindrical specimens fail by “some unresolved combination of tension and shear” stresses. From international standards point of view, the minimum CS value for a water-based restorative cement is 100 MPa, whereas for a water-based lining cement the specified value is 50 MPa (ISO 9917-2:2017). Nevertheless, some commercial GICs, such as the AMALGOMER (manufactured by Advanced Healthcare Ltd, Kent, United Kingdom) considerably surpass mechanical properties specified in the standard and can withstand loads over 400 MPa. 8.6.2 Flexural strength An alternative approach to help us understand the mechanical properties of glass ionomer cements is the assessment of their flexural strength (FS). Flexural strength can be defined by the following equation: σ5 F bd where F is the axial load applied at the fracture point; b is the width of the sample; and d is the thickness of the sample (Fig. 8.3). For the three-point bend measurement purposes (Fig. 8.3), the flexural strength can be expressed by the following equation: σ5 3FL 2bd2
Glass-ionomer cement: chemistry and its applications in dentistry 187 Figure 8.3 Schematic representation of three-point flexural bend test. Figure 8.4 Flexural strength versus PAA molar mass (Fennell and Hill, 2001c). where F is the axial load force applied, in unit newton; L is the length of the sample support span; b is the width of the sample; and d is the thickness of the sample. In the flexural strength measurement, fracture occurs at midpoint under tension. In a study by Fennell and Hill (2001a) it was concluded that the dominant parameters which have an effect on the FS of GICs are the poly(acrylic acid) (PAA) molar mass (Fig. 8.4) and the PAA concentration (Fig. 8.5). Prosser et al. (1984) also reported that the flexural strength of GICs depends on the glass composition and the polyelectrolyte used to prepare them. The authors noted that opaque and opal glasses containing crystallites tended to yield cements with high FS. FS was also found to increase with increasing molecular weight of the poly(acrylic acid). In water-setting cements, the strength of the cements was shown to be critically dependent on the glass/polyacid ratio. In a study by Dowling et al. (2012), compressive strength, three-point flexural strength (TFS), and biaxial flexural strength (BFS) data from three commercially available GICs were
188 Advanced Dental Biomaterials Figure 8.5 Flexural strength versus PAA concentration (Fennell and Hill, 2001b). statistically compared for reproducibility (P 5 .05). Based on the statistical analysis of this data, Dowling et al. (2012) found that TFS and BFS were reproducible; however, the experimental data were found to vary significantly, and as such the study concluded that CS measurement is not a statistically valid measurement. The authors could not show any statistical difference between TFS and BFS. The authors also noted that the specimen fabrication for BFS is relatively simple and less operator-dependent as compared to the specimen fabrication for TFS. 8.7 Esthetics The development of the first glass ionomer cement in 1968 was considered to be a big breakthrough in esthetic dentistry because the new cement was translucent. Since the development of the first generation of GICs, the esthetics of modern GICs have improved, mainly as a result of better understanding of the structureproperty relationship. Highly translucent glass ionomer cements are produced by matching the refractive index (RI) of the glass component to the liquid component to avoid light scattering at the glass-polysalt interfaces. However, other factors such as the poly(acrylic acid) concentration and the particle size can also influence the esthetics as well as other properties of the cement. Most commercial and model laboratory GIC glass compositions contain a significant amount of fluoride. It has been shown that fluoride content in GIC glasses can be correlated to the refractive index of the glasses (Duminis et al., 2018). Therefore, varying the fluoride content in GIC glasses can be used as a tool to fine-tune the refractive index of the core glass component so that it closely matches to the liquid component. 8.8 Chemical adhesion with tooth One of the important properties of GICs is their chemical adhesion with the tooth without the use of intermediate bonding material. The process of chemical adhesion
Glass-ionomer cement: chemistry and its applications in dentistry 189 is through two mechanisms. First is an ion exchange process where the polyacrylic acid reacts with the hydroxyapatite in the tooth and forms bonds with calcium and phosphate. The second mechanism is by hydrogen bonding with the collagen present in dentine. Although bond strength between GIC and tooth is weak, evidence suggests that the bond is “self-healing,” that is, it can reform if it is broken. 8.9 Moisture sensitivity of glass-ionomer cements During the early stages of setting, GICs are very sensitive to moisture/saliva contamination and desiccation as these can disrupt the matrix resulting in a weak cement. It is essential that GICs be protected up to at least 24 hours after placement. This can be achieved by applying Vaseline, cocoa butter, or varnishes over the surface of the material soon after placement. 8.10 Use of glass-ionomer cements in alternative restorative technique Alternative restorative treatment (ART), formerly known as atraumatic restorative treatment, is defined as “a dental caries treatment procedure involving the removal of soft, demineralised tooth tissue using hand instrument alone, followed by restoration of the tooth with an adhesive restorative material, routinely glass ionomer.” This technique has been endorsed by the World Health Organization and the International Association for Dental Research as a means of restoring and preventing caries. ART may be used to restore and prevent caries in young patients, uncooperative patients, or patients with special healthcare needs, or when traditional cavity preparation and/or placement of traditional dental restorations are not feasible. GIC is the material of choice for ART, because of its bonding to enamel and dentin, fluoride release, and ease of use. Van Hof et al. concluded, based on a meta-analysis, that single-surface ART restorations using high-viscosity GIC in both primary and permanent dentitions showed high survival rates, and that medium-viscosity (traditional) GIC should not be used for ART restorations. 8.11 Nanoapatite-filled glass ionomers GICs have undergone extensive developments and modifications since it was first developed as ASPA. These include improvements in radiopacity, esthetics, and mechanical properties. Of notable mention among these recent developments is Glass Carbomer by GCP Dental, Netherlands. The glass ionomer contains nanofluorapatite/hydroxyapatite as a secondary filler in the formulation to enhance the remineralization of the tooth tissue. Phosphate ions are essential in the remineralization of the tooth, however including too much phosphate in the glass composition
190 Advanced Dental Biomaterials results in a slow setting material. In Glass Carbomer the manufacturers overcome this problem by including apatite phase as a secondary filler. The nanofluor/ hydroxyapatite in the material acts as nuclei for apatite formation, thus promoting remineralization. The enhanced remineralization is evident through the formation of a “pseudo-enamel” apatite phase which has been reported by an in vivo study (Duinen et al., 2004). 8.12 Thermo-cured glass ionomers Thermo-curing is a relatively new technique which uses radiant heat to accelerate the setting reaction of conventional GICs. This technique helps overcome the problem of early moisture sensitivity of GICs. Glass Carbomer is a commercially available GIC which has been specifically developed for use with thermo-curing. Thermo-curing is achieved by a special light curing device which emits heat. These portable handheld devices are available from Glass Carbomer manufacturer and are like the LED light curing units used for dental composites. In a clinical setting, thermo-curing is performed soon after placing the filling in the cavity by placing the tip of the light-cured unit on the surface of the GIC. In the case of Glass Carbomer, the manufacturer recommends a curing time of 90 seconds. Various researchers have investigated the effects of thermo-curing on mechanical properties of various GICs and have reported an increase in surface microhardness and FS. One clinical study has reported the beneficial effect of thermo-curing on the success rate of ART restorations for up to 18 months. 8.13 Resin-modified glass-ionomer cements Resin-modified GICs (RMGICs) can be regarded as fast setting GICs with immediate resistance to water uptake and some protection from water loss. Their composition consists of calcium fluoroaluminosilicate glass powder which is mixed with the liquid-containing hydroxyethylmethacrylate (HEMA), polyacrylic acid, and polymerization initiators. On mixing the powder with the liquid, two distinct types of setting reactions occur: G G acidbase neutralization and free-radical addition polymerization. These reactions may lead to the formation of two matrices, an ionomer salt hydrogel and polyHEMA matrix with multiple cross-linking. Full properties reached after maturation occurs as in an autocure system. Due to the presence of resin phase, RMGIC is prone to polymerization exotherm and polymerization shrinkage. Furthermore, since RMGICs contain less polyacrylic acid as compared to conventional GICs, there may be an effect on the chemical adhesion with the tooth. Although RMGICs initially provide improved mechanical properties, over
Glass-ionomer cement: chemistry and its applications in dentistry 191 time these properties tend to deteriorate due to water absorption by polyHEMA. The water uptake also makes the RMGIC prone to intrinsic staining which cannot be removed by polishing. 8.14 Glass ionomer as a “nondental” cement Glass ionomers received considerable attention during the early 1990s as cements in other branches of medicine. Excellent biocompatibility was reported (Sasanaluckit et al., 1993), and in one study Meyer et al. (1993) argued that although aluminum had found its way into the cells, there was no evidence of cytotoxicity, possibly due to the formation of nontoxic complexes with silica. Clinical trials reported good results for use in procedures such as alveolar ridge to build up, reconstruction following removal of bone tumors (Lindeque and Jonck, 1993), fixation of cochlear implants (Ramsden et al., 1992), and craniofacial reconstruction surgery (Zollner et al., 1994). However, these materials received a major setback when they were withdrawn from clinical use in France in 1994 (Renard et al., 1994). This was following severe postoperative problems with two patients who underwent translabrynthic otoneosurgery and bone reconstruction with GICs. Extremely high levels of aluminum were observed in the lumbar cerebrospinal fluid, serum, and urine. Since the case report of this incident is poorly written, it is unclear if this event may have occurred due to operator error in the surgical technique. Following this series of events, researchers have produced several formulations of aluminum-free glasses for use in glass ionomers for such purposes. These glasses either use zinc (Boyd and Towler, 2005) or iron to replace aluminum. References Aboush, Y.E., Jenkins, C.B., 1986. An evaluation of the bonding of glass-ionomer restoratives to dentine and enamel. Br. Dent. J. 161 (5), 179184. Ames, W.V.B., 1892. A new oxyphosphate cement for crown setting. Dent. Cosmos. 24, 906. Anusavice, K.J., 2003. Philips’ Science of Dental Materials, 11th ed. Elsevier, India. Baig, M.S., Fleming, G.J., 2015. Conventional glass-ionomer materials: a review of the developments in glass powder, polyacid liquid and the strategies of reinforcement. J. Dent. 43 (8), 897912. Barry, T.I., Clinton, D.J., Wilson, A.D., 1979. The structure of a glass-ionomer cement and its relationship to the setting process. J. Dent. Res. 58, 10721079. Boyd, D., Towler, M.R., 2005. The processing, mechanical properties and bioactivity of zinc based glass ionomer cements. J. Mater. Sci. Mater. Med. 16 (9), 843850. Causton, B.E., 1981. The physico-mechanical consequences of exposing glass ionomer cements to water during setting. Biomaterials 2 (2), 112115. Crisp, S., Wilson, A.D., 1974a. Reactions in glass ionomer cements: I. Decomposition of the powder. J. Dent. Res. 53 (6), 14081413.
192 Advanced Dental Biomaterials Crisp, S., Wilson, A.D., 1974b. Reactions in glass ionomer cements: III. The precipitation reaction. J. Dent. Res. 53 (6), 14201424. Crisp, S., Pringuer, M.A., Wardleworth, D., Wilson, A.D., 1974. Reactions in glass ionomer cements: II. An infrared spectroscopic study. J. Dent. Res. 53 (6), 14141419. Crisp, S., Ferner, A.J., Lewis, B.G., Wilson, A.D., 1975. Properties of improved glassionomer cement formulations. J. Dent. 3 (3), 125130. Crisp, S., Kent, B.E., Lewis, B.G., Ferner, A.J., Wilson, A.D., 1980. Glass-ionomer cement formulations. II. The synthesis of novel polycarobxylic acids. J. Dent. Res. 59, 10551063. Czarnecka, B., Limanowska-Shaw, H., Nicholson, J.W., 2002. Buffering and ion-release by a glass-ionomer cement under near-neutral and acidic conditions. Biomaterials 23, 27832788. Darvell, B., 1990. Uniaxial compression tests and the validity of indirect tensile-strength. J. Mater. Sci. 25 (2A), 757780. Davidson, C.L., Mjör, I.A., 1999. Advances in Glass-Ionomer Cements. Quintessence Pub. Co. Deb, S., Nicholson, J.W., 1999. The effect of strontium oxide in glass-ionomer cements. J. Mater. Sci. Mater. Med. 10, 471474. De Maeyer, E.A., Verbeeck, R.M., Vercruysse, C.W., 2002. Infrared spectrometric study of acid-degradable glasses. J. Dent. Res. 81 (8), 552555. De Moor, R.J., Verbeeck, R.M., 1998. Effect of acetic acid on the fluoride release profiles of restorative glass ionomer cements. Dent. Mater. 14, 261268. De Witte, A.M., De Maeyer, E.A., Verbeeck, R.M., Martens, L.C., 2000. Fluoride release profiles of mature restorative glass ionomer cements after fluoride application. Biomaterials 21, 475482. Duminis, T., Shahid, S., Karpukhina, N.G., Hill, R.G., 2018. Predicting refractive index of fluoride containing glasses. Dent Mater. 34 (5), e83e88. Available from: https://doi. org/10.1016/j.dental.2018.01.024. el Mallakh, B.F., Sarkar, N.K., 1990. Fluoride release from glass-ionomer cements in deionized water and artificial saliva. Dent. Mater. 6, 118122. Fennell, B., Hill, R., 2001a. The influence of poly(acrylic acid) molar mass and concentration on the properties of polyalkenoate cements—Part II Young’s modulus and flexural strength. J. Mater. Sci. 36 (21), 51775183. Fennell, B., Hill, R., 2001b. The influence of poly(acrylic acid) molar mass and concentration on the properties of polyalkenoate cements—Part III—Fracture toughness and toughness. J. Mater. Sci. 36 (21), 51855192. Fennell, B., Hill, R., 2001c. The influence of poly(acrylic acid) molar mass and concentration on the properties of polyalkenoate cements—Part I—Compressive strength. J. Mater. Sci. 36 (21), 51935202. Fleck, D.J., 1902. The chemistry of oxyphosphates. Dent. Items Int. 24, 906. Forsten, L., 1990. Short- and long-term fluoride release from glass ionomers and other fluoride-containing filling materials in vitro. Scand. J. Dent. Res. 98, 179185. Guida, A., Hill, R.G., Towler, M.R., Eramo, S., 2002. Fluoride release from model glass ionomer cements. J. Mater. Sci. Mater. Med. 13, 645649. Hadley, P.C., Billington, R.W., Pearson, G.J., Williams, J.A., 2000. Effect of monovalent ions in glass ionomer cements on their interaction with sodium fluoride solution. Biomaterials 21, 97102. Hadley, P.C., Billington, R.W., Williams, J.A., Pearson, G.J., 2001. Interactions between glass ionomer cement and alkali metal fluoride solutions: the effect of different cations. Biomaterials 22, 31333138.
Glass-ionomer cement: chemistry and its applications in dentistry 193 Jones, F.H., Hutton, B.M., Hadley, P.C., Eccles, A.J., Steele, T.A., Billington, R.W., et al., 2003. Fluoride uptake by glass ionomer cements: a surface analysis approach. Biomaterials 24, 107119. Kendall, K., 1978. Complexities of compression failure. Proc. R. Soc. London, Ser. A— Math. Phys. Eng. Sci. 361 (1705), 245263. Khouw-Liu, V.H., Anstice, H.M., Pearson, G.J., 1999. An in vitro investigation of a poly (vinyl phosphonic acid) based cement with four conventional glass-ionomer cements. Part 1: Flexural strength and fluoride release. J. Dent. 27, 351357. Lindeque, B.G., Jonck, L.M., 1993. Ionogran—an ionomeric micro implant in bone tumour reconstruction. A clinical evaluation. Clin. Mater. 14 (1), 4956. Matsuya, S., Maeda, T., Ohta, M., 1996. IR and NMR analyses of hardening and maturation of glass-ionomer cement. J. Dent. Res. 75 (12), 19201927. McCabe, J., Watts, D., Wilson, H., Worthington, H., 1990. An investigation of test-house variability in the mechanical testing of dental materials and the statistical treatment of results. J. Dent. 18 (2), 9097. McLean, J.W., Gasser, O., 1985. Glass-cermet cements. Quintessence Int. 16, 333343. McLean, J.W., Nicholson, J.W., Wilson, A.D., 1994. Proposed nomenclature for glassionomer dental cements and related materials. Quintessence Int. 25, 587589. Meyer, U., Szulczewski, D.H., Barckhaus, R.H., Atkinson, M., Jones, D.B., 1993. Biological evaluation of an ionomeric bone cement by osteoblast cell culture methods. Biomaterials 14 (12), 917924. Mount, G.J., 1998. Clinical performance of glass-ionomers. Biomaterials 19, 573579. Review. Mount, G.J., Hume, W.R., 1998. Preservation and Restoration of Tooth Structure. Mosby, London. Mount, G.J., Makinson, O.F., 1982. Glass-ionomer restorative cements: clinical implications of the setting reaction. Oper. Dent. 7, 134141. Nicholson, J.W., 1998. Chemistry of glass-ionomer cements: a review. Biomaterials 19, 485494. Review. Paddon, J.M., Wilson, A.D., 1976. Stress relaxation studies on dental materials. 1. Dental cements. J. Dent. 4 (4), 183189. Pires, R., Nunes, T.G., Abrahams, I., Hawkes, G.E., Morais, C.M., Fernandez, C., 2004. Stray-field imaging and multinuclear magnetic resonance spectroscopy studies on the setting of a commercial glass-ionomer cement. J. Mater. Sci. Mater. Med. 15 (3), 201208. Prosser, H.J., Powis, D.R., Brant, P., Wilson, A.D., 1984. Characterization of glass-ionomer cements. 7. The physical properties of current materials. J. Dent. 12, 231240. Ramsden, R.T., Herdman, R.C., Lye, R.H., 1992. Ionomeric bone cement in neuro-otological surgery. J. Laryngol. Otol. 106 (11), 949953. Renard, J.L., Felten, D., Bequet, D., 1994. Post-otoneurosurgery aluminium encephalopathy. Lancet 344 (8914), 6364. Rudel, C., Zollner, W., 1994. Ionomeric cement—a bone glue for device fixation. Ear Nose Throat J. 73 (3), 189191. Review. Sasanaluckit, P., Albustany, K.R., Doherty, P.J., Williams, D.F., 1993. Biocompatibility of glass ionomer cements. Biomaterials 14 (12), 906916. Setchell, D.J., Teo, C.K., Khun, A.T., 1985. The relative solubilities of four modern glassionomer cements. Br. Dent. J. 158 (6), 220222. Smith, D.C., 1998. Development of glass-ionomer cement systems. Biomaterials 19, 467478.
194 Advanced Dental Biomaterials Thevadass, K.P., Pearson, G.J., Anstice, H.M., Davies, E.H., 1996. Method for enhancing the fluoride release of a glass-ionomer cement. Biomaterials 17, 425429. Van Duinen, R.N., Davidson, C.L., De Gee, A.J., Feilzer, A.J., 2004. In situ transformation of glass-ionomer into an enamel-like material. Am. J. dent. 17, 223227. Wasson, E.A., Nicholson, J.W., 1990. A study of the relationship between setting chemistry and properties of modified glass-poly(alkenoate) cements. Br. Polym. J. 23, 179183. Wasson, E.A., Nicholson, J.W., 1993. New aspects of the setting of glass-ionomer cements. J. Dent. Res. 72 (2), 481483. Williams, J.A., Billington, R.W., Pearson, G.J., 1992. The comparative strengths of commercial glass-ionomer cements with and without metal additions. Br. Dent. J. 172, 279282. Williams, J.A., Billington, R.W., Pearson, G., 1997. Silver and fluoride ion release from metal-reinforced glass-ionomer filling materials. J. Oral Rehabil. 24, 369375. Williams, J.A., Billington, R.W., Pearson, G.J., 1999. The influence of sample dimensions on fluoride ion release from a glass ionomer restorative cement. Biomaterials 20, 13271337. Williams, J.A., Billington, R.W., Pearson, G.J., 2001. A long term study of fluoride release from metal-containing conventional and resin-modified glass-ionomer cements. J. Oral Rehabil. 28, 4147. Williams, J.A., Billington, R.W., Pearson, G.J., 2002. The glass ionomer cement: the sources of soluble fluoride. Biomaterials 23, 21912200. Wilson, A.D., 1989. Developments in glass-ionomer cements. Int. J. Prosthodont. 2, 438446. Wilson, A.D., Kent, B.E., 1972. A new translucent cement for dentistry. The glass ionomer cement. Br. Dent. J. 132, 133135. Wilson A.D., McLean J.W., 1988. Glass-Ionomer Cement. Quintessence, Chicago, IL. Wilson, A.D., Prosser, H.J., 1982. Biocompatibility of the glass ionomer cement. J. Dent. Assoc. S. Afr. 37 (12), 872879. Wilson, A.D., Groffman, D.M., Powis, D.R., Scott, R.P., 1986. A study of variables affecting the impinging jet method for measuring the erosion of dental cements. Biomaterials 7 (3), 217220. Wilson, A.D., Hill, R.G., Warrens, C.P., Lewis, B.G., 1989. The influence of polyacid molecular weight on some properties of glass-ionomer cements. J. Dent. Res. 68 (2), 8994. Wilson, A.D., Paddon, J.M., Crisp, S., 1979. The hydration of dental cements. J. Dent. Res. 58 (3), 10651071. Further reading Brune, D., 1982. Heat treatment of glass ionomer, silicate, zinc phosphate and zinc polycarboxylate cements. Scand. J. Dent. Res. 90 (5), 409412. Fejerskov, O., Ekstrand, J., Burt, B.A. (Eds.), 1996. Fluoride in Dentistry. second ed. Munksgaard, Copenhagen. Fusayama, T., Katayori, T., Nomoto, S., 1963. Corrosion of gold and amalgam placed in contact with each other. J. Dent. Res. 42, 11831197. Hellwig, E., Lennon, A.M., 2004. Systemic versus topical fluoride. Caries Res. 38, 258262. Jenkins, G.N., 1999. Review of fluoride research since 1959. Arch. Oral. Biol. 44, 985992.
Glass-ionomer cement: chemistry and its applications in dentistry 195 Nicholson, J.W., Czarnecka, B., 2004. The interaction of lactic acid-glass cements with aqueous solutions. J. Mater. Sci. Mater. Med. 15 (2), 151154. Nicholson, J.W., Tawfik, H., Czarnecka, B., 2002. A study of cements formed by aqueous lactic acid and aluminosilicate glass. J. Mater. Sci. Mater. Med. 13 (4), 417419. Schmidt, W., Purmann, R., Jochum, P., Gasser, O., 1980. Mixing compounds for glassionomer cements and use of a copolymer for preparing the mixing components. Eur. Pat. Appl. 24, 056. Ten Cate, J.M., 2004. Fluorides in caries prevention and control: empiricism or science. Caries Res. 3825438257. Towler, M.R., Bushby, A.J., Billington, R.W., Hill, R.G., 2001. A preliminary comparison of the mechanical properties of chemically cured and ultrasonically cured glass ionomer cements, using nano-indentation techniques. Biomaterials 22 (11), 14011406. Wilson, A.D., Crisp, S., Abel, G., 1977. Characterization of glass-ionomer cements. 4. Effect of molecular weight on physical properties. J. Dent. 5 (2), 117120.
Impression materials for dental prosthesis 9 Payam Zarrintaj1,2,3, Sahba Rezaei4, Seyed Hassan Jafari4, Mohammad Reza Saeb2,3,5, Shadi Ghalami6, Mahsa Roshandel6, Brouki Milan Peiman7,8, Daghigh Ahmadi Ehsaneh9, Farshid Sefat10,11 and Masoud Mozafari7,8 1 Polymer Engineering Department, Faculty of Engineering, Urmia University, Urmia, Iran, 2 Color and Polymer Research Center (CPRC), Amirkabir University of Technology, Tehran, Iran, 3Advanced Materials Group, Iranian Color Society (ICS), Tehran, Iran, 4School of Chemical Engineering, College of Engineering, University of Tehran, Tehran, Iran, 5 Department of Resin and Additive, Institute for Color Science and Technology, Tehran, Iran, 6Department of Anatomy and Pathology, University of Siena, Siena, Italy, 7Cellular and Molecular Research Center, Iran University of Medical Sciences, Tehran, Iran, 8 Department of Tissue Engineering & Regenerative Medicine, Faculty of Advanced Technologies in Medicine, Iran University of Medical Sciences, Tehran, Iran, 9Centre for Nanohealth, College of Engineering, Swansea University, Swansea, United Kingdom, 10 Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom, 11Interdisciplinary Research Centre in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom Chapter Outline 9.1 Introduction 198 9.2 Elastic impression materials 9.2.1 9.2.2 9.2.3 9.2.4 9.2.5 9.3 Inelastic impression materials 9.3.1 9.3.2 9.3.3 9.3.4 200 Polyethers 200 Polysulfide 201 Alginate 201 Agar 202 Silicones 202 204 Impression wax 204 Impression compound 205 Impression plaster 205 Metallic oxide pastes (zinc oxideeugenol impression paste) 206 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00009-8 Copyright © 2019 Elsevier Ltd. All rights reserved.
198 Advanced Dental Biomaterials 9.4 Characteristics of impression materials 9.4.1 9.4.2 9.4.3 9.4.4 9.4.5 9.5 Conclusion and future perspective References 212 9.1 207 Dimensional accuracy/dimensional stability 207 Wettability 208 Elastic recovery/flexibility 208 Mechanical properties 209 Miscellaneous 209 212 Introduction Biomedical scientists have endeavored to repair the organs, mainly with the aid of regeneration or implanting strategies (Chiu et al., 2017; Zarrintaj et al., 2017b; Bakhshandeh et al., 2017). Dental health issues are related to the esthetic features of people who naturally desire for being perfectly evaluated for their beauty; hence, a wide range of materials have been utilized in dentistry with miscellaneous features from both beauty and health-care views (Hafshejani et al., 2017; Zamanian et al., 2013; Zarrintaj et al., 2018a). Indeed the key imprints of such concerns are impression materials (IMs) that have been widely used in dentistry. IMs have been utilized in prosthodontics (e.g., denture), orthodontics, restoration, maxillofacial prosthetics, diagnosis/treatment planning, and dental implants. IMs have been known as negative imprints of the mouth tissues, which are utilized for the positive formation of the teeth and juxtaposing tissues model. Various materials exhibit the appropriate properties as candidates for IMs, these can be classified into elastic (nonrigid) and inelastic (rigid) materials. Common elastic materials, when examined clinically, are reversible hydrocolloids, such as agar, which work on the basis of phase change caused by temperature rise; irreversible hydrocolloids such as alginate salt; elastomeric materials such as silicones; and polyethers. Plasters, zinc oxide, and eugenol-based impression pastes are categorized as rigid/inelastic IMs. Such materials are used for the dental arches using dental impression trays (Ting-shu and Jian, 2015). Fig. 9.1 depicts the restoration the lost teeth using IMs (Bhakta et al., 2011). Typically, choosing the appropriate substance for a particular clinical usage/ application depends on either expense or precision. For designing more precise materials, when they are needed for a special purpose and not routinely, particular considerations are required. For cases where undercuts are not present among the recording surfaces, stiff substances are preferred. Rigid IMs can also be utilized in edentulous subjects, where soft tissue (compressible) undercuts are present (McCabe and Walls, 2013). The reason for using the impression tray is to prepare the supporting matrix for the IMs before simply inserting the impression into the mouth. In general, two types of trays are available: custom-made and stock trays (Marotti et al., 2014). Stock trays are supplied in various shapes and size in order to help the clinician to choose the tray that fits well to the patient’s mouth. Such trays are often either under- or overextended in relation to the extent of the oral soft and
Impression materials for dental prosthesis 199 Figure 9.1 Exposed fixture head after removal of the healing abutment (A); closed tray impression coping screwed in place (B); light-bodied IM syringed around impression coping (C); impression taken in a stock tray (D); impression with details of soft tissue around the implant and adjacent teeth (E); and impression coping repositioned into the impression (F) (Bhakta et al., 2011). hard tissues, which need to be recorded for the clinical purposes so as to facilitate modification if needed (Pastoret et al., 2017). On the other hand, custom-made trays are constructed on the study casts of a patient’s mouth that have been obtained with IMs, thus having a greater accuracy when compared to the stock trays. For recording the oral soft tissues, such trays should be suitably extended in all orientations. Some methods are used to stick IMs to the trays including puncturing perforation in the tray, utilizing adhesives, and rim lock (McCabe and Walls, 2013; Tripathi et al., 2017). IMs can be applied in liquid state or solid-like state to form the shape of dentition and the surrounding structures of the oral cavity before being set. Materials used should possess the following acceptable features: they should have pleasant esthetic color/ taste/odor with no releasable toxic ingredients, should be biocompatible with no irritation to the tissue, have an appropriate shelf-life, be affordable, be dimensionally stable, and have sufficient mechanical strength. Techniques for taking impressions can be defined as mucostatic (normal resting position), mucocompressive (compression position), and selective pressure techniques (Oh and Morris, 2017). In order to construct a proper cast, IMs should be scrutinized; fulfilling this aim, in this overview, IMs and their properties are discussed. In the following sections, first, we focus on elastic IMs, followed by inelastic IMs, and finally their characteristics.
200 9.2 Advanced Dental Biomaterials Elastic impression materials 9.2.1 Polyethers Among the elastomers, polyethers are the least hydrophilic; therefore they are preferably chosen for moisture condition uses to capture preparation margins adequately; moreover, their wetting feature facilitates the fabrication of a gypsum cast. It can be prepared by the monophase impression technique or syringe-andtray method (Fig. 9.2) (Rafael and Liebermann, 2017; Livaditis, 1998). Polyether has some advantages, such as highly hydrophilic, high-to-moderate stability, good accuracy, impressible in a monophase transition state, available in wide range of viscosity, appropriate tear resistance, high modulus, and proper elastic recovery. On the other hand, it suffers from being too stiff, susceptible to moisture absorption, and with potential for allergic reactions (Von Fraunhofer, 2013). Cationic ring-opening polymerization of ethylene imine controls the polyether setting behavior (Sakaguchi and Powers, 2012). New emerging polyethers exhibit proper flexibility that is essential for its facile removal from the mouth. Because of water absorption characteristics, polyethers should not be fully immersed in the water to prevent distortion (Powers and Wataha, 2017). It was observed that exposing polyether to immersion disinfection deteriorates its wettability potential. Sodium hypochlorite and phenol enhance its wettability, while iodophor has the inverse effect (Shetty et al., 2013). Figure 9.2 Matrix filled with high-viscosity polyether impression material by the use of impression syringe to minimize entrapment of air (Livaditis, 1998).
Impression materials for dental prosthesis 201 9.2.2 Polysulfide Polysulfide, because of its nauseous taste, is not solely preferable for a dental prosthesis. It has been introduced as a paste to the two-component systems after mixing. Condensation polymerization helps the polysulfide setting process to yield a crosslinked rubbery polymer with high molecular weight, while together with polysulfide water forms as a by-product (Shoemaker et al., 2012; Guiraldo et al., 2017). Polysulfide enjoys the advantages of proper tear resistance, dimension stability, appropriate accuracy, and flexibility, whereas unpleasant taste, prolonged setting time, and mixing difficulties limit its capabilities (Levartovsky et al., 2011). Silica and TiO2 are being added as fillers to the base paste and have revealed promise in altering polysulfide viscosity. The set reaction consists of the oxidation of the SH groups that cause cross-linking and chain extension to yield elastomeric properties, which can be easily removed from the mouth, comparably easier than that of polyether (Hamalian et al., 2011; Yang et al., 2016). Its prolonged reaction results in a long-term dimensional change, while its setting time takes a value around 10 minutes with setting shrinkage (Von Fraunhofer, 2013). Due to the hydrophobic nature of polysulfide, it has been desirable to be utilized in milieu therapy without saliva and blood (Hamalian et al., 2011). 9.2.3 Alginate Alginate is an irreversible hydrocolloid resulting from a solgel transition caused by chemical reaction (Serrano-Aroca et al., 2017). Alginate is a polysaccharide presenting excellent biocompatibility with tissues (Atoufi et al., 2017). When exposed to calcium ions (Ca21), alginate aqueous solution cross-links. Sodium phosphate can be used as a retarder for regulating the setting time that has been found to vary from 1 to 5 minutes with a mild rise in water temperature (Fokkinga et al., 2017). Alginate is normally utilized when the accuracy is not so important. Elastomeric materials can be used as a secondary layer over the alginate. Thanks to its low price, easy flow, swift setting time, and minimal displacement, alginate can play the role of an IM. On the other hand, low-dimensional stability, inappropriate tear strength, and bubbling during mixing place some questions on alginate usage. The fraction of powder regulates the properties of the final product, such as gel strength, setting reaction, flow, and stability (Al-Enazi and Naik, 2016). It was reported that the gelation time is governed by the secondary materials such as disinfection liquids. For instance, chlorhexidine increases gelation time increment, while sodium hypochlorite decreases the gelation time. Alginate gelation is related to ion availability so that a higher content of ions results in fast gelation; moreover, ion concentration affects the mechanical stability (de Azevedo Cubas et al., 2014). Chlorhexidine will consume the ions and reduce the cross-link density (CLD); hence, the mechanical stability of alginate plunges down. On the other hand, sodium hypochlorite generates ions leading to the amelioration of the CLD and mechanical stability enhancement (Amalan et al., 2013). Operational conditions,
202 Advanced Dental Biomaterials such as temperature and humidity in storage and transport, manipulation of instructions, and ingredients ratio, govern the final product properties, such as distortion and stability (Kulkarni and Thombare, 2015). 9.2.4 Agar Agar is a kind of polysaccharide obtained from seaweed. It is a reversible hydrocolloids yielded via solgel transition under the influence of temperature, so that exhibits thermogelling behavior (Han et al., 2017; Zarrintaj et al., 2017a). Agar provides high accuracy that makes it effective for use in crowns and bridges (fixed prosthodontics); moreover, because of its thermoreversible behavior it can be used several times. Because of the hydrophilicity of agar, dental drying is not mandatory, and it can be used in the wet state. Agar formation necessitates the water bath and rim-lock trays with coiled edges, which can allow water to pass through to cool down the agar for the sake of setting (Iwasaki et al., 2016). Agar exhibits viscoelastic behavior, and its elastic recovery can be promoted by removing the impression quickly so that the material tolerates the stress for a concise time; furthermore, agar can be torn by applying a very low amount of stress because of poor mechanical properties. The dimensional stability of materials is not desirable due to high water uptake of gel (McCabe and Walls, 2013; Atoufi et al., 2017). Pouring the impression at several time intervals may bring about dimensional fluctuations. On this issue, sequential pouring of the IMs (alginate, agar, and polyvinyl siloxane) for several times on dimensional precision of the impressions was recommended. It was shown that when the materials were poured instantly, their dimensional exactness did not differ noticeably, while when the same materials were repoured after 30 minutes (the second pour), dimensional accuracy of alginate indicated the most alteration compared to the other aforementioned materials (Craig, 1988). 9.2.5 Silicones 9.2.5.1 Polysiloxanes Polysiloxane (PVS), usually called addition silicone, has been utilized widely in advanced restorative dentistry. Its grades are classified based on the filler content that controls its properties, such as thickness and flowability (Wang, 2016). The most well-known forms are extra light-bodied (low filler content), light-bodied, universal or medium-bodied, heavy-bodied, and putty (high filler content). A paste-topaste system and additional polymerization (without by-product) have been used for synthesizing such silicone leading to a stable production. However, the hydrophobicity of such material necessitates accurate moisture control in applying time (Goodall et al., 2015). Monophase polyvinyl siloxane and polyether elastomeric IMs have been widely studied. It was revealed that impregum, Penta, and aquasil acted better under the dry state; also, impregum acted better than the aquasil in both situations (Vadapalli et al., 2016). New hydrophilic elastomeric IMs have been recommended for diminishing the voids and distortion in the impressions. Soft
Impression materials for dental prosthesis 203 Figure 9.3 Polysiloxane (PVS) impression in a custom tray for fixed prosthodontics (Punj et al., 2017). polyether exhibited higher strain and lower tensile strength in comparison with addition silicones. Moreover, the tear and tensile strength of heavy-bodied materials were higher than those of light-bodies (Lu et al., 2004). Cole et al. synthesized thiol- and allyl-functionalized siloxane oligomers using radical-mediated polymerizations to gain swift set elastomeric dental IMs. Thiol-ene siloxane was crosslinked through the redox-initiated reaction. Properties of such dental impression were adjusted with plasticizer and kaolin filler; moreover, it exhibited a high accuracy (Cole et al., 2014). PVS should be cast in appropriate trays to use as an IM (Fig. 9.3) (Punj et al., 2017). 9.2.5.2 Condensation silicone The main materials of condensation silicones are dimethyl siloxane with CaCO3 or silica as a filler. Stannous octoate and alkyl silicate act as catalysts. Ethyl alcohol is the polymerization by-product that leads to high shrinkage during setting. Such silicones set quickly down on the tooth and are considered to get rigid to some extent (Von Fraunhofer, 2013). The condensation silicones are available in putty, paste, or light-bodied forms to be formed accurately. Hydrophobicity, shrinkage, and releasing their by-products, however, are disadvantages of condensation silicones to be taken into account before choosing them for dental treatments. Disinfectant within the silicone-based impression should be carefully controlled because it was reported that the dimensional stability of such products varies formulation to formulation. It is noteworthy that disinfectants containing benzalkonium chloride and glutaraldehyde are not deteriorative to the dimensional consistency of the aforementioned elastomeric materials (Sinobad et al., 2014).
204 Advanced Dental Biomaterials 9.2.5.3 Vinyl polyether siloxane This substrate is a novel IM receiving attention because of the simplicity of removal of additional silicone and the hydrophilicity of polyether simultaneously. Based on such specifications, it is a useful material for certain cases, such as narrow, deep gingival crevices (Punj et al., 2017). It was reported that VPS exhibited the appropriate dimensional stability (Nassar et al., 2013). Vinyl polyether silicone (VPES) and vinylpolysiloxane (VPS) were comparable regarding dimensional stability and surface detail reproduction after the disinfection process and long-time storage (Din et al., 2017). Investigations demonstrated that these materials are stable after 2-week storage; however, VPES showed less dimensional alteration compared to the VPS. Also, the disinfected species of both substrates were more consistent than the pristine ones. In vitro setting revealed that the VPES had enough dimensional stability and surface accuracy (Nassar and Chow, 2015). 9.3 Inelastic impression materials 9.3.1 Impression wax Varieties of natural waxes and resins have been utilized in dentistry for defined applications. Waxes refer to thermoplastic materials solidifying at ambient temperature and being molten without decomposition, which consist of two or more ingredients for the construction of nonmetallic denture bases (Powers and Craig, 1978). Dental waxes have a large thermal expansion coefficient. They can be expanded upon temperature rise and vice versa. The mechanical features of waxes depend inversely with temperature, but overall, their compressive strength and elastic modulus are poor. Waxes can be derived from natural sources, such as mineral (paraffin, microcrystalline, and montan), plant (carnauba, cocoa butter), insect (beeswax), animals (spermaceti), or from synthetic materials such as polyethylene, polyoxyethylene glycol, halogenated hydrocarbon, and hydrogenated (Tinto et al., 2017). Waxes, based on their usage, can be categorized into three groups in dentistry including pattern wax (inlay wax, casting wax, and base plate wax), processing wax (boxing wax, utility wax, and sticky wax), and impression wax (corrective wax and bite plate wax). The greatest disadvantage of wax is its distortion (Tinto et al., 2017). Conventional waxes consist of a paraffin wax with a low melting point and beeswax in a ratio of 3:1. This proportion assures that at mouth temperature, an appropriate flow is expectable. This kind of IM is not common in recording thorough impressions; in fact, they are used for the modification of small defects in other impressions, especially the zinc oxideeugenol impressions (Von Fraunhofer, 2013). Waxes belong to the thermoplastic substances that are able to flow at mouth temperature. These waxes can fill the sections of the impressions not receiving adequate material or imperfections caused by the air blow with the aid of brushing. Prior to utilizing the wax in defective
Impression materials for dental prosthesis 205 parts of the impressions or impression trays, it should be melted. It is necessary to give the material enough time when it is in the mouth to reach the oral temperature as it should undergo plastic flow to record the denture bearing area more precisely (McCabe and Walls, 2013). 9.3.2 Impression compound Impression compounds (ICs) are categorized as thermoplastic substances available in sheet or stick form. Such compounds have been produced by combining waxes, thermoplastic resins, fillers, and coloring agents. Some additives, such as shellac, stearic acid, and gutta-percha, can be added to the compound for plasticity enhancement. ICs softening takes a long time to complete because of low thermal conductivity of the ICs. It was observed that the shrinkage occurs after removal due to the high thermal expansion coefficient of ICs. Low thermal conductivity and high thermal expansion coefficient lead to the creation of internal stresses during the temperature fall form softening to the ambient degree resulting in distortion (Mete et al., 2017). Low-fusing substances have been used as IMs to aid flow at temperatures above 45 C and are available in sheet or stick form. In order to record the impressions of edentulous ridges, the sheet form is preferable, while the stick form is suitable for recording the impressions of single crowns. The method for softening the stick form is using a flame, whereas to soften the sheet a water bath is used. It is important to notice that during softening the stick material’s direct exposure to flame should be avoided to avoid ignition or boiling. Regarding the sheet forms, the softening time in water bath should be controlled, since if the submerged time is long, some important components, such as stearic acid, may be leached out. Higher fusing ICs can be applied for fabricating impression trays (McCabe and Walls, 2013; Von Fraunhofer, 2013; Anusavice et al., 2013). 9.3.3 Impression plaster Impression plasters (IPs) have conventionally been utilized as casting materials and IMs for edentate patients. The components of the IP are calcined, β-calcium sulfate hemihydrate that is mixed with water to trigger a reaction resulting in the formation of calcium sulfate dehydrate (Oppedisano, 2013). The ratio of water/powder affects mixture constancy and setting time. Controlling the behavior of IPs is an important factor; hence, some additives have been added to the compound. In order to decrease the setting expansion of the plasters, antiexpansion agents, such as potassium sulfate (K2SO4), have been utilized. It is reported that such agents speed up the setting reaction. Borax is also combined as a retarder to provide the dentist with an opportunity to control the setting features. Pigments, such as alizarin, can be utilized to make a difference between the impression and the mold (Dai et al., 2014). Conventional plaster casting materials, due to the adhesion to the mold, necessitate the use of releasing agents, but newly blended plaster is fabricated from acrylic resin or shellac to allow
206 Advanced Dental Biomaterials easy separation. The disinfection process of IPs involves immersion in sodium hypochlorite. These materials are fragile, and whenever the pressure or tensile force is applied, they fracture; thus, they cannot be used in undercut conditions. IPs have been used for recording impressions of highly mobile soft tissues overlying the residual alveolar bone. Moreover, IPs should be preserved in the dry state, as water absorbed from the humid state causes setting time extension (Von Fraunhofer, 2013). 9.3.4 Metallic oxide pastes (zinc oxideeugenol impression paste) Impression paste has been utilized to take the secondary impressions for a complete denture and consists of base part (zinc oxide) and catalyst paste (eugenol), which are mixed with a stainless steel spatula for about 1 minute. Zinc oxide is accompanied with vegetable or mineral oil, used for plasticizing, and also helps to neutralize the irritation caused by eugenol. Eugenol is usually available with rosin leading to smooth and homogeneous materials and also facilitating the reaction pace. The color of the zinc oxide paste is white, while eugenol paste is reddish brown. The color contrast helps in determining the perfect mixing, since in this state, the mixture should show a homogenous color (Manappallil, 2015). These substrates are divided into two main groups: hard paste (Type 1) for which the terminal set comes to pass within 10 minutes and soft paste (Type 2) that indicates a final set occurrence within 15 minutes. After the final set the impression can be removed from the mouth. The setting reaction is ionic in nature. Hence, ion concentration is an important factor in the reaction proceeding that can be affected by temperature and humidity, and ionizable salt also affects the reaction rate. Initially, ZnO hydrolyzes and reacts with eugenol to achieve zinc eugenolate salt (2C10H12O2 1 ZnO!Zn(C10H11O2)2 1 H2O) (Luengo et al., 2017). It was reported that an allergic response to the irritation can occur in some patients, as the eugenol leaches out and reaches the soft tissue. In such cases, ZOE-like materials, that is, eugenol-free zinc oxide impression pastes, are useful. ZOE-like materials are the products of the reaction between zinc oxide and different carboxylic acids, such as orthoethoxybenzoic acid. These acids are utilized instead of eugenol. In order to diminish the tissue burning sensation, oil of cloves (including 70%85% eugenol) can be used instead of pure eugenol in the first tube. Regarding disinfection, a 2% alkaline glutaraldehyde is utilized through the steps as previously stated for ICs. In the case of dimensional accuracy the impression pastes are reliable materials as less than 0.1% shrinkage happens during setting. The impression pastes can be maintained without any change in shape arising from the relaxation or another motive of the distortion. This state can be assured if the material used in the tray structure shows dimensional stability (Anusavice et al., 2013). All the mentioned materials should be used with the appropriate tray to exhibit the functional performance. 3Dprinted trays showed better properties, such as uniform thickness distribution of IMs, than conventional ones (Fig. 9.4) (Sun et al., 2017).
Impression materials for dental prosthesis 207 Figure 9.4 Design process of the digital 3D-printed tray [(AD) maxilla and (EH) mandible]. (A and E) The scanned data of the primary impression. (B) and (F) The impression is trimmed to the appropriate range. (C and G) The main part of the tray and the tissue stop. (D and H) A handle is added to the tray. The finished trays [(AiDi) manual tray and (EiHi) digital tray]. (Ai and Ei) A maxillary tray with a tissue stop. (Bi and Fi) A mandibular tray with a tissue stop. (Ci and Gi) A maxillary tray without a tissue stop. (Di and Hi) A mandibular tray without a tissue stop (Sun et al., 2017). 9.4 Characteristics of impression materials 9.4.1 Dimensional accuracy/dimensional stability Viscosity has a crucial role in determining the accuracy of the detail reproduction. In fact, a low viscosity or degree of pseudoplasticity helps to record the surface segments precisely (Hamalian et al., 2011). When fixing the IM into the patient’s mouth, it should be in the fluid state. The capability of the IMs to preserve the material accuracy over time indicates their dimensional stability; however, by the phrase dimensional accuracy, the absence of dimensional change for a short time after removing from the mouth and during setting is considered (Anusavice et al., 2013). The dimensional accuracy of some materials is time dependent, reported for the case of elastomeric IMs, including polyvinyl siloxane, polyether, and polysulfide. In other words the highest level of dimensional accuracy corresponds to the product collected right after the polymerization is completed, while it diminishes during the storage of IMs over a prolonged period (Rubel, 2007). Materials with low shrinkage have been chosen to be utilized as dental impressions (Hamalian
208 Advanced Dental Biomaterials et al., 2011). For instance, polyvinyl siloxane and polyether can keep their dimensional accuracy for about 12 weeks after making the impression (Rubel, 2007). PVS can be infinitesimally poured into the mold during the operation period, whereas since polyether can absorb humidity from the environment and swell, it is better for polyether to be poured within 1 hour after ejection from the mouth. Regarding polysulfide and condensation silicone, as by-products accompany their polymerization during the setting reaction, water and volatile ethyl alcohol should be poured within 30 minutes after ejection from the mouth because such byproducts volatilize from the set impression and cause distortion. All of the elastomeric IMs shrink during polymerization; moreover, because of the generation of by-products during the setting, they show more constriction. Overall, the greatest dimensional change during setting refers to the polysulfide and polyvinyl silicone, and the smallest one pertains to the PVS (Hamalian et al., 2011). 9.4.2 Wettability A hydrophilic nature is one of the essential characteristics an actual IM needs, as this substance is in continuity with the wet tissue. Affinity toward spreading on hydrophilic substances results in flow capability into tiny areas or splitting and recording partial details. IMs with low contact angle flow easily into small gaps and make impressions with fewer voids. Hence, such materials are reliable for utilization in fixed prosthodontics (Hamalian et al., 2011). The sort of IMs that can flow into partial segments in the scale of 2070 μm have been required in the field of fixed partial dentures; on the other hand, IMs that can reproduce details in the scale of 100150 μm are useful in the fields of removable prosthodontics (Rubel, 2007). Hydrophilic IMs enable dental stone to flow smoothly. Accordingly, the casts without bubbles can be obtained. In order to produce more precise casts when the IMs have high contact angle, both the particular pouring technique and attention should be considered. Hydrophobic materials necessitate the surfactant utilization to reduce the contact angle prior to pouring casts; on the contrary, polyether, polysulfide, and hydrocolloids are hydrophilic IMs with low contact angle (Rubel, 2007). 9.4.3 Elastic recovery/flexibility It is of paramount importance for IMs to have an elastic recovery property, allowing it to return to its main dimensions without noticeable distortion upon removal from the mouth (Re et al., 2015). It was indicated that PVS exhibited foremost elastic behavior (with over 99% elastic recovery), followed by polyether and then polysulfide. Instantly after mixing, PVS, because of swift elasticity development, should be applied promptly, particularly at high temperatures. Conversely, polyether kept the plasticity for the more extended duration after mixing. Also, the ultimate stiffness of the polyether is greater than that of PVS and may affect the ease of removing the material from the mouth (Mehta et al., 2014). The flexibility of the impressions can facilitate the removal of the materials from the mouth. Polyether is
Impression materials for dental prosthesis 209 the stiffest IM. Polyvinyl siloxane is moderately rigid, and the rigidity relies on the viscosity of the material (Hamalian et al., 2011). Alginate is evaluated as the most flexible IM (Rubel, 2007). The investigations demonstrated that the viscosity is an important factor in fabricating impressions and die with the least bubbles and maximal details. The increase in the amount of deformation and the time spent for removal of the impression from the mouth have an influence on the accuracy of the impression (Hamalian et al., 2011). 9.4.4 Mechanical properties The main three relevant mechanical features of IMs in clinical terms having a functional impact on dental impression applications are yield strength, strain at yield point, and tear energy. The yield strength is attributed to the point in which the material can tolerate the stress without permanent deformation. The strain at yield illustrates the quantity of undercut that an impression can defeat without permanent deformation (Re et al., 2015). The tear energy determines and shows the degree to which a material maintains resistance against tearing after setting. The material should provide some properties such as appropriate and sufficient elastic recovery and consumption of much energy to commence and spread tearing. Polysulfides are incredibly resistant to tearing but show permanent deformation and do not indicate complete elastic recovery after a critical point of permanent deformation (Al-Enazi and Naik, 2016). Hydrocolloids show low tear strength. Both PVS and polyether have the greatest tear strength, and they tear before they reach their perennial deformation point. Hence, they are more appropriate for clinical use (Hamalian et al., 2011). Since producers have their exclusive formulations, various viscosities and flow features exist. As different producers produce the materials, they also are at various working times available according to standard-set versus quick-set IMs (Rubel, 2007). 9.4.5 Miscellaneous In addition to all the general properties that have been mentioned in previous sections, some other criteria are essential, including evaluating whether the materials are tolerable for the patients, gaining the best consequences for low costs, and utilizing the disinfectants that cause the least dimensional changes. Disinfection of some materials, for example, hydrocolloids, polyethers, and methacrylates, need particular protocols to prevent distortion taking place after setting (Rubel, 2007; Zarrintaj et al., 2018b). Various antibacterials and disinfectants can be used for dental impressions (Hafshejani et al., 2017; Maller et al., 2012). Diluted sodium hypochlorite is a disinfectant but not a sterilizer, accepted by the American Dental Association for all materials except zinc oxideeugenol paste. Disinfection of zinc oxideeugenol impression paste is done using glutaraldehyde (Rubel, 2007). In Table 9.1 the properties of IMs are described.
Table 9.1 Comparison of various types of dental impression properties. Impression materials Type Advantage Polyether Elastic G G G G G G G Polysulfide Elastic G G G G Hydrophilic elastomeric Dimensionally stable Minimal shrinkage Proper accuracy Monophase impression Good tear resistance Low shrinkage Good tear resistance Stable dimensional Proper accuracy Most flexible elastomer Disadvantage G G G G G G G G Agar Elastic G G G Alginate Elastic G G G G G Addition silicone Elastic G G G G G High accuracy Hydrophilic Reusable Easy flow Cheap Reproduction of adequate detail Fast setting time Minimal tissue displacement in the mouth G Good detail reproduction Excellent dimensional stability No shrinkage on set High patient acceptance More than one model can be poured from one cast G G G G G G G G G G G Ref. Too stiff Short working time Guiraldo et al. (2017) Low patient satisfaction Unpleasant taste and odor Long setting time Requires excellent moisture control Difficult to mix Some shrinkage on set with the release of by-product Complex procedural steps Significant start-up cost of the hardware Punj et al. (2017) Poor dimensional stability Poor tear strength Unsupported Distortion Easy to include air during mixing A minimum thickness of 3 mm is required, which is hard to achieve in thin areas in between the teeth Hydrophobic Too accurate Poor tear resistance Expensive Khalid et al. (2015) Iwasaki et al. (2016) Punj et al. (2017) (Continued)
Table 9.1 (Continued) Impression materials Type Advantage Condensation silicone Elastic G G Accurate High patient acceptance Disadvantage G G G G G Plaster Inelastic G G G G G Impression compound Inelastic G G G G Zinc oxide eugenol plaster Inelastic G G G G Hydrophilic Good detail reproduction Excellent dimensional stability (contraction on setting) Good patient tolerance 23 min working time Primary impressions of complete dentures Border molding of trays Extension of trays Achieving mucocompression in the postdam area when working impressions are taken for complete dentures Thermoplastic Can be heated to aid removal from the casting material Good detail reproduction Excellent dimensional stability (0.15% shrinkage on setting) G G G Hydrophobic Requires excellent moisture control Unreliable dimensional stability Difficult to accurately proportion components leading to variable results Marked shrinkage on setting with the release of byproduct Brittle No recovery from deformation. Therefore if an undercut is present, the material will have to be broken off the impression and then glued back together before casting Excess salivation by the patient could have an adverse effect on detail reproduction Ref. Punj et al. (2017) Von Fraunhofer (2013) Von Fraunhofer (2013) G G Rigid Presence of undercuts can distort the final material or cause the section engaged to separate from the resultant impression Luengo et al. (2017)
212 9.5 Advanced Dental Biomaterials Conclusion and future perspective Dental IMs have been utilized for developing the numerous dental and orthodontic applications. They are the first stage in the chain of a dental prosthesis or the initial level of placing a crown or bridge. Due to such aims, various materials have been examined for dental prostheses purposes ranging from elastic to rigid materials. To preventing contamination, various disinfection materials have been added to the IMs to reach a desired product. The maintenance of properties, such as dimensional stability, is found to be among the major properties needed for the suitable functioning of IM. Various properties, such as setting time, accuracy, wettability, and recovery, should be considered in IM usage. IMs preparation necessitates the tray involved in IM casting. Recently, a digital technique has been utilized for the preparation of IMs and to optimize the process by shortening the process time, enhancing the accuracy, and facilitating the setting procedure. To reach the best dental impression, however, the material should be designed so as to meet several requirements simultaneously including low shrinkage, high accuracy, high stability, and proper hydrophilicity, and it must be applied with modern techniques to achieve the appropriate product. In this regard the current chapter gives some insights into the status of IMs used for dental prostheses. In the near future, digital dentistry will cover all aspects of this. The new emerging digital devices are starting to be used in the IM field, and they will become more user friendly, more precise, and smaller in size of wand/equipment. Tomography is used for capturing the basic graph to construct proper restorative implants. Hence, the digital methods in IMs preparation should be evaluated in more detail in future studies to design a road map for IMs selection for the next generation of restorative materials. References Al-Enazi, T.A., Naik, A.V., 2016. Disinfection of alginate and addition silicon rubber-based impression materials. Int. J. Stomatol. Occlusion Med. 8 (1), 4448. Amalan, A., Ginjupalli, K., Upadhya, N., 2013. Evaluation of properties of irreversible hydrocolloid impression materials mixed with disinfectant liquids. Dent. Res. J. 10 (1), 65. Anusavice, K.J., Shen, C., Rawls, H.R., 2013. Phillips’ Science of Dental Materials. Elsevier Health Sciences. Atoufi, Z., Zarrintaj, P., Motlagh, G.H., Amiri, A., Bagher, Z., Kamrava, S.K., 2017. A novel bio electro active alginate-aniline tetramer/agarose scaffold for tissue engineering: synthesis, characterization, drug release and cell culture study. J. Biomater. Sci., Polym. Ed. 143. just accepted. Bakhshandeh, B., Zarrintaj, P., Oftadeh, M.O., Keramati, F., Fouladiha, H., Sohrabi-jahromi, S., et al., 2017. Tissue engineering; strategies, tissues, and biomaterials. Biotechnol. Genet. Eng. Rev. 33 (2), 144172. Bhakta, S., Vere, J., Calder, I., Patel, R., 2011. Impressions in implant dentistry. Br. Dent. J. 211 (8), 361367.
Impression materials for dental prosthesis 213 Craig, R.G., 1988. Review of dental impression materials. Adv. Dent. Res. 2 (1), 5164. Chiu, L.L.Y., Chu, Z., Radisic, M., Mozafari, M., 2017. Tissue engineering. Reference Module in Materials Science and Materials Engineering. Elsevier. Cole, M.A., Jankousky, K.C., Bowman, C.N., 2014. Thiol-ene functionalized siloxanes for use as elastomeric dental impression materials. Dent. Mater. 30 (4), 449455. Dai, J., Cheng, N., Miron, R.J., Shi, B., Cheng, X., Zhang, Y., 2014. Effect of decreased implant healing time on bone (re)modeling adjacent to plateaued implants under functional loading in a dog model. Clin. Oral. Invest. 18 (1), 7786. de Azevedo Cubas, G.B., Valentini, F., Camacho, G.B., Leite, F.R.M., Cenci, M.S., PereiraCenci, T., 2014. Antibacterial efficacy and effect of chlorhexidine mixed with irreversible hydrocolloid for dental impressions: a randomized controlled trial. Int. J. Prosthodont. 27 (4), 363365. Din, S.U., Parker, S., Braden, M., Tomlins, P., Patel, M., 2017. Experimental hydrophilic vinyl polysiloxane (VPS) impression materials incorporating a novel surfactant compared with commercial VPS. Dent. Mater. 33, e301e309. Fokkinga, W.A., Witter, D.J., Bronkhorst, E.M., Creugers, N.H., 2017. Clinical fit of partial removable dental prostheses based on alginate or polyvinyl siloxane impressions. Int. J. Prosthodont. 30 (1), 3337. Goodall, R.H., Darras, L.P., Purnell, M.A., 2015. Accuracy and precision of silicon based impression media for quantitative areal texture analysis, Sci. Rep., 5. p. 10800. Guiraldo, R.D., Berger, S.B., Siqueira, R.M., Grandi, V.H., Lopes, M.B., Gonini-Júnior, A., et al., 2017. Surface detail reproduction and dimensional accuracy of molds: influence of disinfectant solutions and elastomeric impression materials. Acta Odontol. Latinoam. 30 (1), 1318. Hafshejani, T.M., Zamanian, A., Venugopal, J.R., Rezvani, Z., Sefat, F., Saeb, M.R., et al., 2017. Antibacterial glass-ionomer cement restorative materials: a critical review on the current status of extended release formulations. J. Control. Release 262, 317328. Hamalian, T.A., Nasr, E., Chidiac, J.J., 2011. Impression materials in fixed prosthodontics: influence of choice on clinical procedure. J. Prosthodont. 20 (2), 153160. Han, M., Li, Q.-L., Cao, Y., Fang, H., Xia, R., Zhang, Z.-H., 2017. In vivo remineralization of dentin using an agarose hydrogel biomimetic mineralization system. Sci. Rep. 7, 41955. Iwasaki, Y., Hiraguchi, H., Iwasaki, E., Yoneyama, T., 2016. Effects of immersion disinfection of agar-alginate combined impressions on the surface properties of stone casts. Dent. Mater. J. 35 (1), 4550. Khalid, M., Shah, S.N., Chughtai, M.A., 2015. Comparison of mean dimensional measurement of alginate impression using sodium hypochlorite versus gluteraldehyde and benzalkonium chloride for disinfection. Cell 321, 9001101. Kulkarni, M.M., Thombare, R.U., 2015. Dimensional changes of alginate dental impression materials—an in vitro study. J. Clin. Diagn. Res.: JCDR. 9 (8), ZC98ZC102. Levartovsky, S., Folkman, M., Alter, E., Pilo, R., 2011. Elastomeric impression materials. Refu’at ha-peh veha-shinayim (1993) 28 (2), 5464. Livaditis, G.J., 1998. Comparison of the new matrix system with traditional fixed prosthodontic impression procedures. J. Prosthet. Dent. 79 (2), 200207. Lu, H., Nguyen, B., Powers, J.M., 2004. Mechanical properties of 3 hydrophilic addition silicone and polyether elastomeric impression materials. J. Prosthet. Dent. 92 (2), 151154. Luengo, J., Reyes, H., Toscano, I., Garcia, Y., Anaya, M., 2017. Clinical and radiographic evaluation of CTZ (chloramphenicol-tetracycline-zinc eugenol oxide) antibiotic paste in pulp treatment of primary molars. J. Dent. Health Oral Disord. Ther. 8 (1), 00272.
214 Advanced Dental Biomaterials Maller, S.V., Karthik, K., Maller, U.S., Abraham, M.C., Kumar, R.N., Manikandan, R., 2012. Drug and dental impression materials. J. Pharm. Bioallied Sci. 4 (Suppl 2), S316. Manappallil, J.J., 2015. Basic Dental Materials. JP Medical Ltd. Marotti, J., Tortamano, P., Castilho, T.R., Steagall, W., Wolfart, S., Haselhuhn, K., 2014. Accuracy of a self-perforating impression tray for dental implants. J. Prosthet. Dent. 112 (4), 843848. McCabe, J.F., Walls, A.W., 2013. Applied Dental Materials. John Wiley & Sons. Mehta, R., Dahiya, A., Mahesh, G., Kumar, A., Wadhwa, S., Duggal, N., et al., 2014. Influence of delayed pours of addition silicone impressions on the dimensional accuracy of casts. J. Oral Health Commun. Dent. 8, 3. Mete, J.J., Rajguru, V.L., Dange, S.P., 2017. Aluminum barrier laminate or plastic tube as a dispenser for modeling plastic impression compound for border molding. J. Prosthet. Dent. 119, 676677. Nassar, U., Chow, A.K., 2015. Surface detail reproduction and effect of disinfectant and long-term storage on the dimensional stability of a novel vinyl polyether silicone impression material. J. Prosthodont. 24 (6), 494498. Nassar, U., Oko, A., Adeeb, S., El-Rich, M., Flores-Mir, C., 2013. An in vitro study on the dimensional stability of a vinyl polyether silicone impression material over a prolonged storage period. J. Prosthet. Dent. 109 (3), 172178. Oh, W., Morris, H., 2017. The principles of functional and mucostatic impressions for complete denture bases: a review. Compend. Continuing Educ. Dent. (Jamesburg, NJ: 1995) 38 (10), 664. Oppedisano, M., 2013. Delayed Linear Expansion of Two Ultra-low Expansion Dental Stones (Doctoral dissertation). Pastoret, M.-H., Krastl, G., Bühler, J., Weiger, R., Zitzmann, N.U., 2017. Accuracy of a separating foil impression using a novel polyolefin foil compared to a custom tray and a stock tray technique. J. Adv. Prosthodont. 9 (4), 287293. Powers, J.M., Craig, R.G., 1978. Thermal analysis of dental impression waxes. J. Dent. Res. 57 (1), 3741. Powers, J.M., Wataha, J.C., 2017. Dental Materials: Foundations and Applications. Elsevier. Punj, A., Bompolaki, D., Garaicoa, J., 2017. Dental impression materials and techniques. Dent. Clin. North Am. 61 (4), 779796. Rafael, C.F., Liebermann, A., 2017. Clinical characteristics of an allergic reaction to a polyether dental impression material. J. Prosthet. Dent. 117 (4), 470472. Re, D., De Angelis, F., Augusti, G., Augusti, D., Caputi, S., D’Amario, M., et al., 2015. Mechanical properties of elastomeric impression materials: an in vitro comparison. Int. J. Dent. 2015, 428286. Rubel, B.S., 2007. Impression materials: a comparative review of impression materials most commonly used in restorative dentistry. Dent. Clin. North Am. 51 (3), 629642. Sakaguchi, R.L., Powers, J.M., 2012. Craig’s Restorative Dental Materials-E-Book. Elsevier Health Sciences. Serrano-Aroca, Á., Ruiz-Pividal, J.-F., Llorens-Gámez, M., 2017. Enhancement of water diffusion and compression performance of crosslinked alginate films with a minuscule amount of graphene oxide. Sci. Rep. 7 (1), 11684. Shetty, S., Kamat, G., Shetty, R., 2013. Wettability changes in polyether impression materials subjected to immersion disinfection. Dent. Res. J. 10 (4), 539. Shoemaker, D.P., Chung, D.Y., Mitchell, J., Bray, T.H., Soderholm, L., Chupas, P.J., et al., 2012. Understanding fluxes as media for directed synthesis: in situ local structure of molten potassium polysulfides. J. Am. Chem. Soc. 134 (22), 94569463.
Impression materials for dental prosthesis 215 Sinobad, T., Obradović-Ðuričić, K., Nikolić, Z., Dodić, S., Lazić, V., Sinobad, V., et al., 2014. The effect of disinfectants on dimensional stability of addition and condensation silicone impressions. Vojnosanit. Pregl. 71 (3), 251258. Sun, Y., Chen, H., Li, H., Deng, K., Zhao, T., Wang, Y., et al., 2017. Clinical evaluation of final impressions from three-dimensional printed custom trays. Sci. Rep. 7 (1), 14958. Ting-shu, S., Jian, S., 2015. Intraoral digital impression technique: a review. J. Prosthodont. 24 (4), 313321. Tinto, W., Elufioye, T., Waxes, R.J., 2017. Pharmacognosy. Elsevier, pp. 443455. Tripathi,, A., Singh,, S.V., Aggarwal,, H., Gupta,, A., Arya,, D., 2017. A technique to preserve spacer thickness during custom tray fabrication with autopolymerizing polymethylmethacrylate. Int. J. Dent. Res. 2 (1), 12. Vadapalli, S.B., Atluri, K., Putcha, M.S., Kondreddi, S., Kumar, N.S., Tadi, D.P., 2016. Evaluation of surface detail reproduction, dimensional stability and gypsum compatibility of monophase polyvinyl-siloxane and polyether elastomeric impression materials under dry and moist conditions. J. Int. Soc. Prev. Commun. Dent. 6 (4), 302. Von Fraunhofer, J.A., 2013. Dental Materials at a Glance. John Wiley & Sons. Wang, Z., 2016. Comparison of dimensional accuracies using two elastomeric impression materials in casting three-dimensional tool marks. J. Forensic. Sci. 61 (3), 792797. Yang, Z.-Z., Wang, H.-Y., Lu, L., Wang, C., Zhong, X.-B., Wang, J.-G., et al., 2016. Hierarchical TiO2 spheres as highly efficient polysulfide host for lithium-sulfur batteries. Sci. Rep. 6, 22990. Zamanian, A., Yasaei, M., Ghaffari, M., Mozafari, M., 2013. Calcium hydroxide-modified zinc polycarboxylate dental cements. Ceram. Int. 39 (8), 95259532. Zarrintaj, P., Bakhshandeh, B., Rezaeian, I., Heshmatian, B., Ganjali, M.R., 2017a. A novel electroactive agarose-aniline pentamer platform as a potential candidate for neural tissue engineering. Sci. Rep. 7, 17187. Zarrintaj, P., Moghaddam, A.S., Manouchehri, S., Atoufi, Z., Amiri, A., Amirkhani, M.A., et al., 2017b. Can regenerative medicine and nanotechnology combine to heal wounds? The search for the ideal wound dressing. Nanomedicine 12 (19), 24032422. Zarrintaj, P., Manouchehri, S., Ahmadi, Z., Saeb, M.R., Urbanska, A.M., Kaplan, D.L., et al., 2018a. Agarose-based biomaterials for tissue engineering. Carbohydr. Polym. 187, 6684. Zarrintaj, P., Urbanska, A., Gholizadeh, S.S., Goodarzi, V., Saeb, M.R., Mozafari, M., 2018b. A facile route to the synthesis of anilinic electroactive colloidal hydrogels for neural tissue engineering applications. J. Colloid Interface Sci. 516, 5766.
10 Nano glass ionomer cement: modification for biodental applications Shariq Najeeb1, Zohaib Khurshid2, Hani Ghabbani3, Muhammad S. Zafar3,4 and Farshid Sefat5,6 1 Independent Researcher and Private Practitioner, Alberta, Canada, 2Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia, 3Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia, 4Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan, 5Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom, 6Interdisciplinary Research Centre in Polymer Science and Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom Chapter Outline 10.1 Introduction 217 10.2 Applications of glass ionomer cements 219 10.3 Nanomodifications of glass ionomer cement powders 10.3.1 10.3.2 10.3.3 10.3.4 219 Powder-based nanomodification of glass ionomer cements 220 Nanohydroxyapatite and ionomers 220 Glass ionomer cements modified with other nanoparticles 222 Nanomodified resin-modified glass ionomer cements 223 10.4 Conclusion 224 References 224 10.1 Introduction A variety of synthetic and natural biomaterials are used to restore diseased or damaged tissues (Zafar et al., 2015; Husain et al., 2017). Materials such as metals and alloys have been used in dentistry to repair teeth for thousands of years. Historical records indicated that Plaster of Paris might have been used as a bone substitute material in as early as the 19th century. Alloys of silver, called dental amalgams, have been used since the late 1800s. More recently, a variety of materials, such as titanium, zirconia, and polyetheretherketone (Najeeb et al., 2016b,c), have been used as dental implant materials. Dental materials are used to replace or restore hard and soft oral tissues. There are two main types of dental materials: direct Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00010-4 Copyright © 2019 Elsevier Ltd. All rights reserved.
218 Advanced Dental Biomaterials dental materials and indirect dental materials. Direct dental materials are used directly in the oral cavity to restore dental tissues; they do not need major processing or modification in the dental laboratory before use. Some examples of direct dental materials are resin composites, glass ionomer cement (GIC), compomer, fissure sealants, and topical fluoride. Indirect dental materials have been modified, physically and/or chemically, in the dental laboratory or chair-side before being used in the oral cavity. Examples of indirect materials are acrylic dentures, dental implants, indirect resin composite crowns and veneers, and maxillofacial prostheses. Indeed, modern dentistry involves the combined use of these indirect and direct materials, along with natural biomaterials to improve the functions, esthetics, and biomechanics of oral tissues. GICs are mainly used as direct restorative dental materials. Directly speaking, GICs are a group of restorative materials which contain aluminofluorosilicate glasses as the vital component. First developed by Wilson and Kent in the early 1970s (Wilson and Kent, 1971), GICs have come a long way since their inception. A number of modifications in their composition, introduction of various bioactive fillers, and changing the particle size of the powders have been combined to tailor the physical, chemical, and optical properties of GICs. Conventionally, GICs have been composed of two main components: a fluoroaluminosilicate glass and an aqueous solution of polyacids (Sakaguchi and Powers, 2012). In addition to the main ingredients, other acids such as maleic acid and itaconic acid (IA) may be added to the formulation to improve the handling properties by altering the setting time. Incorporation of salts of radiopaque metals such as barium and strontium makes GICs visible in radiographs, making a clinical diagnosis and posttreatment assessment of restorations easier. The presence of fluoride ions in the glass component of GICs may impart antibacterial properties by releasing fluoride into the oral cavity (Zafar and Ahmed, 2015; Zafar, 2013). Therefore fluoride-releasing restorative materials are used as vehicles for delivering fluoride to the oral cavity (Zafar and Ahmed, 2015; Ullah and Zafar, 2015). Indeed, the inherent adhesive qualities of GICs, coupled with their white color and translucency, make them an acceptable material for restoring deciduous teeth. Conventional glass ionomers are set via a reaction between polyacrylic acid (PAA) and the glass particles. The reaction initially results in the gelation of the acid and glass components (Sakaguchi and Powers, 2012). The structure can be best described as unreacted glass particles embedded in the silica gel. Indeed, changing the particle size, the type of the acid and adding other additives to the glass and/or acid component of the cement can alter the physical, mechanical, and chemical properties. The hardening phase of the GIC takes place when the polyacid chains cross-link with calcium and aluminum ions in the powder (Anusavice and Phillips, 2003). Research has indicated that the majority of sodium and fluoride ions remain unreacted in the set GIC. The final stage of GIC setting is called “maturation.” During this stage, which may last for up to 1 year, water is bound by the acid and glasses. Additionally, aluminum may also slowly leach out. In addition to dental applications, research is being conducted to make GICs suitable for bone applications. Their bioactive and adhesive properties are an advantage which may
Nano glass ionomer cement: modification for biodental applications 219 prove to be useful in bone regeneration and implant applications. However, there are several limitations such as poor surface mechanical properties and wear resistance (Zafar and Ahmed, 2014; Zafar, 2018), that may impede the use of GICs. In this chapter, we discuss various modifications that have been made to GICs to improve their biological and mechanical properties. 10.2 Applications of glass ionomer cements Over the last several decades, GICs have seen a variety of applications in dentistry. GICs have been used to restore deciduous (milk) teeth due to ease of application. Because of their white color and translucency, GICs have also been used to restore anterior (front) teeth, either alone or in combination with other tooth-colored materials, such as resin composites. Due to their adhesive properties, they may also be used for the temporary or permanent cementation of crowns and bridges, in addition to luting of orthodontic devices. Indeed, their bioactive properties have also allowed them to be placed in deep carious lesions, following removal of infected dentine. Because infected dentine is removed, but bacteria-free and caries-affected dentine is left as much as possible, this technique is termed as atraumatic restorative treatment. Due to its special dentine bonding properties, glass ionomers can be used in endodontic treatment (obturation) as a sealer and as a core material. Activ GP (gutta percha) contains gutta percha cones saturated with GIC on the external surface. Although the sealer has minimal antimicrobial activity compared to other sealers (Heling and Chandler, 1996), it creates a monoblock and allows adhesion with canal walls. Also, obturating the root canal with a glass ionomer based sealer makes the root more resistant to vertical root fracture because of its ability of dentin bonding (Trope and Ray, 1992). 10.3 Nanomodifications of glass ionomer cement powders Nanotechnology is best described as the use of systems and modifications of materials in the size range of 1 100 nm to improve their mechanical properties, function, and esthetic attributes (Zafar et al., 2016, 2017; Khurshid et al., 2015). Over the past few years, nanotechnology has been used to enhance osseointegration of implants (Najeeb et al., 2015), reinforcement of polymeric materials (Najeeb et al., 2016d), prevention of dental caries (Huang et al., 2011), and to enhance the efficacy of tissue regeneration and drug delivery (Qasim et al., 2017). Two main techniques have been employed for the production of nanosized materials. “Top-down” is the production of particles or systems in the nanometer range by removal of bulk material, leaving behind nanometer-sized particles. Milling, lithography, and machining are examples of top-down manufacturing. Conversely,
220 Advanced Dental Biomaterials “bottom-up” nanomanufacturing would be best described as fabrication of nanometer-sized materials atom by atom, such as in the case of the coating of dental implants, tissue regeneration, and additive 3D printing. Nanomodification of GICs is mainly produced by the top-down nanofabrication of glasses, oxides, and bioceramics. A composite material is a material composed of two or more materials with vastly different chemical and physical properties. Teeth and bone are an example of natural composites. They both are composites of inorganic apatites embedded in an organic matrix. In dentistry and surgery, a wide variety of synthetic and natural composite materials are used. Resin composites, dental implants, and regenerative tissue materials are just some of several composites used. Indeed, technically speaking, GICs are also composite materials. Unreacted glass particles are embedded in a cross-linked matrix of organic acids, much akin to bone and teeth. Hence, modification of the particle size and chemistry has a significant effect on the mechanical, physical, and optical properties of GICs. Several nanomodifications have been proposed to improve the quality, life, and function of GICs. This chapter aims to summarize these modifications and ascertain their present status and the clinical use of these nanomodified GIC formulations. 10.3.1 Powder-based nanomodification of glass ionomer cements Nanomodification of GICs has been achieved in three ways. First, the size of the glass powder itself is altered to change the mechanical and physical properties of GICs. Indeed, the initial study by De Caluwe et al. (2014) demonstrated the effect of introducing nanosized glass powder on the setting characteristics and mechanical properties of GIC. They observed that the addition of nanosized fluoroaluminosilicate glass not only decreased the setting time but also improved the mechanical properties of the set GIC. A number of procedures are used to replicate the detrimental effect of the oral environment on the restorative materials. It has been observed that thermocycling has a more negative impact on the mechanical properties of nanomodified materials than those of conventional ionomers. Over the past few years, several modifications have been suggested, aimed at increasing the mechanical, physical, biological, and optical properties of GICs. Nanomodification of GICs mainly involves the incorporation of nanosized fillers to the powder component. Significant nanomodifications are presented below, and a summary of the mechanical properties of various nanomodified GICs are presented in Table 10.1. 10.3.2 Nanohydroxyapatite and ionomers Hydroxyapatite (HAp) is found in bone and teeth as the mineral component. Due to its osteoconductive effects and remineralizing attributes, HAp has been used as a dental implant coating to prevent dental caries and in regenerative therapy
Table 10.1 A summary of the mechanical properties of nanomodified glass ionomer cements (GICs) compared with those of conventional GICs. Liquid Powder Nanoparticles (% and size) Compressive strength (MPa) Tensile strength (MPa) Flexural strength (MPa) Reference PAA copolymer PAA copolymer FAS FAS 161 178 11.8 19 14.8 31 Moshaverinia et al. (2008a) Moshaverinia et al. (2008b) PAA copolymer FAS 179 23 33 Moshaverinia et al. (2008b) Polymer of AA, NVP, IA PAA copolymer FAS 183.8 23.5 36 Moshaverinia et al. (2008a) FAS None nHAp (5 wt.% 100 200 nm) nFAp (5 wt.% 100 200 nm) HAp (5 wt.% 100 200 nm) 3% 23.17 PAA copolymer FAS Garcia-Contreras et al. (2015); Elsaka et al. (2011) Kumar et al. (2017) Chitosan (10 wt.% 110 235 nm) 176.27 21.26 FAS, Fluoroaluminosilicate; HAp, hydroxyapatite; IA, itaconic acid; nHAp, nanohydroxyapatite; NVP, N-vinylpyrrolidone; PAA, polyacrylic acid.
222 Advanced Dental Biomaterials (Huang et al., 2011; Kettenberger et al., 2017). Utilizing their bioactive nature, both HAp and fluorapatite (FAp) have been incorporated into resin composites and glass ionomers (Moshaverinia et al., 2008b; Zakir et al., 2013). Indeed, the addition of nano-HAp (nHAp) and nano-FAp (nFAp) to the GIC powder results in significantly improved flexural, tensile, and compressive strengths of the set GIC (Moshaverinia et al., 2008b). There are several mechanisms by which apatites may improve the mechanical properties of the GICs. Spectroscopic studies have revealed that incorporation of apatites decreases the solubility of the set GIC by increasing the crystallinity of the composite (Moshaverinia et al., 2008a,b). When compared with nHAp-modified GICs, nFAp-modified GICs have demonstrated higher crystallinity leading to better mechanical properties (Moshaverinia et al., 2008a,b). Fractures and sensitivity to moisture are two of the most common causes of their failure. Hence, the addition of nanoapatites to GICs may overcome their limitations and prolong their intraoral life. The mechanical properties of nFAp-modified GICs can be further enhanced by substituting the PAA copolymer with a mixture of IA and N-vinylpyrrolidone (NVP) copolymers. The mechanism behind this improvement is the interaction between the apatite and the NVP chains, resulting in a more crystalline, stronger chemical structure (Moshaverinia et al., 2008a). Another advantage of incorporating nanoapatites is the increased adhesive bond strength between the tooth and the GIC restoration. This is thought to be a result of two observations. First, the chemical similarity of apatites present in teeth and the nanoapatites incorporated in the GIC may give rise to the ionic interactions between the tooth and the GIC structures (Lucas et al., 2003). Second, the small size of the nanometer-sized apatite crystals makes it easier for them to infiltrate dentinal tubules and enamel pores (Lee et al., 2010). 10.3.3 Glass ionomer cements modified with other nanoparticles Zirconia (zirconium oxide, ZrO2) is a ceramic which has been used in dentistry for the production of crowns (Kolbeck et al., 2008), dentures, and dental implants (Calvo-Guirado et al., 2015). A study by Gu et al. (2005) evaluated the effect on the mechanical properties of incorporating nanosized HAp and zirconia (nZrO2) to conventional GIC. In the study, it was observed that the addition of nHAp and nZrO2 at a concentration of 4% by volume resulted in significant improvement in the mechanical properties of the cement. Nevertheless, in the same study, when scanning electron micrographs of conventional set GIC and set nHAp/nZrO2modified GIC were compared, more cracks were observed in the latter (Gu et al., 2005). This could have a detrimental effect on the long-term viability of these modified cements, because the cracks may lead to marginal failure and subsequent failure of the restorations. Other nanosized particles, such as calcium fluoride (CaF2) and titania (titanium oxide; TiO2), have also been used to modify GICs (Elsaka et al., 2011; Moreau and Xu, 2010). Addition of CaF2 to resin-modified GICs (RMGICs) may prolong the longevity of these cements by improving the mechanical properties, but it may also have an adverse effect on fluoride release from the cement (Xu et al., 2010).
Nano glass ionomer cement: modification for biodental applications 223 Although nano-TiO2 (nTiO2) has been observed to improve the mechanical properties and enhance the antibacterial attributes of GICs, more studies are required to determine the safety of these modified restorations, because titania is known to be cytotoxic (Hall et al., 2009). Nanoparticles of chitosan, a deacylated form of chitin, a polysaccharide found in shells of crustaceans, have been observed to improve the flexural strength of GICs significantly due to increased interaction between the chitosan chains and the GIC matrix (Kumar et al., 2017). 10.3.4 Nanomodified resin-modified glass ionomer cements To improve the mechanical properties of GICs, they may be combined with a self-curing or light-activated resin polymer. These RMGICs are less susceptible to flexural fractures and expansion due to water sorption and have better esthetics (McCabe, 1998). However, RMGICs still have some disadvantages. First, they have inferior mechanical properties when compared to resin composites (McCabe and Walls, 2008). Moreover, they have higher creep and lower fluoride release when compared to conventional GICs. More recently, nanoclusters of silica have been incorporated to the powder formulation of the RMGICs to improve the mechanical properties and fluoride release from the set cement (Coutinho et al., 2009). They are supplied with a primer that aids bonding with tooth structure. RMGICs bond to the tooth structure through a combination of micromechanical interactions due to the infiltration of resins into the etched tooth surface and ionic bonding between the HAp crystals present in tooth and the acids RMGICs (Fig. 10.1). The aim of adding Figure 10.1 The schematic presentation of the structure of nanomodified GIC and its bonding with dental hard tissues (Najeeb et al., 2016a). GIC, Glass ionomer cement.
224 Advanced Dental Biomaterials nanofillers to the RMGICs is to increase the infiltration of resin into the pores in the etched tooth surface and enhance the binding. However, studies indicate that there is no significant difference between the tensile bond strength of nanoRMGICs and conventional GIC (Coutinho et al., 2009). Indeed, clinical studies also showed that there is no difference between the survival rates of RMGICs and nano-RMGICs (Perdigao et al., 2012). Although nano-RMGIC is available commercially (e.g., Ketac N100/Ketac Nano; 3M ESPE, St. Paul, Minnesota, United States), more research is needed before nano-RMGICs find more use in the clinical setting. Abrasion produced by toothbrush simulation produces less surface roughness on the surface of nano-RMGICs than on conventional RMGICs; degradation produced by bacteria and chemicals produce comparable surface roughnesses on both (de Paula et al., 2011, 2014; de Fúcio et al., 2012). This indicates that in long-term use, nano-RMGICs do not hold any significant advantage over conventional RMGICs as far as the surface properties are concerned. To date, no statistically significant difference has been observed between the fluoride release from nano-RMGICs and conventional RMGICs. 10.4 Conclusion Although some nanomodifications have been proposed for GICs, more studies and long-term clinical trials are vital to ascertain their advantages over conventional GICs and RMGICs. The failure of the bioceramic glass interface in the set GICs is thought to be a major concern that should be studied more. Not many studies have been carried out to observe the effect of nanomodified GICs on dental pulp cells. However, with the recent improvements in the resins and fillers, nanomodified GICs may find widespread usage in dental clinics. References Anusavice, K., Phillips, R., 2003. Phillips’ Science of Dental Materials. Saunders, St. Louis, MO, Great Britain. Calvo-Guirado, J.L., Aguilar Salvatierra, A., Gargallo-Albiol, J., Delgado-Ruiz, R.A., Maté Sanchez, J.E., Satorres-Nieto, M., 2015. Zirconia with laser-modified microgrooved surface vs. titanium implants covered with melatonin stimulates bone formation. Experimental study in tibia rabbits. Clin. Oral Implants Res. 26, 1421 1429. Coutinho, E., Cardoso, M.V., De Munck, J., Neves, A.A., Van Landuyt, K.L., Poitevin, A., et al., 2009. Bonding effectiveness and interfacial characterization of a nano-filled resinmodified glass-ionomer. Dent. Mater. 25, 1347 1357. De Caluwe, T., Vercruysse, C.W., Fraeyman, S., Verbeeck, R.M., 2014. The influence of particle size and fluorine content of aluminosilicate glass on the glass ionomer cement properties. Dent. Mater. 30, 1029 1038.
Nano glass ionomer cement: modification for biodental applications 225 de Fúcio, S.B., de Paula, A.B., de Carvalho, F.G., Feitosa, V.P., Ambrosano, G.M., Puppin-Rontani, R., 2012. Biomechanical degradation of the nano-filled resin-modified glass-ionomer surface. Am. J. Dent. 25, 315 320. de Paula, A.B., Fucio, S.B., Ambrosano, G.M., Alonso, R.C., Sardi, J.C., Puppin-Rontani, R., 2011. Biodegradation and abrasive wear of nano restorative materials. Oper. Dent. 36, 670 677. de Paula, A.B., de Fucio, S.B., Alonso, R.C., Ambrosano, G.M., Puppin-Rontani, R., 2014. Influence of chemical degradation on the surface properties of nano restorative materials. Oper. Dent. 39, E109 E117. Elsaka, S.E., Hamouda, I.M., Swain, M.V., 2011. Titanium dioxide nanoparticles addition to a conventional glass-ionomer restorative: influence on physical and antibacterial properties. J. Dent. 39, 589 598. Garcia-Contreras, R., Scougall-Vilchis, R., Contreras-Bulnes, R., Sakagami, H., MoralesLuckie, R., Nakajima, H., 2015. Mechanical, antibacterial and bond strength properties of nano-titanium-enriched glass ionomer cement. J. Appl. Oral Sci. 23, 321 328. Gu, Y.W., Yap, A.U., Cheang, P., Khor, K.A., 2005. Effects of incorporation of HA/ZrO(2) into glass ionomer cement (GIC). Biomaterials 26, 713 720. Hall, S., Bradley, T., Moore, J.T., Kuykindall, T., Minella, L., 2009. Acute and chronic toxicity of nano-scale TiO2 particles to freshwater fish, cladocerans, and green algae, and effects of organic and inorganic substrate on TiO2 toxicity. Nanotoxicology 3, 91 97. Heling, I., Chandler, N.P., 1996. The antimicrobial effect within dentinal tubules of four root canal sealers. J. Endod. 22, 257 259. Huang, S., Gao, S., Cheng, L., Yu, H., 2011. Remineralization potential of nanohydroxyapatite on initial enamel lesions: an in vitro study. Caries Res. 45, 460 468. Husain, S., Al-Samadani, K.H., Najeeb, S., Zafar, M.S., Khurshid, Z., Zohaib, S., et al., 2017. Chitosan biomaterials for current and potential dental applications. Materials 10, 602. Kettenberger, U., Luginbuehl, V., Procter, P., Pioletti, D.P., 2017. In vitro and in vivo investigation of bisphosphonate-loaded hydroxyapatite particles for peri-implant bone augmentation. J. Tissue Eng. Regen. Med. 11, 1974 1985. Khurshid, Z., Zafar, M., Qasim, S., Shahab, S., Naseem, M., AbuReqaiba, A., 2015. Advances in nanotechnology for restorative dentistry. Materials 8, 717 731. Kolbeck, C., Behr, M., Rosentritt, M., Handel, G., 2008. Fracture force of tooth tooth-and implant tooth-supported all-ceramic fixed partial dentures using titanium vs. customised zirconia implant abutments. Clin. Oral Implants Res. 19, 1049 1053. Kumar, R.S., Ravikumar, N., Kavitha, S., Mahalaxmi, S., Jayasree, R., Kumar, T.S., et al., 2017. Nanochitosan modified glass ionomer cement with enhanced mechanical properties and fluoride release. Int. J. Biol. Macromol. 104, 1860 1865. Lee, J., Lee, Y., Choi, B., Lee, J., Choi, H., Son, H., et al., 2010. Physical properties of resin-reinforced glass ionomer cement modified with micro and nano-hydroxyapatite. J. Nanosci. Nanotechnol. 10, 5270 5276. Lucas, M.E., Arita, K., Nishino, M., 2003. Toughness, bonding and fluoride-release properties of hydroxyapatite-added glass ionomer cement. Biomaterials 24, 3787 3794. McCabe, J.F., 1998. Resin-modified glass-ionomers. Biomaterials 19, 521 527. McCabe, J.F., Walls, A.W.G., 2008. Applied Dental Materials. Oxford, Blackwell. Moreau, J.L., Xu, H.H., 2010. Pmc2948642; fluoride releasing restorative materials: effects of pH on mechanical properties and ion release. Dent. Mater. 26, e227 e235.
226 Advanced Dental Biomaterials Moshaverinia, A., Ansari, S., Movasaghi, Z., Billington, R.W., Darr, J.A., Rehman, I.U., 2008a. Modification of conventional glass-ionomer cements with N-vinylpyrrolidone containing polyacids, nano-hydroxy and fluoroapatite to improve mechanical properties. Dent. Mater. 24, 1381 1390. Moshaverinia, A., Ansari, S., Moshaverinia, M., Roohpour, N., Darr, J.A., Rehman, I., 2008b. Effects of incorporation of hydroxyapatite and fluoroapatite nanobioceramics into conventional glass ionomer cements (GIC). Acta Biomater. 4, 432 440. Najeeb, S., Khurshid, Z., Matinlinna, J.P., Siddiqui, F., Nassani, M.Z., Baroudi, K., 2015. Nanomodified peek dental implants: bioactive composites and surface modification—a review. Int. J. Dent 2015, 381759. Najeeb, S., Khurshid, Z., Zafar, M.S., Khan, A.S., Zohaib, S., Martı́, J.M.N., et al., 2016a. Modifications in glass ionomer cements: nano-sized fillers and bioactive nanoceramics. Int. J. Mol. Sci. 17, 1134. Najeeb, S., Khurshid, Z., Zohaib, S., Zafar, M.S., 2016b. Bioactivity and osseointegration of PEEK are inferior to those of titanium—a systematic review. J. Oral Implantol. 42, 512 516. Najeeb, S., Zafar, M.S., Khurshid, Z., Siddiqui, F., 2016c. Applications of polyetheretherketone (PEEK) in oral implantology and prosthodontics. J. Prosthodont. Res. 60, 12 19. Najeeb, S., Khurshid, Z., Agwan, A.S., Zafar, M.S., Alrahabi, M., Qasim, S.B., et al., 2016d. Dental applications of nanodiamonds. Sci. Adv. Mater. 8, 2064 2070. Perdigao, J., Dutra-Correa, M., Saraceni, S.H., Ciaramicoli, M.T., Kiyan, V.H., 2012. Randomized clinical trial of two resin-modified glass ionomer materials: 1-year results. Oper. Dent. 37, 591 601. Qasim, S.B., Najeeb, S., Delaine-Smith, R.M., Rawlinson, A., Rehman, I.U., 2017. Potential of electrospun chitosan fibers as a surface layer in functionally graded GTR membrane for periodontal regeneration. Dent. Mater. 33, 71 83. Sakaguchi, R.L., Powers, J.M., 2012. Craig’s Restorative Dental Materials. Elsevier/Mosby, Philadelphia, PA. Trope, M., Ray, H.L., 1992. Resistance to fracture of endodontically treated roots. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. 73, 99 102. Ullah, R., Zafar, M.S., 2015. Oral and dental delivery of fluoride: a review. Fluoride 48, 195 204. Wilson, A.D., Kent, B.E., 1971. The glass-ionomer cement, a new translucent dental filling material. J. Appl. Chem. Biotechnol. 21, 313. Xu, H.H., Moreau, J.L., Sun, L., Chow, L.C., 2010. Pmc3077944; novel CaF(2) nanocomposite with high strength and fluoride ion release. J. Dent. Res. 89, 739 745. Zafar, M.S., 2013. Effects of surface pre-reacted glass particles on fluoride release of dental restorative materials. World Appl. Sci. J. 28, 457 462. Zafar, M.S., 2018. Wear behavior of various dental restorative materials. Mater. Technol. 33, 25 31. Zafar, M.S., Ahmed, N., 2014. Effects of wear on hardness and stiffness of restorative dental materials. Life Sci. J. 11, 11 18. Zafar, M.S., Ahmed, N., 2015. Therapeutic roles of fluoride released from restorative dental materials. Fluoride 48, 184 194. Zafar, M.S., Khurshid, Z., Almas, K., 2015. Oral tissue engineering progress and challenges. Tissue Eng. Regener. Med. 12, 387 397. Zafar, M.S., Najeeb, S., Khurshid, Z., Vazirzadeh, M., Zohaib, S., Najeeb, B., et al., 2016. Potential of electrospun nanofibers for biomedical and dental applications. Materials 9, 73.
Nano glass ionomer cement: modification for biodental applications 227 Zafar, M.S., Khurshid, Z., Najeeb, S., Zohaib, S., Rehman, I.U., 2017. Chapter 26— Therapeutic applications of nanotechnology in dentistry. In: Andronescu, E., Grumezescu, A.M. (Eds.), Nanostructures for Oral Medicine. Elsevier. Zakir, M., Al Kheraif, A.A., Asif, M., Wong, F.S., Rehman, I.U., 2013. A comparison of the mechanical properties of a modified silorane based dental composite with those of commercially available composite material. Dent. Mater. 29, e53 e59.
Enamel etching and dental adhesives 11 Ahmed Talal1, Hafiz Muhammad Owais Nasim2 and Abdul Samad Khan1 1 Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia, 2Department of Dental Materials, Sharif Medical and Dental College, Lahore, Pakistan Chapter Outline 11.1 Introduction 229 11.2 Indications of adhesives 230 11.3 Composition of adhesives 231 11.3.1 Etchant 231 11.3.2 Primer 231 11.3.3 Bonding 231 11.4 Types of enamel etching 232 11.4.1 Acid etching 232 11.4.2 Laser etching 235 11.4.3 Self-etching 236 11.5 Classifications of adhesives 237 11.5.1 Classification based on generations 237 11.5.2 Classification based on clinical steps 240 11.5.3 Classification based on interaction with smear layer 241 11.6 Dentin bonding 241 11.7 Advancement in adhesives 242 11.7.1 Antibacterial properties 242 11.7.2 Bioactive properties 245 11.8 Conclusion 247 References 248 11.1 Introduction The development of adhesives was a turning point in dentistry that revolutionized restorative and preventive dentistry. The concept of cavity preparation changed with the use of adhesives, and it was no longer required to extend the cavity preparation for resistance and retention of form at the cost of sound tooth structure. In addition, advancements in adhesives substantially reduced recurrent caries which is a major cause of restorations failure (Sofan et al., 2017a; Van Meerbeek et al., 1992). Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00011-6 Copyright © 2019 Elsevier Ltd. All rights reserved.
230 Advanced Dental Biomaterials The mineralized tooth structure, that is, enamel and dentin, has a unique composition and morphology. Enamel consists of hydroxyapatite (HA) (96 wt.%), water (3 wt.%), and organic matrix, that is, proteins and lipids (1 wt.%) and has a solid homogeneous crystalline structure with strong intermolecular forces and high surface energy which makes bonding relatively easy and predictable (Baier, 1992), whereas dentin consists of HA (70 wt.%; 40% 45% in vol.), water (10 wt.%; 20% 25% in vol.), and organic matrix (20 wt.%; 30% in vol.) and presents a challenge to achieve optimal adhesion due to its chemical composition, level of moisture, and morphological variation. In addition, morphological and biomechanical changes which occur in the dentin due to the physiological and pathological processes can make the adhesion process complex and challenging (Vinagre and Ramos, 2016). The success or failure of resin restorations is substantially dependent on the quality of adhesives which in turn is crucially dependent on the correct isolation of the operative site from the fluid (dentinal fluid and saliva) and water. The main goals of adhesion are (1) restoration retention and stability; (2) absorption and counteraction of shrinkage stresses; (3) pulp dentin complex sealing; (4) postoperative sensitivity reduction; (5) tooth structure reinforcement; and (6) reduction or elimination of microleakage (Barkmeier and Cooley, 1992; Lutz et al., 1996). Adhesives used in restorations perform the following functions (Brenna, 2012; Bowen and Marjenhoff, 1992): G G G Long-term durability of the restoration by providing a perfect seal Functional rehabilitation of the tooth by distributing forces Provide retention to the restoration and counter shrinkage stresses In the last three decades, much progress has been made in dental adhesives with substantial improvement in their properties and ease of application. However, still, bond failure and recurrent caries at the restoration tooth interface constitute the major cause of restoration replacement. New developments in dental adhesives are focused on imparting antibacterial and bioactive properties to improve their clinical performance further. 11.2 Indications of adhesives With the advancement in adhesives and development of restorative materials, the use of adhesives has widened substantially, and they are being used in numerous dental procedures, listed as follows (Perdigão and Sezinando, 2013): 1. 2. 3. 4. 5. 6. restoration of carious and noncarious defects; bonding of indirect restorations (crowns, inlay, and onlay); bonding of partial and full resin veneers in diastema closure; pit and fissure sealant; orthodontic bracket cementation;
Enamel etching and dental adhesives 231 7. desensitizing of the exposed tooth root; 8. sealing of root canal; and 9. bonding of fiber or metal posts. 11.3 Composition of adhesives The main components of adhesive systems include the one described in the following subsections (Sofan et al., 2017a; Erickson, 1992). 11.3.1 Etchant The etchant used in adhesive systems demineralizes the tooth structure and produces microporosities in which primer and bonding flow to produce micromechanical bonding. The etchant used in etch-and-rinse systems is 35% 37% phosphoric acid. Whereas in self-etch systems an acidic monomer acts as both an etchant and a primer. 11.3.2 Primer Hydrophilic monomers in water-soluble solvents (acetone, ethanol, and water) for improved flow and penetration into porous tooth substrate are used as primers. These primers increase the wettability of adhesives on the etch tooth surface. 11.3.3 Bonding Unfilled fluid resins have been used as a bonding agent between the tooth and composite and are mostly light-cured. The basic mechanism behind the bonding of dental restorations with the dentin and enamel is by the exchange of minerals or ions from the dental hard tissue with resin of the restoration. To fulfill this concept, etching is the first line of action. Enamel etching technique is used to remove the smear layer and create a rough surface. The “smear layer” term was first introduced by McComb and Smith (1975). According to them, an amorphous smear layer was produced due to the instrumentation during the cutting and shaping of enamel and dentin. This smear layer consists of inorganic and organic parts of the dental tissue. The bonding between the restoration and the dentin is compromised due to the presence of the smear layer, which is 0.5 5.0 μm thick and has the ability to block the dentinal tubules. On the other hand, it reduces the permeability of the dentin and acts as a “diffusion barrier.” Some studies showed that the presence of a smear layer might reduce or inhibit the bacterial growth due to reduced dentin permeability. A study showed that with the application of sodium hypochlorite (NaOCl) and mild hypochlorous acid (HOCl) solutions, the organic component on the smear layer-covered dentin could be removed, which could eliminate the hybridized smear layer created by self-etch adhesive, leading to the reduction of nanoleakage within
232 Advanced Dental Biomaterials the hybrid layer (HL) (Thanatvarakorn et al., 2014). Therefore the main objective of enamel etching is to remove the smear layer and produce a rough surface by dissolving the enamel rods or prisms. It encourages the adhesive to flow between the indentations and provide the micromechanical retention for the resin-based composite materials and restorations. However, there is no scientific consensus on the efficacy of the removal of the smear layer in the root canal treatment (Moszner et al., 2005). 11.4 Types of enamel etching Mainly, enamel etching has been divided into the following: G G acid etching and laser etching. 11.4.1 Acid etching The acid etching technique was first presented by Buonocore in 1955 for etching the enamel surface of the tooth (Buonocore et al., 1968). According to his hypothesis, industrial phosphoric acid, which was used in paints and acrylic coating for metals, can be used to adhere the acrylic resin with the enamel of the tooth after etching with 85% phosphoric acid for 30 seconds. The acid etching technique can remove the superficial layer of enamel up to 10 μm and produce a rough porous surface, where it is expected that the average pore size is 5 50 μm, which helps in the formation of “resin tags” and also increases the wettability of the enamel surface (Devarasa et al., 2012). It is important to know the pattern of acid etching of enamel. There are three major types of pattern produced by acid etching (Gwinnett, 1994), which are as follows: G G G Type 1: Removal of the enamel prism core, the peripheral prism remains intact. Type 2: Reverse of the type 1 process; the peripheries of the cores are removed, leaving the enamel prism core intact. Type 3: A somewhat total-etch technique in which prism morphology is less distinctive. Furthermore, acid etching increases the bond strength between the restoration and the dental hard tissue. It also reduces the microleakage around the restoration and reduces the chances of secondary caries. However, a drawback associated with etching is that the outermost strongest layer of the enamel surface is lost, which makes it more prone to acid attacks. Acid etching procedure needs to be further simplified, with a reduction in the steps and the amount of enamel loss should be reduced, as the conservative approach is appreciated in the modern restorative dentistry (Peumans et al., 2005). To etch the enamel surface, many types of inorganic acids and chelators have been used; however, it is important to determine which acid treatments would yield favorable results consistently on tooth surfaces. Among
Enamel etching and dental adhesives 233 the many types of etchants that have been used clinically, phosphoric acid-based etchants have gained popularity. 11.4.1.1 Phosphoric acid In the etch-and-rinse technique, 30% 40% phosphoric acid is applied on the tooth surface for 30 60 seconds and then rinsed off. A range of concentrations of the phosphoric acid have been used in the etching process. By using 37% phosphoric acid the prismatic and interprismatic crystal on the enamel surface is dissolved and a rough and porous surface is formed, which may be 5 50 μm in depth. If the phosphoric acid concentration is less than 27%, then it forms dicalcium phosphate dihydrate precipitate which is difficult to remove. In contrast, if the concentration of the phosphoric acid is increased to more than 40% it reduces the dissolution of the calcium which results in the shallow etching pattern (Silverstone, 1974). Therefore commercially available etchant contains phosphoric acid in a 37% concentration. The recommended time for acid etching is 15 20 seconds; however, a comparative study showed that the 15 seconds etching provided similar surface roughness as compared to 60 seconds. However, a shorter etching time resulted in less bonding strength. It is important to know that the acid-treated surfaces are more prone to caries and the strength of the enamel may reduce (Swift et al., 1995). Another application of the 37% phosphoric acid is to apply it onto the tooth surface for 60 seconds before bleaching. It helps in better penetration of the bleaching agent into the enamel surface and gives better results. It is established that there are morphological and histological differences between the primary and permanent teeth; therefore the infiltration of resins is reduced in primary dentition due to aprismatic enamels. It is reported that no significant differences in the micromorphology of the enamel surface of primary teeth were found for etching times ranging from 15 to 120 seconds (Garcia-Godoy and Gwinnett, 1991). Another study reported no significant difference in bond strength for primary enamel when etched for 60 240 seconds (Smutka et al., 1978). However, Bozalis et al. (1979) suggested that primary enamel should not be etched for more than 180 seconds. 11.4.1.2 Citric acid Citric acid is considered as a weak organic acid and is used in periodontal infection or diseases after the flap surgeries or on instrumentation to recover the cementum and reduce the healing time. It has been applied as a conditioner and weak etchant and it is reported that the use of 50% citric acid is effective and significantly removes the surface smear layer (Prado et al., 2011). It can also be used as an irrigating material for the root canal treatment, where the best results can be obtained by using citric acid of pH 1.1 to remove the smear layer (Haznedaroğlu, 2003). Another study showed that the most effective pH of citric acid used for the complete removal of smear layer is between pH 1.1 and 1.9 (Di Lenarda et al., 2000). The advantages the citric acid are (1) antibacterial properties; (2) cytotoxic effect;
234 Advanced Dental Biomaterials Figure 11.1 SEM of dentin surface etched with 50 wt.% citric acid at a pH of 0.7, the acid has produced a porous and clean dentinal surface (Breschi et al., 2002). and (3) removal of the smear layer (Hennequin and Douillard, 1995). Fig. 11.1 shows the SEM image of dentinal tubules after etching with 50% citric acid of pH 0.7 and it exhibits good etching of the smear layer with the production of porous and clean intertubular dentin. Other alternative materials for acid etching are 10% melic acid and 2.5% nitric acid which are as effective as the above. Melic acid is used for conditioning dentin and enamel surfaces. 11.4.1.3 Ferric chloride solution Though the effect of ferric chloride is not completely understood, 3% ferric chloride solution has been used for acid etching before applying the adhesive and the placement of the restoration. It is suggested that the use of ferric chloride may have the following advantages (Saeki et al., 2001): G G G G It can promote polymerization (Bowen, 1980). It may help in the cross-linking and stabilizing the etch collagen network. It protects the demineralized matrix from degradation during the drying procedure. It helps in the denaturation of the collagen matrix during the process of etching (Mizunuma, 1986). Amalgambond (Parkell Dental, United States) was first developed by Nakabayashi et al. (1982) and used in the dentin bonding system containing 10% citric acid and 3% ferric chloride. Amalgambond has the ability to remove the smear layer over the dentin and produce the rough surface up to 2 μm in depth. It can also be used to bond the amalgam with the dentin. The shear bond strength of this bond system is less than or equal to 20 MPa (Barbour and Rees, 2004).
Enamel etching and dental adhesives 235 11.4.1.4 Sodium hypochlorite The first bleaching agent proposed was sodium hypochlorite which has been used as a microabrasive material to remove the superficial layer of the enamel. It has been used to remove superficial stains on the tooth surface which do not extend to the deeper layer of the enamel. In the case of moderate to severe fluorosis, the conservative treatment of surficial discoloration includes treatment by using sodium hypochlorite alone or with the bleaching agent. However, it is not an alternative to the definitive restorative management including veneers, which are placed by reducing the enamel surface. This sodium hypochlorite etching technique can be useful in treating young patients who are more concerned about the esthetic and color of teeth. The sodium hypochlorite solution has been used extensively to remove or clean the organic structure by oxidizing it. The hypomineralized enamel lesion can be degraded, and the chromogenic organic materials present on the enamel surface can be removed by applying the sodium hypochlorite. It is recommended to apply 5% sodium hypochlorite to the superficial layer of the enamel for 5 10 minutes with the bleaching agent, if the staining or discoloration persists then reetch for 60 seconds and rinse. 11.4.1.5 Ethylenediamine tetra-acetic acid In older patients, the mineral content, tertiary dentin and/or “sclerotic dentin,” become resistant to acid etching techniques. At the microscopic level, the dentinal tubules are constricted due to the calcification and thickening of peripheral dentin. For complete removal of the smear layer, a mixture of hypochlorite, chelating agents (ethylene diamine tetra-acetic acid (EDTA) or quaternary ammonium bromide (cetrimide) added EDTA), and acids (maleic, citric, and polyacrylic acid) have been used to dissolve the organic and inorganic part of the dental hard tissue. Sodium hypochlorite and EDTA can be used as an alternative for the irrigation in root canal treatment. It is proven that the EDTA completely removes the smear layer if applied for 2 3 minutes on the tooth structure. The main drawback of the EDTA is prolonged working time for etching or endodontic treatment (Prado et al., 2011). Moreover the EDTA with aqueous citric acid can also be used to irrigate the root canal in endodontic treatment (Yamaguchi et al., 1996). 11.4.2 Laser etching Due to the limitation of the acid etch techniques, “laser irradiations” can be used to remove the smear layer after preparation of the tooth cavity. One of the alternatives is the erbium, chromium:yttrium scandium gallium garnet (Er,Cr:YSGG) radiation from the laser which creates many small pores or a rough surface by melting the superficial layer of the enamel and recrystallizing the enamel rods or prisms. The pattern obtained after the laser etching is somewhat similar to that seen in the type 3 acid etching technique (Ghaffari et al., 2017). The recrystallization of the
236 Advanced Dental Biomaterials Figure 11.2 Er,Cr:YSGG laser-irradiated dentin surface showing scaly, irregular, and rugged appearance (Tseng et al., 2007). Er,Cr:YSGG, Erbium, chromium: yttrium scandium gallium garnet. enamel surface shows the fungiform appearance of the enamel surface, which sufficiently increases the micromechanical or chemical retention of the restoration with the tooth structure. Fig. 11.2 shows the Er,Cr:YSGG laser-irradiated dentin surface with a scaly, irregular, and rugged appearance (Tseng et al., 2007). Due to these properties, it may be a good alternative to the conventional acid etching technique because it does not induce any pain, heat, or vibration (Lee et al., 2003). The results of bond strength between the restoration and tooth structure by laser etching are still controversial. 11.4.3 Self-etching This is the new technique which contains acidic monomer in the polymerizable form and has a dual function as conditioner and primer for enamel and dentin (Moszner et al., 2005). By using the self-etching technique, the number of steps is reduced, as rinsing of phosphoric acid from enamel and dentin is not required. Its action is similar to the conventional etching and rinse technique to remove the smear layer from the dental hard tissue, however, in comparison, the enamel surface is less retentive, as shown in Fig. 11.3. Its application involves two steps: first application of self-etching primer (SEP) on dentin and enamel followed by adhesive resin application. These SEPs are known as sixth generation (Kugel and Ferrari, 2000). In recent advancements, one-bottle and one-step self-etching have been introduced to the market, which is known as seventh generation (Cal-Neto and Miguel, 2006). Methacrylate-based SEP or bonding agents having the pH 1.5 2.5 are currently used. There are some limitations or drawbacks in the one component self-etching, that is, it tends to absorb more water and hydrolytically degrade the dimethacrylate-based resins.
Enamel etching and dental adhesives 237 Figure 11.3 Enamel surface after etching with a self-etching primer solution; the enamel surface is less retentive than that obtained with phosphoric acid (Kugel and Ferrari, 2000). 11.5 Classifications of adhesives Dental adhesives can be classified based on various factors, which include the following: G G G different generations as they developed; number of clinical steps involved; and their action on the smear layer. 11.5.1 Classification based on generations 11.5.1.1 First-generation adhesives Developed in the 1960s by Bowen, Cervident (S White Burs, Inc, Lakewood, New Jersey) was the first-generation dentin bonding agent. It consisted of N-phenylglycine-glycidyl methacrylate comonomer. It was claimed to have the ability to chelate with calcium in the tooth structure. However, the clinical results were not very promising with a bond strength of only 2 3 MPa (Kugel and Ferrari, 2000; Sofan et al., 2017a; Barkmeier and Cooley, 1992). 11.5.1.2 Second-generation adhesives Clearfil Bond System F introduced in 1978 was the first product of secondgeneration dentin adhesives. It was suggested that second-generation systems make bonds to the dentin surface through an ionic bond between calcium and chlorophosphate groups. Although the produced bond was stronger than that of the first generation, however, the bond strength of second-generation dentin bonding agent was 1 5 MPa which was still below the 10 MPa required for acceptable in vivo retention. The bond produced by the second generation was prone to hydrolysis when
238 Advanced Dental Biomaterials exposed to saliva or moisture in the dentin, which could result in debonding of restoration. In addition, as no etching was involved in earlier generations, the adhesion was due to bonding to the smear layer which was weak and unreliable (Kugel and Ferrari, 2000; Barkmeier and Cooley, 1992). 11.5.1.3 Third-generation adhesives In 1979 Fusayama et al. introduced acid etching of dentin before application of bonding agent. However, acid etching did not produce much improvement in the dentin bonding due to the hydrophobic nature of bonding resin. In addition, it was thought that acid etching might produce pulpal inflammation, which limited its use. Most of the third-generation bonding systems developed at that time did not remove the entire smear layer but modified it. These materials modify the smear layer and slightly demineralize the underlying intertubular dentin. Third-generation adhesives produced both poor and good results. Scotchbond 2 (3M ESPE, Germany) was the first material of this kind which was fully accepted by American Dental Association (ADA) (Kugel and Ferrari, 2000; Heymann et al., 2013). 11.5.1.4 Fourth-generation adhesives Fourth-generation dentin adhesive systems were developed to be used on acidetched dentin and consisted of etchant, primer, and bonding, which were applied to the tooth one after another. Acid application on dentin removes the smear layer, demineralizes intertubular and peritubular dentin, opens the dentin tubules, and exposes the collagen fiber network (Fig. 11.4). Acid etching changes the mineral content of dentin substrate and reduces its surface free energy. As dissolved HA has high surface energy and the exposed collagen network has low energy surface, the primer used in these systems increases the surface energy of dentin. The bonding mechanism between the adhesive material and etched dentin substrate is achieved by the formation of resin tags and adhesive lateral branches in the dentin porosities. Figure 11.4 Scanning electron micrograph of etched dentin showing exposed collagen fibers and open dental tubules (Pashley et al., 2011).
Enamel etching and dental adhesives 239 Figure 11.5 SEM image of the resin dentin interface, showing RC, the AR, HL, adhesive resin tags (T), and dentin (D) (Mortazavi et al., 2012). AR, Adhesive resin; HL, hybrid layer; RC, resin composite. Primer and bonding resin penetrate etched intertubular dentin producing a resin dentin interdiffusion zone or HL (Fig. 11.5) (Mortazavi et al., 2012). Depending on the acid type, concentration, and application time, dentin is demineralized up to 7.5 μm. These systems produce bond strength between 17 and 30 MPa. These adhesives systems are however time-consuming and technique-sensitive. As in the fourth-generation adhesives, enamel and dentin are etched simultaneously; these systems were initially called the “total-etch” technique, now they are commonly known as the “three-step etch-and-rinse” technique (Kugel and Ferrari, 2000). 11.5.1.5 Fifth-generation adhesives Fifth-generation adhesives are two-step etch-and-rinse systems. They simplify the bonding procedure by combining the primer and bonding agent in a single solution and are applied simultaneously. They are also called “one-bottle” systems. However, a separate etching step is still required. They have an acceptable bond strength, are less technique-sensitive, easy to use, and have reduced postoperative sensitivity (Sofan et al., 2017a). 11.5.1.6 Sixth-generation adhesives The sixth-generation adhesive systems are two-steps self-etch systems. Two-steps self-etch systems employ SEPs which do not require rinsing. These systems are also described as “nonrinsing conditioners” or “self-priming etchant.” In the SEP system, etching and priming of enamel and dentin are done simultaneously. Since the rinsing and drying step is eliminated in SEP, the chances of over- or underdrying of dentin are reduced. These systems are less technique sensitive than etch-andrinse systems. There are less chances of discrepancy between demineralization and resin infiltration depth as both happen simultaneously. In addition, there may be less chances of postoperative sensitivity with SEP as compared to an etch-and-rinse
240 Advanced Dental Biomaterials system, as SEPs do not eradicate the smear layer. Although these adhesives produce good bond strength to dentin, the bond strength to enamel is comparatively less than that produced by fourth- and fifth-generation adhesives. 11.5.1.7 Seventh-generation adhesives In seventh-generation systems, etching, priming, and bonding have been combined in a single solution and as a single step. These are also called “one-step self-etch” or “all-in-one” adhesive systems. Although the clinical procedure is simplified in the seventh-generation systems, their highly hydrophilic nature makes them liable to water degradation and the bond strength of seventh-generation adhesives is lowest among all the adhesives systems (Sofan et al., 2017b). 11.5.1.8 Eighth-generation adhesives Futurabond DC (Voco, United States) containing nano-sized fillers were introduced in 2010 by Voco America as eighth-generation adhesives systems. These adhesives systems, owing to the presence of nanofillers, have higher penetration into the etched dentin and produce a thicker HL with improved mechanical properties. They have increased bond strength to enamel and dentin and better stress absorption. As the filler acts as a cross-linking agent, these systems have better dimensional stability (Sofan et al., 2017a). However, a study showed that increasing the nanofiller content beyond 1 wt.% increases the viscosity of the adhesive which may cause them to accumulate as clusters which can lead to cracks and bond disintegration (Kasraei et al., 2009). 11.5.2 Classification based on clinical steps Dental adhesives can also be classified based on the number of clinical steps involved. They are classified as either one-, two-, or three-step adhesive systems. 11.5.2.1 Three-step etch-and-rinse systems These adhesive systems involve three steps, which include etching, priming, and bonding, all of which done separately. These adhesives systems are supplied in three bottles, each for etchant, bonding, and primer. Among all the adhesive systems, three-step etch-and-rinse systems produce the most durable bonding and are considered the gold standard (Van Meerbeek et al., 2003a). 11.5.2.2 Two-step systems Two-step adhesive systems simplified the bonding procedure by combining two steps. These systems are further subdivided as follows: G Two-step etch-and-rinse system. In these systems the priming and bonding steps are combined, as primer and bonding are supplied together in one bottle which is applied after etching and rinsing.
Enamel etching and dental adhesives G 241 Two-step self-etch systems are also known as “self-priming etchant.” These systems combine etching and priming steps, which are followed by the application of bonding without rinsing. 11.5.2.3 One-step self-etch system One-step adhesive systems combine etching, priming, and bonding steps. SEP and bonding are supplied as one solution, supplied in a single bottle. These systems further simplify the bonding process (Sofan et al., 2017a). 11.5.3 Classification based on interaction with smear layer Mechanically prepared dentin is covered by loosely bound smear layer, for a reliable bond between restorative resin and dentin; this smear layer must either be removed or modified (Asmussen and Uno, 1992). Adhesive systems can be classified according to their effect on the smear layer (Brenna, 2012). 11.5.3.1 Smear layer removing systems The adhesive systems which involve etch-and-rinse technique can be regarded as smear layer removing systems. They are either three-step systems (etching 1 primer 1 bonding) or two-step systems (etching, primer 1 bonding). These systems involve phosphoric acid etching of tooth substrates (enamel and dentin), which is followed by rinsing with water for complete removal of acid. Acid application with subsequent rinsing removes the smear layer and exposes surface roughness, enamel prisms, dentinal tubules, and collagen network. 11.5.3.2 Smear layer dissolving systems The adhesive systems that incorporate SEPs can be regarded as smear layer dissolving systems. These systems include two-step self-etch systems and one-step selfetching “all-in-one” systems which do not involve rinsing after application. These adhesive systems do not remove the smear layer completely, rather they dissolve the smear layer by infiltration. 11.6 Dentin bonding Dentin is composed of HA, water, and organic matrix. It has a heterogeneous structure with low intermolecular forces and low surface energy. In addition, it is structurally and chemically altered by pathological processes, all this and the wet surface due to fluid seepage make the bonding between dentin and restorative resin complicated. Moreover, the tubular build-up of dentin and the resulting outward pulpal water current in vital teeth turn dentin into a complex substrate (Baier, 1992; Erickson, 1992; Van Meerbeek et al., 2003b; Perdigão, 2010). Tooth structure cutting during cavity preparation results in the formation of the smear layer which is loosely attached to the tooth surface. This smear layer is
242 Advanced Dental Biomaterials composed of organic and inorganic components from enamel, dentin, or cementum, which in clinical conditions is usually contaminated with saliva, blood, and bacteria. It covers the dentin surface and blocks the dentinal tubules by forming smear plugs (Tao et al., 1988). Two strategies are generally employed to address the presence of the smear layer; these include complete removal of the smear layer before bonding or the use of bonding agents which can penetrate the smear layer while incorporating it. However, the removal of the smear layer and smear plugs increases the dentinal tubules permeability, allowing fluid flow from and to the pulp chamber (Pashley and Carvalho, 1997; Van Meerbeek et al., 2003b). Bond strength of resin monomer to caries-affected dentin is lower than in sound dentin. The HL produced in caries-affected dentin is thicker than that produced in normal dentin, as affected dentin is more susceptible to acid, which leads to deeper penetration of adhesive resins, whereas the penetration of adhesive resin into the dentinal tubules is obstructed by acid-resistant mineral deposits in the dentinal tubules. Lack of resin tag hybridization to peritubular dentin, due to acid-resistant mineral deposits, decreases the cohesive strength of caries-affected dentin and leads to lower bond strength to affected dentin (Say et al., 2005; Wei et al., 2008). Studies showed that dentin aging does not significantly affect the bond strength between dentin and adhesive resins (Lopes et al., 2011; Ozer et al., 2005). The inherently wet nature of dentin makes it difficult to clinically dry it completely. Excessive drying of dentin can lead to the collapse of the collagen network which prevents the infiltration of adhesive resin into the nanochannels produced due to HA dissolution between collagen fibers. Subsequently, it results in reduced bonding between the resin and dentin. This led to the development of the “moist bonding” technique in which dental adhesives can be used on a moist dentin surface. These adhesives incorporate organic solvents such as acetone and ethanol with primers or adhesives. These solvents enhance the penetration of resin monomers into the dentin collagen network, as they displace water from the moist dentin surface and intact collagen network, which results in enhancing bond strength (Kanca, 1992; Perdigao, 2002). 11.7 Advancement in adhesives 11.7.1 Antibacterial properties Secondary caries is the major cause of restoration failures and replacements. Bacterial invasion between the restoration and tooth structure, due to microleakage or the residual bacteria in the cavity, can lead to secondary caries formation (Kim and Shin, 2014). It is desirable to have an adhesive system with antibacterial properties which can prevent colonization of bacteria and inhibit caries, resulting in increased longevity of the restoration (Amin et al., 2014; Imazato et al., 2014; Chen et al., 2014). Various studies investigated the antibacterial potential of adhesive systems by incorporating nanoparticles of silver (NAg), nanoparticles of amorphous calcium phosphate (NACP), fluorine ions, zinc oxide, quaternary ammonium
Enamel etching and dental adhesives 243 dimethacrylate (QADM), 12-methacryloyloxydodecyl pyridinium bromide (MDPB), benzalkonium chloride, and chlorhexidine (Cheng et al., 2015; Zhang et al., 2013; Melo et al., 2013; Passariello et al., 2014; Andre et al., 2015; Wiegand et al., 2007). However, the antibacterial activity of silver particles and fluorine ions decreases over time and may adversely affect the material’s physical properties. In contrast, another study showed that the silver nanoparticles could have positive effects on bond strength of both etch-and-rinse and self-etch adhesive systems. It is further reported that the best results of silver nanoparticles are achieved if they are applied before acid etching (Fatemeh et al., 2017). In a study, 0.1 wt.% NAg was added to the primer which was produced with pyromellitic glycerol dimethacrylate and 2-hydroxyethyl methacrylate, whereas the adhesive consisted of bisphenol-A-glycerolate dimethacrylate and triethylene glycol dimethacrylate, in which 0.1% NAg, 10% QADM, and 0% 40% NACP were added. The results showed that incorporation of NAg, QADM, and NACP into the bonding system greatly reduced human saliva microcosm biofilm viability, metabolic activity, lactic acid production, and colony-forming units of microcosm biofilms (Melo et al., 2013). It is believed that the bactericidal action of quaternary ammonium (QA) is due to three processes: (1) interaction between negatively charged bacterial cell and positively charged QA which results in the increase in osmotic pressure; (2) diffusion of QA through the cell wall and binding to the cytoplasmic membrane; and (3) cytoplasmic membrane disruption which leads to release of cytoplasmic constituents and cell death (Cocco et al., 2015; Han et al., 2017). The proposed mechanism of antibacterial action of QA is given in Fig. 11.6. In another study, quaternary ammonium monomer dimethylaminododecyl methacrylate (DMADDM) and NACP were incorporated into commercially available bonding agent (Scotchbond Multi-Purpose; 3M, St. Paul, Minnesota) to investigate Figure 11.6 Proposed mechanism of antibacterial action of QA (Cocco et al., 2015). QA, Quarternary ammonium.
244 Advanced Dental Biomaterials its antibacterial potential, the results showed that DMADDM and NACP substantially reduced metabolic activity, lactic acid formation, and the total streptococci and Streptococcus mutans of the biofilm (Chen et al., 2014). Another in vitro study also investigated the incorporation of 2-methacryloyloxyethyl phosphorylcholine, dimethylaminohexadecyl methacrylate (DMAHDM), and NACP into Scotchbond Multi-Purpose bonding agent for root dentin coating to prevent caries with promising antibacterial and remineralizations results (Zhang et al., 2015b). A recent study showed that DMAHDM and NACP added into a bonding for root caries application has a potent effect on periodontal and endodontic pathogens (Wang et al., 2016). In an in vitro study, Passariello et al. (2014) compared antibacterial properties of multiple bonding agents by placing disks samples in culture media seeded with Streptococcus gordonii, Streptococcus sanguinis, S. mutans, or Lactobacillus acidophilus. The study showed that the bonding agents containing MDPB, benzalkonium chloride, and chlorhexidine have the highest antibacterial activities for the longest duration of time among all the materials assessed. Another study compared the antibacterial properties of three commercially available adhesives systems with different antibacterial components. The adhesive systems compared were Gluma 2Bond (Glutaraldehyde), Clearfil SE Protect (MDPB), and Peak Universal Bond (Chlorhexidine). These adhesives systems were applied on the dentin surface of extracted teeth which were then restored with composites. The antibacterial activity of these adhesives was evaluated by direct contact method against four strict anaerobic and four facultative bacteria for up to 24 hours. The results showed that among the adhesive systems, MDPB coating adhesive systems have the highest antibacterial activity (Andre et al., 2015). Clearfil Protect Bond (Kuraray Medical, Okayama, Japan), an antibacterial adhesive containing MDPB, has bactericidal properties in uncured form and when applied into the cavity acts as a disinfectant. Once cured it inhibits the growth of the bacteria by destroying their cell membranes which come in direct contact with the restoration. When applied on the dentin surface Clearfil Protect Bond can penetrate to a depth of 500 μm. One of the major causes of secondary dentin is residual bacteria. Since MDPB-containing adhesive system has cavity disinfecting properties, it may improve the long-term prognosis of the restoration (Andre et al., 2015; Imazato et al., 2014). In a clinical study, Clearfil Protect Bond used for orthodontic bracket cementation showed reduced enamel demineralization around orthodontic brackets compared to a conventional adhesive (Transbond XT, 3M Unitek, Monrovia, California) during a 30-day period (Uysal et al., 2010). It is reported that MDBP can be used to improve bonding durability by inhibiting matrix metalloproteases and cathepsins (Tezvergil-Mutluay et al., 2011, 2015). Li et al. (2009) incorporated QA groups, that is, methacryloxylethylcetyl dimethyl ammonium chloride, into the dental adhesive and observed that the modified adhesive influenced the growth, adherence, and membrane integrity of S. mutans. Jedrychowski et al. (1983) reported that the small concentrations of chlorhexidine increased the antibacterial activity without compromising the mechanical properties. Su et al. incorporated nisin (an antibacterial peptide) in dental adhesives. Nisin has been approved as generally recognized as safe by the US Food and Drug
Enamel etching and dental adhesives 245 Administration since 1988 because of its long-term use in food preservation (E234) without causing health problems. It has effective antibiotic activity against cariogenic microorganisms. The results of this study showed that the nisin-incorporated adhesive significantly suppressed the growth of S. mutans. However, with the increase in concentration of these peptides, the bond strength was decreased (Su et al., 2018). Zinc oxide (ZnO) has been widely used in health sciences (Sarwar et al., 2017). In dental materials it shows antibacterial properties, inhibition of metalloproteinases, and biocompatibility (Wu et al., 2018; Henn et al., 2011; Toledano et al., 2012). It is reported that the antibacterial activity of ZnO is higher at the nanoscale (Mirzaei and Darroudi, 2017; Raghupathi et al., 2011); however, the mechanism behind these results is not well established. ZnO has been synthesized at the nanometric dimension in quantum dots (QDs) and these presented enhanced properties compared to material in the “bulk” form. Recently, ZnO QDs (ZnOQDs) were incorporated in experimental dental adhesive and investigated for antibacterial properties. It was found that the addition of ZnOQDs reduced biofilm formation in experimental adhesive resin surface by about 50%. It is anticipated that ZnO nanoparticles bound on the microorganism’s surface and disorganized the bacterium membrane by modifying the molecular structure of phospholipids, entering into the cell matrix and leading to modification of the cell shape and to leakage of intracellular constituents (Garcia et al., 2018). In another study, nanocapsules containing indomethacin and triclosan were added to commercial and experimental dental adhesives and this study confirmed the favorability of adding nanocapsules for achieving a reliable long-term bond performance. Addition of nanocapsules in dental adhesives can prevent the inflammatory process of pulp tissue in deep cavities and protect against caries through controlled drug release via these nanocapsules. Furthermore, these experimental adhesives had no influence on the in situ degree of conversion or the immediate bond strength. However, the bond strength decreased after 1 year of aging (Genari et al., 2018). Geraldeli et al. introduced various concentrations of arginine in experimental dental adhesives. Arginine is an amino acid found in a variety of foods and is also naturally produced by the human body and secreted in saliva in free form or as salivary peptides. It was found that the adhesive system containing 7% arginine was able to retain appropriate physical and mechanical properties. Moreover, it showed controlled release and moderate recharge of arginine over a prolonged period of time. Arginine was released from the adhesive system at a rate and concentration to exhibit antibacterial effect (Geraldeli et al., 2017). 11.7.2 Bioactive properties Since their development, dental adhesives have come a long way regarding their bonding ability and ease of application. However, the tooth restoration interface remains the weakest link in the restorations. Microgap formation at the interface leads to microleakage, which results in secondary caries formation. To increase the
246 Advanced Dental Biomaterials longevity of the restorations by further improving the clinical performance of adhesive, recent developments include imparting bioactive properties in addition to antibacterial properties to the bonding resins (Zhang et al., 2015a; Taubock et al., 2014; Tay and Pashley, 2009; Chen et al., 2014). Synthetic calcium phosphate materials, that is, amorphous calcium phosphate, HA, monocalcium phosphate monohydrate, and β-tricalcium phosphate, have been incorporated in dental adhesives (Eliaz and Metoki, 2017; Xie et al., 2017; Pei et al., 2017; Sfalcin et al., 2017; Yokota et al., 2016). Bioactive glass (BAG) is another type of bioceramic, which has gained popularity since its invention in 1969 by Prof. Hench. It consists mainly of silicon, calcium, sodium, phosphorous, and oxygen (Yoshihara et al., 2017) and has also been used as dental adhesive to remineralize the tooth surface (De Morais et al., 2018). These bioactive fillers can bond with the living tissue chemically by forming a calcium phosphate layer at the toothmaterial interface that renders the restoration durable and prevents it from bacterial ingression (Khan et al., 2008). Dental adhesives with remineralizing capabilities have been developed. An adhesive with the ability to release Ca and P ions may enhance the HA formation which is the main mineral component of the tooth structure. Ca and P ions released from the adhesive can facilitate in the remineralization of the dentin HL and at the margins of the restoration and the tooth surface. The remineralization of the HL can make it more resistant to oral environment degradation and capable of neutralizing biofilm acid. In addition, Ca and P released from an adhesive may contribute to the remineralization of infected and affected cavity lesions (Imazato et al., 2014; Zhang et al., 2015a). A 30-day in vitro study employed demineralization/remineralization cyclic regimen to assess the remineralization potential of NACP nanocomposites. The results showed that the remineralization effects of NACP nanocomposite on human enamel were fourfold that of a commercial fluoride-releasing composite (Weir et al., 2012). In another study, NACP was incorporated into commercially available bonding agent (Scotchbond Multi-Purpose) and its remineralization potential was investigated. The results showed that incorporation of NACP at 40% filler content resulted in increased release of Ca and P ions which have remineralization potential (Chen et al., 2014). BAGs have been widely used for bone defects repair and as an implant material due to their bone regeneration ability. BAG has the ability to release calcium and phosphate ions, which have remineralization effects on the demineralized tooth structure. In addition, dissolution of BAG in an aqueous environment releases alkali ions, resulting in an increase in pH which gives BAG antibacterial and acid neutralizing properties (Yang et al., 2013, 2016; Xie et al., 2008; Zehnder et al., 2004). Remineralization of mineral-depleted areas by BAG reduces the micropermeability along the dentin-bonded interface (Sauro et al., 2012). Formation of incipient caries, known as white spot lesions (WSL) around orthodontic brackets, is a common side effect of fixed orthodontic appliances. These lesions manifest demineralization of enamel due to bacterial action. Cariogenic bacteria which surround orthodontic appliances use fermentable carbohydrate to produce acids which lead to the reduction of the local plaque pH, resulting in the depletion of calcium and phosphate
Enamel etching and dental adhesives 247 ions from the enamel structure and resulting in WSL (Mitchell, 1992). Various studies have reported on the incorporation of BAG in the adhesive systems for orthodontic brackets. BAG addition in the adhesives has been shown to prevent demineralization of enamel due to the Ca and P ion releasing ability, thereby preventing WSL (Yang et al., 2016; Brown et al., 2011; Manfred et al., 2013). Sauro et al. (2012) investigated the beneficial effects of Bioglass 45S5 and zincpolycarboxylated BAG-based dental adhesives on the bonded dentin interface. The concentration of these particles was 33 wt.%. The study showed an increase in nanomechanical properties, while simultaneously reducing the micropermeability along the dentin-bonded interface. This occurred by means of the therapeutic remineralization of imperfect mineral-depleted areas. Furthermore, the surface of the specimen exhibited several “funneled” dentinal tubules with no exposed collagen fibrils and it was mainly protected by residual resin and mineral crystals which were embedded within a resin/collagen network. Both BAG and fluoride-based BAG (F-BAG) resins induced the precipitation of mineral contents inside the dentinal tubules and onto the surface. Also, the F-BAG resins showed a higher modulus of elasticity (13%) and an increase in stiffness of demineralized dentin when compared to BAG-based samples and control samples. This might be due to the release of fluoride and phosphate ions, which can accelerate the remineralization process. Recently, F-BAG and BAG have been added in dimethacrylate-based resins and it has been found that F-BAG-based composite shows more bioactivity than conventional BAG, as well as accelerated dentin remineralization. F-BAG furthermore showed a higher ability to reduce the solubilization of C-terminal cross-linked telopeptides and C-terminal telopeptides (Tezvergil-Mutluay et al., 2017). Khvostenko et al. synthesized BAG (15 wt.%) and incorporated it into resin matrices along with strontium glass, and their in vitro study showed a significant antimicrobial effect by reducing the extent of bacterial biofilm penetration into preexisting marginal gaps. Recently, niobium-based BAG has been incorporated in dental adhesives and results showed that the commercial adhesive at 30% niobium-based BAG concentrations did not compromise the bond strength of the dental adhesive, but did show high radiopacity and presented mineral precipitation, which characterized the experimental adhesives as bioactive materials (Carneiro et al., 2018). Another study showed that the incorporation of boron nitride nanotubes up to 0.15 wt.% improved the chemical and mechanical properties of dental adhesives and promoted mineral deposition (Degrazia et al., 2017). The dental adhesives based on these bioceramics have shown their potential to be used in clinical dentistry. 11.8 Conclusion It is well established that over time dental adhesives have shown quite acceptable results. However, concern has been raised related to their mechanical and physical properties. Therefore in the future, more research is required to improve their physical and mechanical properties without decreasing the biological
248 Advanced Dental Biomaterials potential. Furthermore, the improvement of their composition to enhance their antibacterial properties can help in minimizing the risks of secondary caries formation. The use of bioceramics can improve biological and mechanical properties; an ionocovalent structural model of the calcium phosphate family seems promising as reinforcing fillers. Surface modification during in situ synthesis and doping with amine, bromine, and with metal ions can also lead to improvements in the properties of fillers of dental adhesives. In conclusion we can say that so far the advancements in dental adhesives are tremendous, but further research is required to optimize the performance in clinical application fully. References Amin, S., Shetty, H.K., Varma, R.K., Amin, V., Nair, P.M., 2014. Comparative evaluation of antibacterial activity of total-etch and self-etch adhesive systems: an ex vitro study. J. Conserv. Dent. 17, 266 270. Andre, C.B., Gomes, B.P., Duque, T.M., Stipp, R.N., Chan, D.C., Ambrosano, G.M., et al., 2015. Dentine bond strength and antimicrobial activity evaluation of adhesive systems. J. Dent. 43, 466 475. Asmussen, E., Uno, S., 1992. Adhesion of restorative resins to dentin: chemical and physicochemical aspects. Oper. Dent. (Suppl. 5), 68 74. Baier, R.E., 1992. Principles of adhesion. Oper. Dent. (Suppl. 5), 1 9. Barbour, M., Rees, J., 2004. The laboratory assessment of enamel erosion: a review. J. Dent. 32, 591 602. Barkmeier, W.W., Cooley, R.L., 1992. Laboratory evaluation of adhesive systems. Oper. Dent. (Suppl. 5), 50 61. Bowen, R., 1980. Adhesive bonding of various materials to hard tooth tissues. XXII. The effects of a cleanser, mordant, and polySAC on adhesion between a composite resin and dentin. J. Dent. Res. 59, 809 814. Bowen, R.L., Marjenhoff, W.A., 1992. Development of an adhesive bonding system. Oper. Dent. (Suppl. 5), 75 80. Bozalis, W., Marshall, J.G., Cooley, R., 1979. Mechanical pretreatments and etching of primary-tooth enamel. ASDC J. Dent. Child. 46, 43 49. Brenna, F., 2012. Restorative Dentistry: Treatment Procedures and Future Prospects. Elsevier/Mosby, St. Louis, MO. Breschi, L., Gobbi, P., Mazzotti, G., Falconi, M., Ellis, T.H., Stangel, I., 2002. High resolution SEM evaluation of dentin etched with maleic and citric acid. Dent. Mater. 18, 26 35. Brown, M.L., Davis, H.B., Tufekci, E., Crowe, J.J., Covell, D.A., Mitchell, J.C., 2011. Ion release from a novel orthodontic resin bonding agent for the reduction and/or prevention of white spot lesions. An in vitro study. Angle Orthod. 81, 1014 1020. Buonocore, M., Matsui, A., Gwinnett, A., 1968. Penetration of resin dental materials into enamel surfaces with reference to bonding. Arch. Oral Biol. 13, 61 70. Cal-Neto, J.P., Miguel, J.A.M., 2006. Scanning electron microscopy evaluation of the bonding mechanism of a self-etching primer on enamel. Angle Orthod. 76, 132 136.
Enamel etching and dental adhesives 249 Carneiro, K.K., Araujo, T.P., Carvalho, E.M., Meier, M.M., Tanaka, A., Carvalho, C.N., et al., 2018. Bioactivity and properties of an adhesive system functionalized with an experimental niobium-based glass. J. Mech. Behav. Biomed. Mater. 78, 188 195. Chen, C., Weir, M.D., Cheng, L., Lin, N.J., Lin-Gibson, S., Chow, L.C., et al., 2014. Antibacterial activity and ion release of bonding agent containing amorphous calcium phosphate nanoparticles. Dent. Mater. 30, 891 901. Cheng, L., Zhang, K., Weir, M.D., Melo, M.A., Zhou, X., Xu, H.H., 2015. Nanotechnology strategies for antibacterial and remineralizing composites and adhesives to tackle dental caries. Nanomedicine 10, 627 641. Cocco, A.R., Rosa, W.L., Silva, A.F., Lund, R.G., Piva, E., 2015. A systematic review about antibacterial monomers used in dental adhesive systems: current status and further prospects. Dent. Mater. 31, 1345 1362. De Morais, R.C., Silveira, R.E., Chinelatti, M., Geraldeli, S., De Carvalho Panzeri Pires-DeSouza, F., 2018. Bond strength of adhesive systems to sound and demineralized dentin treated with bioactive glass ceramic suspension. Clin. Oral Investig. 22, 1923 1931. Degrazia, F.W., Leitune, V.C.B., Samuel, S.M.W., Collares, F.M., 2017. Boron nitride nanotubes as novel fillers for improving the properties of dental adhesives. J. Dent. 62, 85 90. Devarasa, G., Subba Reddy, V., Chaitra, N., Swarna, Y., 2012. Self-etching adhesive on intact enamel, with and without pre-etching. Microsc. Res. Tech. 75, 650 654. Di Lenarda, R., Cadenaro, M., Sbaizero, O., 2000. Effectiveness of 1 mol L21 citric acid and 15% EDTA irrigation on smear layer removal. Int. Endod. J. 33, 46 52. Eliaz, N., Metoki, N., 2017. Calcium phosphate bioceramics: a review of their history, structure, properties, coating technologies and biomedical applications. Materials 10. Erickson, R.L., 1992. Surface interactions of dentin adhesive materials. Oper. Dent. (Suppl. 5), 81 94. Fatemeh, K., Mohammad Javad, M., Samaneh, K., 2017. The effect of silver nanoparticles on composite shear bond strength to dentin with different adhesion protocols. J. Appl. Oral. Sci. 25, 367 373. Garcia, I.M., Leitune, V.C.B., Visioli, F., Samuel, S.M.W., Collares, F.M., 2018. Influence of zinc oxide quantum dots in the antibacterial activity and cytotoxicity of an experimental adhesive resin. J. Dent. 73, 57 60. Garcia-Godoy, F., Gwinnett, A., 1991. Effect of etching times and mechanical pretreatment on the enamel of primary teeth: an SEM study. Am. J. Dent. 4, 115 118. Genari, B., Leitune, V.C.B., Jornada, D.S., Aldrigui, B.R., Pohlmann, A.R., Guterres, S.S., et al., 2018. Effect on adhesion of a nanocapsules-loaded adhesive system. Braz. Oral Res. 32, e008. Geraldeli, S., Soares, E.F., Alvarez, A.J., Farivar, T., Shields, R.C., Sinhoreti, M.A.C., et al., 2017. A new arginine-based dental adhesive system: formulation, mechanical and anticaries properties. J. Dent. 63, 72 80. Ghaffari, H., Mirhashemi, A., Baherimoghadam, T., Azmi, A., Rasooli, R., 2017. Effect of surface treatment on enamel cracks after orthodontic bracket debonding: Er,Cr:YSGG laser-etching versus acid-etching. J. Dent. 14, 259. Gwinnett, A.J., 1994. Chemically conditioned dentin: a comparison of conventional and environmental scanning electron microscopy findings. Dent. Mater. 10, 149 155. Han, Q., Li, B., Zhou, X., Ge, Y., Wang, S., Li, M., et al., 2017. Anti-caries effects of dental adhesives containing quaternary ammonium methacrylates with different chain lengths. Materials 10, 643.
250 Advanced Dental Biomaterials Haznedaroğlu, F., 2003. Efficacy of various concentrations of citric acid at different pH values for smear layer removal. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 96, 340 344. Henn, S., Nedel, F., De carvalho, R.V., Lund, R.G., Cenci, M.S., Pereira-Cenci, T., et al., 2011. Characterization of an antimicrobial dental resin adhesive containing zinc methacrylate. J. Mater. Sci. Mater. Med. 22, 1797 1802. Hennequin, M., Douillard, Y., 1995. Effects of citric acid treatment on the Ca, P and Mg contents of human dental roots. J. Clin. Periodontol. 22, 550 557. Heymann, H., Swift, E.J., Ritter, A.V., Sturdevant, C.M., 2013. Sturdevant’s Art and Science of Operative Dentistry. Elsevier/Mosby, St. Louis, MO. Imazato, S., Ma, S., Chen, J.H., Xu, H.H., 2014. Therapeutic polymers for dental adhesives: loading resins with bio-active components. Dent. Mater. 30, 97 104. Jedrychowski, J.R., Caputo, A.A., Kerper, S., 1983. Antibacterial and mechanical properties of restorative materials combined with chlorhexidines. J. Oral Rehabil. 10, 373 381. Kanca 3rd, J., 1992. Resin bonding to wet substrate. 1. Bonding to dentin. Quintessence Int. 23, 39 41. Kasraei, S., Atai, M., Khamverdi, Z., Nejad, S.K., 2009. The effect of nanofiller addition to an experimental dentin adhesive on microtensile bond strength to human dentin. J. Dent. Tehran Univ. Med. Sci. 6, 36 41. Khan, A.S., Ahmed, Z., Edirisinghe, M.J., Wong, F.S., Rehman, I.U., 2008. Preparation and characterization of a novel bioactive restorative composite based on covalently coupled polyurethane-nanohydroxyapatite fibres. Acta Biomater. 4, 1275 1287. Kim, S.R., Shin, D.H., 2014. Antibacterial effect of self-etching adhesive systems on Streptococcus mutans. Restor. Dent. Endod. 39, 32 38. Kugel, G., Ferrari, M., 2000. The science of bonding: from first to sixth generation. J. Am. Dent. Assoc. 131, 20S 25S. Lee, B.-S., Hsieh, T.-T., Lee, Y.-L., Lan, W.-H., Hsu, Y.-J., Wen, P.-H., et al., 2003. Bond strengths of orthodontic bracket after acid-etched, Er: YAG laser-irradiated and combined treatment on enamel surface. Angle Orthod. 73, 565 570. Li, F., Chen, J., Chai, Z., Zhang, L., Xiao, Y., Fang, M., et al., 2009. Effects of a dental adhesive incorporating antibacterial monomer on the growth, adherence and membrane integrity of Streptococcus mutans. J. Dent. 37, 289 296. Lopes, G.C., Vieira, L.C., Araujo, E., Bruggmann, T., Zucco, J., Oliveira, G., 2011. Effect of dentin age and acid etching time on dentin bonding. J. Adhes. Dent. 13, 139 145. Lutz, F., Krejci, I., Oddera, M., 1996. Advanced adhesive restorations: the post-amalgam age. Pract. Periodontics Aesthet. Dent. 8, 385 394. quiz 398. Manfred, L., Covell, D.A., Crowe, J.J., Tufekci, E., Mitchell, J.C., 2013. A novel biomimetic orthodontic bonding agent helps prevent white spot lesions adjacent to brackets. Angle Orthod. 83, 97 103. Mccomb, D., Smith, D.C., 1975. A preliminary scanning electron microscopic study of root canals after endodontic procedures. J. Endod. 1, 238 242. Melo, M.A., Cheng, L., Weir, M.D., Hsia, R.C., Rodrigues, L.K., Xu, H.H., 2013. Novel dental adhesive containing antibacterial agents and calcium phosphate nanoparticles. J. Biomed. Mater. Res. B Appl. Biomater. 101, 620 629. Mirzaei, H., Darroudi, M., 2017. Zinc oxide nanoparticles: biological synthesis and biomedical applications. Ceram. Int. 43, 907 914. Mitchell, L., 1992. Decalcification during orthodontic treatment with fixed appliances—an overview. Br. J. Orthod. 19, 199 205.
Enamel etching and dental adhesives 251 Mizunuma, T., 1986. Relationship between bond strength of resin to dentin and structural change of dentin collagen during etching. Influence of ferric chloride on structure of the collagen. J. Jpn. Dent. Mater. 5, 54 64. Mortazavi, V., Fathi, M., Ataei, E., Khodaeian, N., Askari, N., 2012. Shear bond strengths and morphological evaluation of filled and unfilled adhesive interfaces to enamel and dentine. Int. J. Dent. 2012, 858459. Moszner, N., Salz, U., Zimmermann, J., 2005. Chemical aspects of self-etching enamel dentin adhesives: a systematic review. Dent. Mater. 21, 895 910. Nakabayashi, N., Kojima, K., Masuhara, E., 1982. The promotion of adhesion by the infiltration of monomers into tooth substrates. J. Biomed. Mater. Res. Part A 16, 265 273. Ozer, F., Sengun, A., Ozturk, B., Say, E.C., Tagami, J., 2005. Effect of tooth age on microtensile bond strength of two fluoride-releasing bonding agents. J. Adhes. Dent. 7, 289 295. Pashley, D.H., Carvalho, R.M., 1997. Dentine permeability and dentine adhesion. J. Dent. 25, 355 372. Pashley, D.H., Tay, F.R., Breschi, L., Tjäderhane, L., Carvalho, R.M., Carrilho, M., et al., 2011. State of the art etch-and-rinse adhesives. Dent. Mater. 27, 1 16. Passariello, C., Sannino, G., Petti, S., Gigola, P., 2014. Intensity and duration of in-vitro antibacterial activity of different adhesives used in orthodontics. Eur. J. Oral Sci. 122, 154 160. Pei, D.D., Liu, S.Y., Yang, H.Y., Gan, J., Huang, C., 2017. Effect of a nano hydroxyapatite desensitizing paste application on dentin bond strength of three self-etch adhesive systems. Zhonghua Kou Qiang Yi Xue Za Zhi 52, 278 282. Perdigao, J., 2002. Dentin bonding as a function of dentin structure. Dent. Clin. North Am. 46, 277 301. vi. Perdigão, J., 2010. Dentin bonding—variables related to the clinical situation and the substrate treatment. Dent. Mater. 26, e24 e37. Perdigão, J., Sezinando, A., 2013. Enamel and dentin bonding for adhesive restorations. NonMetallic Biomaterials for Tooth Repair and Replacement. Elsevier. Peumans, M., Kanumilli, P., De munck, J., Van landuyt, K., Lambrechts, P., Van meerbeek, B., 2005. Clinical effectiveness of contemporary adhesives: a systematic review of current clinical trials. Dent. Mater. 21, 864 881. Prado, M., Gusman, H., Gomes, B.P., Simão, R.A., 2011. Scanning electron microscopic investigation of the effectiveness of phosphoric acid in smear layer removal when compared with EDTA and citric acid. J. Endod. 37, 255 258. Raghupathi, K.R., Koodali, R.T., Manna, A.C., 2011. Size-dependent bacterial growth inhibition and mechanism of antibacterial activity of zinc oxide nanoparticles. Langmuir 27, 4020 4028. Saeki, K., Marshall, S., Gansky, S., Marshall, G., 2001. Etching characteristics of dentin: effect of ferric chloride in citric acid. J. Oral Rehabil. 28, 301 308. Sarwar, S., Ali, A., Pal, M., Chakrabarti, P., 2017. Zinc oxide nanoparticles provide anticholera activity by disrupting the interaction of cholera toxin with the human GM1 receptor. J. Biol. Chem. 292, 18303 18311. Sauro, S., Osorio, R., Watson, T.F., Toledano, M., 2012. Therapeutic effects of novel resin bonding systems containing bioactive glasses on mineral-depleted areas within the bonded-dentine interface. J. Mater. Sci. Mater. Med. 23, 1521 1532. Say, E.C., Nakajima, M., Senawongse, P., Soyman, M., Ozer, F., Tagami, J., 2005. Bonding to sound vs caries-affected dentin using photo- and dual-cure adhesives. Oper. Dent. 30, 90 98.
252 Advanced Dental Biomaterials Sfalcin, R.A., Correr, A.B., Morbidelli, L.R., Araujo, T.G.F., Feitosa, V.P., Correr-Sobrinho, L., et al., 2017. Influence of bioactive particles on the chemical-mechanical properties of experimental enamel resin infiltrants. Clin. Oral Investig. 21, 2143 2151. Silverstone, L.M., 1974. Fissure sealants. Caries Res. 8, 2 26. Smutka, S., Jedrychowski, J., Caputo, A., 1978. An evaluation of primary enamel pretreatments and their effects on resin retention. J. Dent. Res. 57, 796 799. Sofan, E., Sofan, A., Palaia, G., Tenore, G., Romeo, U., Migliau, G., 2017a. Classification review of dental adhesive systems: from the IV generation to the universal type. Ann. Stomatol. (Roma) 8, 1. Sofan, E., Sofan, A., Palaia, G., Tenore, G., Romeo, U., Migliau, G., 2017b. Classification review of dental adhesive systems: from the IV generation to the universal type. Ann. Stomatol. (Roma) 8, 1 17. Su, M., Yao, S., Gu, L., Huang, Z., Mai, S., 2018. Antibacterial effect and bond strength of a modified dental adhesive containing the peptide nisin. Peptides 99, 189 194. Swift, E.J., Perdigao, J., Heymann, H.O., 1995. Bonding to enamel and dentin: a brief history and state of the art, 1995. Quintessence Int. 26, 95 110. Tao, L., Pashley, D.H., Boyd, L., 1988. Effect of different types of smear layers on dentin and enamel shear bond strengths. Dent. Mater. 4, 208 216. Taubock, T.T., Zehnder, M., Schweizer, T., Stark, W.J., Attin, T., Mohn, D., 2014. Functionalizing a dentin bonding resin to become bioactive. Dent. Mater. 30, 868 875. Tay, F.R., Pashley, D.H., 2009. Biomimetic remineralization of resin-bonded acid-etched dentin. J. Dent. Res. 88, 719 724. Tezvergil-Mutluay, A., Agee, K.A., Uchiyama, T., Imazato, S., Mutluay, M.M., Cadenaro, M., et al., 2011. The inhibitory effects of quaternary ammonium methacrylates on soluble and matrix-bound MMPs. J. Dent. Res. 90, 535 540. Tezvergil-Mutluay, A., Agee, K.A., Mazzoni, A., Carvalho, R.M., Carrilho, M., Tersariol, I. L., et al., 2015. Can quaternary ammonium methacrylates inhibit matrix MMPs and cathepsins? Dent. Mater. 31, e25 e32. Tezvergil-Mutluay, A., Seseogullari-Dirihan, R., Feitosa, V.P., Cama, G., Brauer, D.S., Sauro, S., 2017. Effects of composites containing bioactive glasses on demineralized dentin. J. Dent. Res. 96, 999 1005. Thanatvarakorn, O., Nakajima, M., Prasansuttiporn, T., Ichinose, S., Foxton, R.M., Tagami, J., 2014. Effect of smear layer deproteinizing on resin-dentine interface with self-etch adhesive. J. Dent. 42, 298 304. Toledano, M., Yamauti, M., Ruiz-Requena, M.E., Osorio, R., 2012. A ZnO-doped adhesive reduced collagen degradation favouring dentine remineralization. J. Dent. 40, 756 765. Tseng, W.Y., Chen, M.H., Lu, H.H., Lin, C.W., Hsieh, T.T., Chen, C.H., et al., 2007. Tensile bond strength of Er, Cr:YSGG laser-irradiated human dentin to composite inlays with two resin cements. Dent. Mater. J. 26, 746 755. Uysal, T., Amasyali, M., Koyuturk, A.E., Ozcan, S., Sagdic, D., 2010. Amorphous calcium phosphate-containing orthodontic composites. Do they prevent demineralisation around orthodontic brackets?. Aust. Orthod. J. 26, 10 15. Van Meerbeek, B., Lambrechts, P., Inokoshi, S., Braem, M., Vanherle, G., 1992. Factors affecting adhesion to mineralized tissues. Oper. Dent. (Suppl. 5), 111 124. Van Meerbeek, B., De munck, J., Yoshida, Y., Inoue, S., Vargas, M., Vijay, P., et al., 2003a. Adhesion to enamel and dentin: current status and future challenges. Oper. Dent. 28, 215 235. Van Meerbeek, B., De munck, J., Yoshida, Y., Inoue, S., Vargas, M., Vijay, P., et al., 2003b. Buonocore memorial lecture. Adhesion to enamel and dentin: current status and future challenges. Oper. Dent. 28, 215 235.
Enamel etching and dental adhesives 253 Vinagre, A., Ramos, J., 2016. Adhesion in restorative dentistry. Adhesives-Applications and Properties. InTech. Wang, L., Xie, X., Weir, M.D., Fouad, A.F., Zhao, L., Xu, H.H., 2016. Effect of bioactive dental adhesive on periodontal and endodontic pathogens. J. Mater. Sci. Mater. Med. 27, 168. Wei, S., Sadr, A., Shimada, Y., Tagami, J., 2008. Effect of caries-affected dentin hardness on the shear bond strength of current adhesives. J. Adhes. Dent. 10, 431 440. Weir, M.D., Chow, L.C., Xu, H.H.K., 2012. Remineralization of demineralized enamel via calcium phosphate nanocomposite. J. Dent. Res. 91, 979 984. Wiegand, A., Buchalla, W., Attin, T., 2007. Review on fluoride-releasing restorative materials—fluoride release and uptake characteristics, antibacterial activity and influence on caries formation. Dent. Mater. 23, 343 362. Wu, Y.Z., Sun, J., Yang, H., Zhao, X., He, D., Pu, M., et al., 2018. Biosynthetic mechanism of luminescent ZnO nanocrystals in the mammalian blood circulation and their functionalization for tumor therapy. ACS Appl. Mater. Interfaces 10, 105 113. Xie, D., Zhao, J., Weng, Y., Park, J.G., Jiang, H., Platt, J.A., 2008. Bioactive glass-ionomer cement with potential therapeutic function to dentin capping mineralization. Eur. J. Oral Sci. 116, 479 487. Xie, X., Wang, L., Xing, D., Zhang, K., Weir, M.D., Liu, H., et al., 2017. Novel dental adhesive with triple benefits of calcium phosphate recharge, protein-repellent and antibacterial functions. Dent. Mater. 33, 553 563. Yamaguchi, M., Yoshida, K., Suzuki, R., Nakamura, H., 1996. Root canal irrigation with citric acid solution. J. Endod. 22, 27 29. Yang, S.Y., Piao, Y.Z., Kim, S.M., Lee, Y.K., Kim, K.N., Kim, K.M., 2013. Acid neutralizing, mechanical and physical properties of pit and fissure sealants containing meltderived 45S5 bioactive glass. Dent. Mater. 29, 1228 1235. Yang, S.-Y., Kim, S.-H., Choi, S.-Y., Kim, K.-M., 2016. Acid neutralizing ability and shear bond strength using orthodontic adhesives containing three different types of bioactive glass. Materials 9, 125. Yokota, Y., Fujita, K.N., Uchida, R., Aida, E., Aoki, N.T., Aida, M., et al., 2016. Quantitative evaluation of MDP-Ca salt and DCPD after application of an MDP-based one-step self-etching adhesive on enamel and dentin. J. Adhes. Dent. 18, 205 213. Yoshihara, K., Nagaoka, N., Maruo, Y., Sano, H., Yoshida, Y., Van meerbeek, B., 2017. Bacterial adhesion not inhibited by ion-releasing bioactive glass filler. Dent. Mater. 33, 723 734. Zehnder, M., Soderling, E., Salonen, J., Waltimo, T., 2004. Preliminary evaluation of bioactive glass S53P4 as an endodontic medication in vitro. J. Endod. 30, 220 224. Zhang, K., Li, F., Imazato, S., Cheng, L., Liu, H., Arola, D.D., et al., 2013. Dual antibacterial agents of nano-silver and 12-methacryloyloxydodecylpyridinium bromide in dental adhesive to inhibit caries. J. Biomed. Mater. Res. B Appl. Biomater. 101, 929 938. Zhang, L., Weir, M.D., Hack, G., Fouad, A.F., Xu, H.H.K., 2015a. Rechargeable dental adhesive with calcium phosphate nanoparticles for long-term ion release. J. Dent. 43, 1587 1595. Zhang, N., Melo, M.A., Chen, C., Liu, J., Weir, M.D., Bai, Y., et al., 2015b. Development of a multifunctional adhesive system for prevention of root caries and secondary caries. Dent. Mater. 31, 1119 1131.
Endodontic materials: from old materials to recent advances 12 Mai Saleh Ali1,2 and Bassel Kano3 1 Faculty of Dentistry, University of Toronto, Toronto, ON, Canada, 2Private Practice, Amman, Jordan, 3Division of Endodontics, Faculty of Dentistry, McGill University, Montreal, QC, Canada Chapter Outline 12.1 Introduction 256 12.2 Materials used in vital pulp therapy 12.2.1 12.2.2 12.2.3 12.2.4 12.2.5 12.3 Materials used as root canal irrigants 12.3.1 12.3.2 12.3.3 12.3.4 12.3.5 12.3.6 12.3.7 12.3.8 12.3.9 12.3.10 Sodium hypochlorite 263 Ethylenediamine tetra-acetic acid 265 Chlorhexidine 266 Citric acid 267 MTAD 267 Tetraclean 269 Hydrogen peroxide 269 Iodine potassium iodide 269 1-Hydroxyethylidene-1,1-bisphosphonate QMiX 270 12.4 Intracanal medicaments 12.4.1 12.4.2 12.4.3 12.4.4 12.4.5 256 Mineral trioxide aggregates 257 Biodentine (Septodont, Saint-Maur-des-Fosses, France) 262 Bioaggregate (Innovative Bioceramix, Vancouver, BC, Canada) 262 Mineral Trioxide Aggregate Angelus (Londrina, PR, Brazil) 262 Endosequence (Brasseler USA, Savanah, Georgia, United States) 262 263 269 271 Calcium hydroxide 271 Chlorhexidine 272 Ledermix 273 Triple antibiotics pastes 273 Bioactive glass 274 12.5 Root canal obturation materials 274 12.5.1 Core obturation materials 274 12.5.2 Root canal sealers (cementing medium) 277 12.6 Root-end filling materials 12.6.1 12.6.2 12.6.3 12.6.4 12.6.5 290 Amalgam 290 Zinc oxide eugenol cements 290 Composite resins (Retroplast) 290 Glass ionomer cements 292 Diaket (3M/ESPE, Seefeld, Germany) 292 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00012-8 Copyright © 2019 Elsevier Ltd. All rights reserved.
256 Advanced Dental Biomaterials 12.6.6 Resin ionomer suspension and compomer 292 12.6.7 Other types of cement 292 12.7 Perforation repair materials 12.8 Summary 293 References 293 Further reading 299 12.1 292 Introduction The aim of endodontics is the diagnosis, prevention, and treatment of diseases and injuries of the dental pulp along with the associated periradicular structures. Vital pulp therapy, nonsurgical treatment of the root canal system, periradicular surgery for persistent infections, and repair of iatrogenic perforations are a few of the endodontic procedures that require the use of materials to reach a favorable result. Fortunately, the field of endodontics has greatly benefited from the recent advances and technologies in material sciences and equipment, thus providing the clinician with options that can facilitate treatment with high success rates. This chapter will cover materials used in vital pulp therapy, irrigants, intracanal medicaments, obturation materials, root-end fillings, and repair materials. 12.2 Materials used in vital pulp therapy In the cases of teeth with reversible pulpitis or partially inflamed pulps caused by chemical, mechanical, or traumatic injury, where the healthy remaining pulpal tissue can be conserved, vital pulp therapy is recommended to generate a hard tissue barrier to preserve, seal, and protect the pulp from microbial insult (Fig. 12.1). This can prolong tooth survival and delay or avoid root canal treatment. Indirect pulp capping, direct pulp capping, and pulpotomy are all different vital pulp therapy procedures. Caries extension and pulp exposure are factors affecting the decision of which procedure is to be approached. Properties of an ideal material for vital pulp therapy are summarized in Box 12.1; nevertheless, none of those available satisfy them all (Qureshi et al., 2014). Table 12.1 summarizes some of the materials that have been used for vital pulp therapy through the years. However, recently bioceramic materials, such as calcium silicate materials, calcium aluminate cements, and calcium phosphate cements, are among the most widely studied materials due to their bioactivity (Jitaru et al., 2016), regenerative abilities, and their promising results not only as materials for vital pulp therapy, but also as root repair materials and as root-end fillings. Among the increasing number of bioceramic materials, calcium silicate cements are the most studied. Therefore the following section will focus on a few materials that belong to this family of cements.
Endodontic materials: from old materials to recent advances 257 Figure 12.1 Vital pulp therapy: (A) indirect pulp capping can be defined as “a procedure performed in a tooth with a deep carious lesion approximating the pulp but without signs or symptoms of pulp degeneration” and “indicated in a permanent tooth diagnosed with no signs or symptoms of pulpitis or with a diagnosis of reversible pulpitis”; (B) direct pulp capping is defined by the American Association of Endodontists (2012) as “placing a material directly on a mechanical or traumatic vital pulp exposure to induce reparative dentin formation”; and (C) pulpotomy, defined by the AAE as “the removal of the coronal portion of the vital pulp as a means of preserving the vitality of the remaining radicular portion.” AAE, American Association of Endodontists. Box 12.1 Properties of an ideal vital pulp therapy material G G G G G G G G G G G G G G Stimulates reparative dentin formation Maintains pulpal vitality Releases fluoride to prevent secondary caries Bactericidal or bacteriostatic Adheres to dentin Adheres to restorative material Resists forces during restoration placement and during the life of restoration Sterile Radiopaque Provides a seal against bacteria and the oral environment Insoluble in biological fluids Dimensionally stable Nonresorbable, nontoxic, noncarcinogenic, nongenotoxic Biocompatible and bioactive 12.2.1 Mineral trioxide aggregates Mineral trioxide aggregate (MTA) was introduced in the mid-1990s by Dr. Mahmoud Torabinejad and associates as a promising pulp capping material (Ford et al., 1996). ProRoot MTA Gray (Dentsply Tulsa Dental Specialties, Johnson City, Tennessee, United States) was the first material to be marketed and it is composed of 75% Type I Portland cement, 20% bismuth oxide, and 5% calcium sulfate
Table 12.1 Some of the materials used for vital pulp therapy. Material Description Polycarboxylate cement Corticosteroids and antibiotics It is not used anymore due to its lack of calcific bridge formation and antibacterial effect (McWalter et al., 1976) The rationale is that those materials will reduce pulpal inflammation. However, there is lack of research to their viability as a vital pulp therapy material CH: Dycal (Dentsply, United States), Life (Kerr, United States) It has been considered as the universal standard for vital pulp therapy materials; however, its drawbacks can no longer make it the preferred one RMGI G G Adhesive systems Clearfil SE Bond (Kuraray Medical, Japan) Optibond S (Kerr, United States) Prime & Bond 2.1 (Dentsply, United States) Composition G G G G Base paste: titanium dioxide, calcium tungstate 1,3-butylene glycol disalicylate Catalyst paste: CH, zinc oxide, zinc stearate, ethyl toluene sulfonamide 1. Reparative hard tissue bridge formation subjacent to CH has tunnel defects (Andelin et al., 2003) 2. It has been shown CH can induce pulp cell apoptosis 3. Dimensionally unstable with weak marginal adaptation to dentin 4. CH degrades and dissolutes with time allowing microleakage and further pulpal injury 5. CH can cause primary tooth resorption If used as a direct pulp capping material, there is lack of dentin bridge formation and the pulp exhibits chronic inflammation (Tarim et al., 1998) Primer: MDP, dimethacrylate monomer, HEMA, silica, N,Ndiethanol-p-toluidine, camphorquinone Bond: HEMA, dimethacrylate monomer, bis-GMA, N,N-diethanolp-toluidine silica, camphorquinone Cytotoxic to pulp cells (Silva et al., 2013) Best used as an indirect pulp capping material with no direct contact with the pulp Conventional glass ionomer is less toxic than the resin-modified formulations G G G Disadvantages (Continued)
Table 12.1 (Continued) Material Zinc oxide eugenol IRM (Dentsply, United States) Temp Bond (Kerr, United States) Relix Tempo (3M ESPE, United States) Formocresol Buckley’s Formocresol (Sultan Healthcare, United States) Description Composition G G G Powder: zinc oxide, PMMA pigment Liquid: eugenol, acetic acid G G G G G Ferric sulfate: ViscoStat (Ultradent, United States), Astringedent (Ultradent, United States) Collagen Hydroxyapatite G G G GF BMP which belongs to super family TGF-β Recombinant insulin like GF-I G G G It is mostly used in pulpotomies in primary teeth It is applied on the radicular stump for 1 5 min Formaldehyde, ortho-cresol, glycerin, ethyl alcohol It is mostly used in pulpotomies in primary teeth as an alternative to formocresol It is applied on the radicular stump for 10 15 s (Loh et al., 2004) Ferric sulfate, silica, aqueous vehicle There is lack of research regarding hydroxyapatite but a study showed. It can be used as scaffolding for the newly formed mineralized tissue (Hayashi et al., 1999) Disadvantages 1. Direct contact with pulp tissue causes chronic inflammation and ultimately necrosis 2. It lacks the formation of a calcific barrier (Glass and Zander, 1949) G G It is a caustic material and causes severe burns in contact The use of formaldehyde is controversial due to its toxicity and corrosiveness (Lewis and Chestner, 1981) It does not form a thick dentin bridge (Dick and Carmichael, 1980) More research is needed More research is needed (Lianjia et al., 1993; Lovschall et al., 2001) (Continued)
Table 12.1 (Continued) Material Description Propolis (Russian penicillin) EMD It is an enamel matrix derivative. It has an important role during periodontal tissue formation and stimulates PDL regeneration It comes as an injectable gel solution (Al-Hezaimi et al., 2011; Kaida et al., 2008) COB cement is a natural polyol containing three hydroxyl radicals Used for vital pulp therapy, root and perforation repair, and as a rootend filling. Introduced by Asgary et al. (2006) G G COB cement Calcium phosphate cement CEM cement (BioniqueDent, Iran) G Composition Disadvantages Contains flavonoids, phenolics, iron, zinc, and other various aromatic compounds (Parolia et al., 2010) It consists of enamel matrix proteins (amelogenin), water and a carrier (propylene glycol alginate). It also contains BMP-like molecules and BMP-expressing cells More research is needed 81% 96% Triglyceride of ricinoleic acid Powder: calcium compounds such as calcium oxide, calcium phosphate, calcium carbonate, calcium silicate, calcium sulfate, CH, calcium chloride Liquid: distilled water Powder: aluminum oxide, calcium oxide, silicon dioxide, magnesium oxide, iron oxide, bismuth oxide Liquid: distilled water More research is needed (Camargo et al., 2009) More research is needed (Ambard and Mueninghoff, 2006; Asgary et al., 2015) G Some clinical trials failed to prove that EMD improves tertiary dentin formation when used as direct pulp capping material in primary teeth and pulpotomy in permanent teeth (Kiatwateeratana et al., 2009) G Calcium aluminate cement EndoBinder (Binderware, Brazil) Capasio (Primus Consulting, Bradenton, FL) Quick-Set (Primus Consulting, Bradenton, FL) G Used for vital pulp therapy, root and perforation repair, and as a rootend filling G More research is needed G G Bis-GMA, Bisphenol A glycidyl methacrylate; BMP, bone morphogenetic protein; CEM, calcium enriched mixture; CH, calcium hydroxide; COB, castor oil bean; EMD, Emdogain; GF, growth factors; HEMA, hydroxyethyl methacrylate; IRM, intermediate restorative material; MDP, 10-methacryloyloxydecyl dihydrogen phosphate; PDL, periodontal ligament; PMMA, polymethyl methacrylate; RMGI, resin-modified glass ionomer cements; TGF-β, transforming growth factor beta.
Endodontic materials: from old materials to recent advances 261 dehydrate (gypsum). Portland cement is composed of tricalcium silicate, dicalcium silicate, tricalcium aluminate tetracalcium aluminoferrite, magnesium oxide, calcium sulfate, and calcium oxide. Bismuth oxide is added as a radiopacifier, while calcium sulfate acts as a setting modifier (Camilleri et al., 2005). In 2002 white MTA was introduced to prevent the tooth discoloration caused by gray MTA, by reducing the amount of iron and magnesium. Moreover, white MTA has smaller, finer, and more homogenous powder particles and thus allows easier handling properties than gray MTA (Parirokh and Torabinejad, 2010a). One of the positive characteristics of MTA is that it is a hydraulic material which means it sets in the presence of moisture (i.e., moisture affects its setting favorably). Nevertheless, excessive moisture may cause a detrimental effect (Fridland and Rosado, 2003). The reaction of MTA and water is considered as an exothermic hydration reaction. The main ingredients that are hydrated during this reaction are tricalcium silicate and dicalcium silicate which forms calcium silicate hydrates and calcium hydroxide. Both reaction products play a role in the bioactivity of MTA. The calcium silicate hydrates form a colloidal gel that solidifies into a hard structure with calcium hydroxide nucleated in an amorphous matrix (Parirokh and Torabinejad, 2010a). MTA has a working time of 5 minutes and sets over a period of 4 hours. When used as a capping material, it is recommended to place a rigid restorative material such as adhesive fillings over MTA in the subsequent clinical appointment after a few days. This restorative material will provide adequate strength to withstand occlusal forces and allow MTA to complete its hydration reaction and gain adequate strength. Table 12.2 lists the advantages and disadvantages of MTA. Table 12.2 Advantages and disadvantages of mineral trioxide aggregates (MTA). Advantages Disadvantages 1. Its ability to stimulate reparative dentinogenesis and hard tissue formation by inducing the release of dentin matrix components that are essential for hard tissue repair (Yasuda et al., 2008) 2. MTA has a superior marginal adaptation to dentin since it forms an interfacial layer at the surface of dentin that is similar to hydroxyapatite (Min et al., 2008) 3. Excellent biocompatibility (Yavari et al., 2009) 4. Excellent sealing ability (Estrela et al., 2011) 5. Antibacterial due to its high pH (Torabinejad et al., 1995) 6. Radiopaque (Borges et al., 2011) 1. Discoloration of tooth structure especially with gray MTA 2. Expensive 3. Difficult handling characteristics 4. Prolonged setting time (Parirokh and Torabinejad, 2010b; Kratchman, 2004; Camilleri et al., 2005) MTA, Mineral trioxide aggregate.
262 Advanced Dental Biomaterials 12.2.2 Biodentine (Septodont, Saint-Maur-des-Fosses, France) Biodentine was manufactured as a “dentin replacement” material. It is a powder/ liquid material mixed by trituration. It consists of tricalcium silicate, dicalcium silicate, calcium carbonate, iron oxide, and zirconium oxide (radiopacifier), whereas the liquid contains calcium chloride as an accelerator and a hydrosoluble polymer that serves as a water reducing agent. Like other calcium silicate materials, it has a hydration setting reaction with calcium hydroxide as one of the reaction products, it is bioactive with alkaline pH, biocompatible and has excellent sealability (Grech et al., 2013). Unlike MTA, biodentine is characterized by a short setting time (9 12 minutes) and color stability (Keskin et al., 2015). 12.2.3 Bioaggregate (Innovative Bioceramix, Vancouver, BC, Canada) Bioaggregate, also known as DiaRoot, consists of tricalcium silicate, dicalcium silicate, calcium phosphate monobasic, amorphous silicon dioxide, and tantalum pentoxide (as a radiopacifier) (Camilleri et al., 2015). Bioaggregate has a fine nanoparticle aluminum-free powder that is mixed with deionized water. It is bioinductive, highly biocompatible, and antibacterial. Compared to MTA, bioaggregate demonstrates higher fracture resistance and a greater resistance to dislodgement in an acidic environment (Hashem and Wanees Amin, 2012). 12.2.4 Mineral Trioxide Aggregate Angelus (Londrina, PR, Brazil) It is composed of 80% Portland cement and 20% bismuth oxide (Duarte et al., 2003). It has a lower radiopacity compared to gray MTA and its lack of calcium sulfate provides a short setting time (10 minutes). 12.2.5 Endosequence (Brasseler USA, Savanah, Georgia, United States) Endosequence comes in two forms: the first is endosequence root repair material (ERRM) in a preloaded paste consistency in a syringe, and the second is endosequence root repair putty (ERRP)—a premixed moldable putty (Lovato and Sedgley, 2011). Like other bioceramics and calcium silicates, it is bioactive, biocompatible with antibacterial activity, and it can be used as a pulp capping material in addition to root repair. Both ERRM and ERRP are bioceramic materials that come as bright white nanosphere particles that are composed of a combination of calcium silicate, calcium phosphate monobasic, zirconium oxide, tantalum oxide, filler, and thickening agents. They are also premixed and ready to use. There are several calcium silicate and bioceramic materials that are regularly being introduced to the market—to name a few, Ortho MTA (BioMTA, Seoul, Korea), MTA Bio (Angelus, Londrina, PR, Brazil), Fluoride-Doped MTA Cement, Cimento Endodontico Rapido, Endocem, Ceramicrete-D (Tulsa Dental
Endodontic materials: from old materials to recent advances 263 Specialties/Argonne National Laboratory, Argonne, Illinois, United States), Generex A (Dentsply Tulsa Dental Specialties, Tulsa, Oklahoma, United States), Tech Biosealer capping (Isasan Srl, Rovello Porro, Co, Italy), TheraCal (Bisco Inc, Schaumburg, Illinois, United States), and many others. All the abovementioned materials demonstrate promising bioactive potential and possess somewhat similar properties with various modifications to their composition in hopes of overcoming the shortcomings of MTA. 12.3 Materials used as root canal irrigants The next step after a straight-line access cavity is the removal of pulpal tissue, microbes, and dentine shavings from the canal. This is achieved via mechanical preparation and chemical irrigation. Root canal irrigation plays an important role in the disinfection of the root canal system and is critical for the treatment’s success. According to Zehnder (2006), an ideal irrigant should possess all the properties mentioned in Box 12.2, yet none of the available ones combine all those characteristics. Accordingly, clinicians often use the irrigant that has many of the most important properties or a combination of different irrigants that complement each other’s action. Table 12.3 summarizes the properties and effects of different irrigants used during root canal treatment. 12.3.1 Sodium hypochlorite Sodium hypochlorite (NaOCl) is the most commonly used irrigant worldwide since it satisfies most of the properties stated in Box 12.2 (Mohammadi, 2008). It is an alkaline, hypertonic solution with a pH . 11. In water, NaOCl ionizes to produce Box 12.2 Ideal properties of a root canal irrigant G G G G G G G G G G G G Broad antimicrobial spectrum High efficacy against anaerobic and facultative microorganisms organized in biofilms Able to dissolve necrotic pulp tissue remnants Able to inactivate endotoxins Able to prevent the formation of a smear layer during instrumentation or to dissolve it once formed Systemically, nontoxic, nonantigenic, and noncarcinogenic if it comes in contact with vital tissues Noncaustic to periodontal tissues Little potential to cause an anaphylactic reaction Has low surface tension Does not stain tooth structures Has no adverse effects on the sealing ability of filling material Relatively inexpensive
Table 12.3 Summary of the properties and effects of different irrigants. Irrigant Antibacterial effect Dissolution of tissues Smear layer removal Concentration (%) NaOCl Effective antibacterial agent by breaking down proteins into amino acids Dissolves necrotic tissue at low concentration At high concentration, it dissolves both vital and necrotic tissues No dissolution capacity Removes the organic part of smear layer (Goldman et al., 1976) 0.5 6.25 Removes the inorganic part of smear layer No smear layer removal 17 EDTA Limited antibacterial effect CHX Effective antibacterial agent against Gram negative, Gram, Gram positive, and yeasts Antibacterial effect due to doxycycline No antibacterial effect Antibacterial effect Antibacterial effect Very minimal antibacterial effect Antibacterial effect MTAD Citric acid H2O2 IKI HEBP QMiX G G No dissolution capacity No dissolution capacity No dissolution capacity No dissolution capacity No dissolution capacity No dissolution capacity No dissolution capacity Removes smear layer through doxycycline and citric acid Removes smear layer None None Weak decalcifying property Removes smear layer 2 10 50 3 5 2 5 CHX, Chlorhexidine; EDTA, ethylenediamine tetra-acetic acid; HEBP, 1-hydroxyethylidene-1,1-bisphosphonate; IKI, iodine potassium iodide; MTAD, a mixture of doxycycline, citric acid, and a detergent; NaOCl, sodium hypochlorite.
Endodontic materials: from old materials to recent advances 265 sodium ion (Na1) and hypochlorite ion (OCl2) which establishes equilibrium with hypochlorous acid HOCL according to the following equation: NaOCl1H2 O2NaOH 1 HOCl2Na1 1OH2 1H 1 1OCl2 Once NaOCl comes in contact with tissue protein, nitrogen, formaldehyde, and acetaldehyde are formed. This protein disintegration enables the replacement of hydrogen in the amino groups ( NH ) by chlorine, and thus forming chloramines ( NCl ). Chloramination contributes to the antimicrobial properties effect of NaOCl and makes this irrigant distinguished by its ability to dissolve both vital and necrotic pulp tissues. However, NaOCl is more efficient at dissolving necrotic tissue. Moreover, breaking down proteins into amino acids is essential in dissolving the organic part of the smear layer and biofilms, which, in turn, allows the irrigant to reach inaccessible areas in the root canal system. In addition to chloramination, NaOCl degrades fatty acids which transform into fatty acid salts (soap) and glycerol (alcohol). This saponification reaction reduces the surface tension of the solution allowing better penetration of dentinal tubules (Estrela et al., 2002). NaOCl also acts as a lubricant and flushes away all the debris packed inside the canal. Probably, two of its most important characteristics are that it is readily available and inexpensive. On the other hand, NaOCl also has some disadvantages. Prolonged irrigation using NaOCl affects the flexural strength of dentin (Grigoratos et al., 2001). Moreover, NaOCl accident is one of the complications that may encounter the clinician. To avoid this complication, it is recommended to use specially designed irrigation needles which should be inserted loosely instead of being wedged inside the root canal (Bettina and Markus, 2012). Allergic reactions to NaOCl are rare, but if present then the irrigant of choice is iodine potassium iodide (IKI) assuming the patient does not have an allergy to it as well. NaOCl is available in different concentrations varying from 0.5% to 6.25%. With lower concentrations, it can dissolve necrotic tissues only, while in higher concentrations, it dissolves both vital and necrotic tissue but it is more toxic (Hand et al., 1978; Spangberg et al., 1973). To avoid the negative impact of NaOCl toxicity, we can increase the effectiveness of a low concentration through different ways: G G G G Increase the volume of low-concentration NaOCl (Siqueira et al., 2000). Increase the temperature of low-concentration NaOCl (Zehnder, 2006). However, the efficacy of this method remains questionable since there are no clinical studies showing that heating NaOCl is effective (Bettina and Markus, 2012). Ultrasonic activation of NaOCl to create cavitational effects, preferably after the canal preparation (Macedo et al., 2014). Adding a detergent to the NaOCl solution to decrease surface tension which will enhance the penetration of NaOCl into the dentinal tubules and hence have a broader effect on eliminating bacteria (Stojicic et al., 2010). 12.3.2 Ethylenediamine tetra-acetic acid Ethylenediamine tetra-acetic acid (EDTA) is a polyaminocarboxylic acid (Fig. 12.2), colorless, and water soluble. EDTA activity lasts for extended periods
266 Advanced Dental Biomaterials Figure 12.2 Chemical structure of polyaminocarboxylic acid. Figure 12.3 Chemical structure of polybiquanide family. of time after direct exposure by extracting bacterial surface proteins and combining with metal ions present in the bacterial cell envelope which can eventually lead to cell death (Hulsmann et al., 2003). It also sequesters di- and tri-cationic metal ions such as Ca21 and Fe31. Furthermore, EDTA is a well-known chelator that reacts with the calcium ions in dentine and forms soluble calcium chelates. One of its major uses is to open occluded, very fine canals since it is capable of decalcifying dentin to a depth of 20 30 μm in 5 minutes. EDTA is commonly used in a concentration of 17%, which demonstrates limited antiseptic activity. However, since it removes the inorganic part of smear layers and biofilms within less than 1 minute after being in contact with root canal wall, it is often used synergistically with NaOCl (Goldman et al., 1976). Even though EDTA has a self-limited action (Hulsmann et al., 2003), it should be used with care following the manufacturer’s instructions as it can erode dentin if used for long periods of time. EDTA retains its ability to form calcium complexes when mixed with NaOCl. On the other hand, EDTA diminishes the tissue-dissolving capacity of NaOCl (Grawehr et al., 2003). Therefore it is suggested to use EDTA and NaOCl separately and copious amounts of NaOCl should be administered to wash out remnants of the EDTA. Overall, EDTA is preferably used at the end of a procedure to remove the smear layer. 12.3.3 Chlorhexidine Chlorhexidine (CHX) is a broad-spectrum antimicrobial agent and has been widely used as a disinfectant for a long time. It belongs to the polybiguanide antibacterial family, consisting of two symmetric four chlorophenyl rings and two biguanide groups connected by a central hexamethylene chain (Fig. 12.3). CHX is a positively charged hydrophobic and lipophilic molecule that interacts with phospholipids and lipopolysaccharides (LPS) on the negatively charged bacterial cell membrane and
Endodontic materials: from old materials to recent advances 267 penetrates the cell through a certain type of active or passive transport mechanism (Davies, 1973). This process will disrupt the microorganism’s permeability altering its equilibrium and making CHX antimicrobial. Unlike NaOCl, CHX overcomes some of the drawbacks of NaOCl including the bad odor, periapical reactions when extruded beyond the root apex, and bleaching spots on patient’s clothes. However, CHX does not possess the tissue-dissolving capacity of NaOCl. Hence, although NaOCl and CHX have comparable antibacterial effects, NaOCl dissolving capacity renders it more superior. CHX is effective against both Gram-positive and Gram-negative bacteria as well as yeasts (Block, 2001), although its activity against Gram-negative bacteria is not as strong as against Gram-positive bacteria. CHX is available as gel, liquid, or controlled-release devices. Some studies reported that CHX gel is more efficient than liquid. Nevertheless, the mode of application is not considered an important factor in the effectiveness of CHX; its concentration is the pivotal one (Basrani et al., 2003) and 2% CHX has better antibacterial efficacy than 0.12%. High concentrations are bactericidal since it damages the cell membrane which causes precipitation of the cytoplasm. While low concentrations are bacteriostatic, low molecular weight substances such as potassium and phosphorus will leak out without the cell being irreversibly damaged (Basrani and Lemonie, 2005). Substantivity is considered a unique characteristic for CHX (Basrani et al., 2002). Due to its cationic nature, it can be absorbed by anionic substrates present in the oral cavity, such as the apatite molecules present in teeth. Substantivity prolongs the antimicrobial effect of CHX, which contributes to delaying the coronal recontamination. Furthermore, CHX improves the resin dentin bond stability due to its broad-spectrum matrix metalloproteinase inhibitory effect (Carrilho et al., 2007). When CHX and NaOCl interact, they form para-chloroaniline (PCA) which is a neutral toxic insoluble precipitate that may interfere with the seal of obturation, causing color change (brownish-reddish) and cyanosis due to methemoglobin formation (Basrani et al., 2007). Table 12.4 presents some possible interactions between the different irrigants. 12.3.4 Citric acid Citric acid is an organic acid and a chelator used in concentrations between 10% and 50%. It is more potent as a chelator than EDTA at similar concentrations and is highly efficient in removing the smear layer and detaching bacterial biofilms (Smith and Wayman, 1986). Although EDTA is a weak antibacterial agent, it is stronger than citric acid in inhibiting bacterial growth (Siqueira et al., 1998). 12.3.5 MTAD MTAD is a mixture of 3% doxycycline hyclate, 4.25% citric acid, and 0.5% polysorbate-80 detergent (Tween 80) (Table 12.5). It is mixed as a liquid and powder and recommended as a final rinse for 5 minutes. MTAD is the first product that can both remove the smear layer and disinfect the root canal system. The whole
Table 12.4 Summary of the possible interactions between different irrigants. Interactions NaOCl EDTA NaOCl G G EDTA NaOCl loses its tissue-dissolving capacity, with virtually no free chlorine detected in combinations A bubbling effect reflects the chemical reaction but has no effect on cleansing PCA a precipitate that causes color change (brownishreddish) and may interfere with the seal of obturation Chlorhexidine EDTA retains its calciumcomplexing ability when mixed with NaOCl A bubbling effect reflects the chemical reaction but has no effect on cleansing White precipitate that involves the formation of salt rather than undergoing a chemical reaction G G Chlorhexidine PCA a precipitate that causes color change (brownishreddish) and may interfere with the seal of obturation White precipitate that involves the formation of salt rather than undergoing a chemical reaction EDTA, Ethylenediamine tetra-acetic acid; NaOCl, sodium hypochlorite; PCA, para-chloroaniline. Table 12.5 Composition of MTAD. Doxycycline G G G G G Isomer of tetracycline Broad-spectrum antibiotic Bacteriostatic by inhibition of protein synthesis In high concentrations it is bactericidal Has an effect in smear layer removal Citric acid G G Smear layer removal Addition of NaOCl is important to remove organic part of smear layer Tween 80 1. Main action is lowering surface tension 2. May increase antibacterial effect a. directly by affecting the bacterial cell membrane and b. indirectly through facilitating the penetration of MTAD into dentin 3. It may decrease antibacterial effect a. by being a nutrient to some bacteria b. by inactivating CHX and iodine effect if CHX and IKI were used as irrigants CHX, Chlorhexidine; IKI, iodine potassium iodide; MTAD, a mixture of doxycycline, citric acid, and a detergent; NaOCl, sodium hypochlorite.
Endodontic materials: from old materials to recent advances 269 mixture may have a synergistic effect on the disruption of the bacterial cell wall and on the cytoplasmic membrane (Torabinejad et al., 2003). MTAD components impart its mode of action which includes: G G G lowering the surface tension through Tween 80 which allows the irrigant to penetrate the dentinal tubules and inaccessible areas; smear layer removal with doxycycline and citric acid; and antibacterial efficacy due to doxycycline. 12.3.6 Tetraclean Tetraclean is similar to MTAD, but they differ in the concentration of antibiotics and the kind of detergent used. In MTAD, doxycycline has the concentration of 150 mg/5 mL, while the concentration of doxycycline in Tetraclean is 50 mg/5 mL. The detergent of MTAD detergent is Tween 80, while polypropylene glycol is used in Tetraclean (Giardino et al., 2005). 12.3.7 Hydrogen peroxide H2O2 is a clear and colorless liquid that is used widely as a disinfectant; however, it has a poor popularity among endodontists and it is not recommended to be used as a routine irrigant. H2O2 is available in concentrations ranging from 3% to 5% and produces hydroxyl-free radicals (HO2) that attack microbial components such as proteins and DNA. Moreover, it is active against bacteria, viruses, and yeasts but some bacteria produce catalase or superoxide dismutase which can provide them with partial protection against H2O2 (Bettina and Markus, 2012). 12.3.8 Iodine potassium iodide It is a halogen releasing agent consisting of 2% iodine in 4% potassium iodide (Siren et al., 2004). IKI is used as an irrigant due to its excellent antibacterial effect and low cytotoxicity (Spangberg et al., 1973). In addition, it is an oxidizing agent that cleaves the disulfide bonds present in the free sulfhydryl groups of bacterial enzymes. In the rare cases of hypersensitivity to NaOCl and CHX, IKI is considered the ideal irrigant due to its low cytotoxicity. Nevertheless, there are more allergy concerns linked to IKI compared to NaOCl and CHX. IKI also stains dentin which is one of its major shortcomings. 12.3.9 1-Hydroxyethylidene-1,1-bisphosphonate 1-Hydroxyethylidene-1,1-bisphosphonate (HEBP) is also called etidronic acid. It is used systematically in patients suffering from osteoporosis and Paget’s disease. As an irrigant, it is a chelator like EDTA and citric acid; however, it is considered a weak decalcifying agent and thus cannot be an effective final rinse. It is preferred to use HEBP in combination with (NaOCl), particularly since HEBP does not affect its proteolytic or antimicrobial properties (Zehnder, 2006).
270 Advanced Dental Biomaterials 12.3.10 QMiX QMiX was first introduced in 2011 as a new combination irrigation solution (Stojicic et al., 2012). It is recommended to be used as a final rinse after NaOCl irrigation and complete instrumentation. QMiX is available as a ready-to-use solution and contains a CHX-analog, triclosan (N-cetyl-N,N,N-trimethylammonium bromide), and EDTA. Since it contains CHX, it is recommended to rinse out the NaOCl from the canal with saline before using QMiX to prevent PCA formation. Moreover, it was shown that QMiX has comparable antibacterial properties and ability to remove the smear layer when compared to EDTA (Dai et al., 2011). Nevertheless, most studies on QMiX are ex vivo and in vivo, and further clinical studies are needed to ascertain the efficacy of QMiX. Table 12.6 describes some of the latest available irrigants. Table 12.6 Some of the newer available irrigants. Irrigant Smear Clear (SybronEndo) Description G G G G Chlor-XTRA (Vista Dental) CHX-Plus (Vista Dental) Hypoclean A and Hypoclean B HEBP G G G G G G G G G G G QMiX G G G G Clear and odorless Water soluble solution Water 1 17% EDTA 1 salts 1 a cationic surfactant (cetrimide) 1 anionic surfactant Adding detergents or surface-active agents is to increase the wettability of the solution and reduce the surface tension 6% NaOCl with surface modifiers Adding detergents or surface-active agents is to increase the wettability of the solution and reduce the surface tension 2% CHX with surface modifiers Adding detergents or surface-active agents is to increase the wettability of the solution and reduce the surface tension 5.25% NaOCl solutions modified with surfactants Adding detergents or surface-active agents is to increase the wettability of the solution and reduce the surface tension Used in combination with NaOCl with little interference with NaOCl activity Alternative to citric acid and EDTA Prevents bone resorption Used systematically in patients suffering osteoporosis and Paget’s disease Needs to be studied with further investigations Contains a CHX-analog, Triclosan (N-cetyl-N,N,N-trimethylammonium bromide), and EDTA Ready to use Used as a final rinse after NaOCl It lowers surface tension, removes smear layer, and has an antibacterial effect CHX, Chlorhexidine; EDTA, ethylenediamine tetra-acetic acid; HEBP, 1-hydroxyethylidene-1,1-bisphosphonate; NaOCl, sodium hypochlorite.
Endodontic materials: from old materials to recent advances 12.4 271 Intracanal medicaments It has been a trend among clinicians to use intracanal medicaments in cases where root canal treatment could not be completed in one visit. According to Athanassiadis et al. (2007) (Box 12.3), intracanal medicaments are necessary adjuncts to inhibit bacterial regrowth between appointments due to surviving bacteria and leaky temporary restorations. There are various materials that can be used as intracanal medicaments to enhance the antimicrobial effect of our cleaning procedures. Table 12.7 summarizes old intracanal medicaments that are not commonly used now. 12.4.1 Calcium hydroxide Calcium hydroxide has a particular importance in dentistry. It has been extensively used as an intracanal medicament since Hermann introduced it in 1920. It is a slurry of Ca(OH)2 in a water base but can be mixed with either water or saline and should have a thick consistency in order to carry as many Ca(OH)2 particles as possible. Chemically, calcium hydroxide is classified as a strong base with a high pH (approximately 12.5 12.8) that kills bacteria by direct contact (Safavi et al., 1990); hence, it should fill the entire length of the canal homogenously in order to exert its maximum biological effect. Fig. 12.4 summarizes the physical and chemical effects of Ca(OH)2. In general, the biological properties of calcium hydroxide are exerted through the ionic dissociation of Ca21 and OH2 ions, this contributes to a high pH which makes Ca(OH)2 antimicrobial (Bystrom and Sundqvist, 1985). It also can inactivate and hydrolyze bacterial LPS especially in necrotic cases (Safavi and Nichols, 1993). Promoting hard tissue formation is yet another significant action exerted by Ca(OH)2 (Frank, 1966). However, Ca(OH)2 has some limitations, for example, the rate of diffusion of hydroxyl ions is slowed due to the inherent buffering capacity of the dentine (Mohammadi and Dummer, 2011). There is controversy regarding the use of calcium hydroxide since it was shown that it does not reduce the number of bacteria significantly when used as an interappointment medicament. Moreover, 20% 45% residue is left on the canal walls even after irrigating with NaOCl or EDTA (Lambrianidis et al., 1999); this residue affects the sealing ability of zinc oxide eugenol (ZOE) sealers (Margelos et al., 1997). Box 12.3 Rationale of intracanal medicaments use G G G G G To eliminate or destroy any remaining viable bacteria in the root canal system that have not been destroyed by the chemomechanical preparation process (i.e., instrumentation and irrigation) To reduce periradicular inflammation and hence reduce pain To help eliminate apical exudate if it is present To prevent or arrest inflammatory root resorption if it is present To prevent reinfection of the root canal system by acting as both a chemical and a physical barrier if the temporary or interim restoration breaks down
272 Advanced Dental Biomaterials Table 12.7 Intracanal medicaments that are not recommended anymore. Phenol compounds Formaldehyde Examples Mode of action Disadvantages Several phenolic preparations have been used for disinfecting root canals such as phenol, camphorated phenol, paramonochlorophenol, thymol, cresol, camphorated parachlorophenol Applying those compounds through cotton pellets or paper points allows their vaporization. Those vapors come in direct contact with microorganisms once they penetrate the root canal system Formaldehyde is a volatile material that produces antimicrobial vapors when applied via a cotton pellet The antimicrobial effect accompanied with those compounds does not balance the cytotoxicity caused by them. For that reason, their use has been decreased significantly (Spangberg et al., 1973) It is cytotoxic and has a mutagenic effect. Its use as an antimicrobial medicament is not justified due to its toxicity (Spangberg et al., 1973; Lewis and Chestner, 1981) It can stain dentin and is considered a severe allergen G G Halogens G G G Formocresol: formaldehyde (ranges between 19% and 37%) Tricresol formalin: formaldehyde component is 90% and contains 10% tricresol Chloramine-T N-Chloro tosylamide sodium salts IKI can be used as both an irrigant and intracanal medicament It releases antimicrobial vapors which can disinfect dentin in 5 min and has low toxicity (Safavi et al., 1990) IKI, Iodine potassium iodide. In addition to being used as an intracanal medicament, Ca(OH)2 is used in apexification procedures for immature teeth before obturating the canal with gutta-percha (GP). Moreover, it has been recently used as a substitute for triple antibiotic pastes in revascularization procedures of immature teeth (Law, 2013). 12.4.2 Chlorhexidine Please refer to CHX in section 12.3.3 for more details. CHX has a substantive antimicrobial activity that prevents bacterial colonization for prolonged periods of time,
Endodontic materials: from old materials to recent advances 273 Figure 12.4 The physical and chemical effects of Ca(OH)2. and this property is behind the rationale for using CHX as an intracanal medicament in a concentration of 2%. 12.4.3 Ledermix Schroeder first developed this material as a combination of antibiotic/corticosteroids paste. Ledermix paste is a nonsetting, water-soluble paste material with a combination of a tetracycline antibiotic, demeclocycline HCl (at a concentration of 3.2%), and a corticosteroid, triamcinolone acetonide (concentration 1%), in a polyethylene glycol base. In patients experiencing severe endodontic pain, Ledermix (Riemser Arzneimittel AG, Insel Riems, Germany) can sometimes be used as an initial medicament in order to alleviate the symptoms (Schroeder, 1975). Furthermore, some studies have shown that Ledermix can inhibit root resorption and promotes favorable root healing (Chen et al., 2008). However, there is still controversy regarding products containing corticosteroids due to concern related to their release to the blood circulation (Abbott, 1992). 12.4.4 Triple antibiotics pastes As the name indicates, this medicament is composed of three antibiotics in one paste (metronidazole, ciprofloxacin, and minocycline). This paste can be effective in eradicating bacteria and has also been used in regenerative endodontics (Windley et al., 2005). Indeed, it was shown clinically that it promoted root formation in young permanent teeth with an open apex despite the presence of a periapical radiolucency and a draining buccal sinus tract (She et al., 2016). It should be noted that there are still concerns related to its use, such as tooth discoloration from minocycline and bacterial resistance.
274 Advanced Dental Biomaterials 12.4.5 Bioactive glass Bioactive glass consists of a SiO2 Na2O CaO P2O5 system (Zehnder et al., 2004). It is considered a promising material due its antimicrobial effect and osteoinductive ability. Its antibacterial mechanism of action is achieved through its high pH, its ability to increase the osmotic pressure above 1% which can inhibit many bacteria, and through inducing mineralization on bacterial surfaces. 12.5 Root canal obturation materials Throughout the history of endodontics, several materials were used as root filling materials ranging from tin foil, cotton pellets, plaster of Paris, to leaving the root canal unfilled. Grossman outlined in 1987 the properties of an ideal obturation material and provided criteria that can help identify the optimal obturation material (Box 12.4). After proper chemomechanical debridement, the disinfected root canal should be sealed in order to prevent reinfection. Accordingly, obturation materials should seal the root canal: G G G apically to block the exit of surviving microorganisms to the periradicular tissue; coronally to prevent reinfection of the pulp space from the oral cavity; and throughout the canal to seal the openings of accessory and lateral canals. The obturation step involves the use of two types of materials (Fig. 12.5): one is the core obturation material in the solid form (e.g., GP and resilon) and another semisolid cementing medium (i.e., sealers). 12.5.1 Core obturation materials 12.5.1.1 Silver points (or cones) Jasper introduced silver points in 1933 claiming that they are as successful as GP. Silver points are rigid and have the same diameter and taper as the instruments Box 12.4 Properties of an ideal obturation material G G G G G G G G G G G Easily manipulated and provides ample working time Dimensionally stable with no shrinkage once inserted Seals the canal laterally and apically, conforming to its complex internal anatomy Nonirritating to the periapical tissues Impervious to moisture and nonporous Unaffected by tissue fluids—no corrosion or oxidization Inhibits bacterial growth Radiopaque and easily discernible on radiographs Does not discolor tooth structure Sterile Easily removed from the canal if necessary
Endodontic materials: from old materials to recent advances 275 Figure 12.5 Diagrammatic illustration of the root canal fillings. Table 12.8 Advantages and disadvantages of silver points. Advantages of silver points G G G G Rigid Have the same taper and diameter as the files used to shape the canal Easily placed inside the canal More predictable length control Disadvantages of silver points G G G G Nonadaptable and unable to fill canal irregularities and this permits leakage It corrodes when in contact with tissue fluids or saliva The cytotoxic corrosion by-products impair periapical healing Difficult to remove if retreatment is indicated used to shape the canal, this property facilitated their placement and allowed a more predictable length control. Silver points possess most of the properties Grossman identified, yet their rigidity precludes proper adaptation to the root canal, and thus inability to form a proper seal apically and laterally. Table 12.8 lists the main advantages and disadvantages of silver points. Furthermore, the leakage of the sealers caused the corrosion of silver points and formation of cytotoxic silver salts (Brady and del Rio, 1975). With those major disadvantages of silver points their use nowadays is considered as below the standard of care. 12.5.1.2 Gutta-percha GP is considered the most widely used obturation material. It is derived from various species of rubber trees from Malaysia, Borneo, Indonesia, and South America, mainly from Brazil. It is the trans-isomer of 1,4-polyisoprene (rubber) and it exists in two crystalline forms (α and β) (Goodman et al., 1974) (Fig. 12.6). The β phase is the unheated solid form that can be compacted. When heated, it transforms into the α phase which exhibits superior thermoplastic properties since it becomes pliable and can flow when pressure is applied. This characteristic is utilized in heat-softened GP obturation techniques (e.g., Thermafil and Obtura) and is now considered the gold standard.
276 Advanced Dental Biomaterials Figure 12.6 cis and trans isomers of 1,4-polyisoprene (natural rubber and gutta-percha). Box 12.5 Ideal properties of a root canal irrigant G G G G 59% 76% Zinc oxide 18% 22% Gutta-percha 1% 18% Heavy metals salts 1% 4% Wax or resin Table 12.9 Advantage and disadvantages of gutta-percha points. Advantages G G G G G G Plasticity Ease of manipulation Minimal toxicity (inert) Radiopaque Can be removed with heat or solvents Adapts with compaction to canal irregularities Disadvantages G G G G Lack of adhesion to dentin Shrinkage on cooling (1% 2%) Elasticity (rebounds, pulls away) Oxidizes to air and light and becomes more brittle According to Friedman et al. (1975), GP is not the main constituent of these obturation points and contains zinc oxide as a filler, wax, and resins as plasticizers and radiopacifier heavy metals such as barium sulfate (Box 12.5). It satisfies most of the properties identified by Grossman; however, its main disadvantage is that it lacks adequate rigidity. Another disadvantage is its inability to seal without a sealer. Table 12.9 summarizes the advantages and disadvantages of GP. GP is available in standardized International Standards Organization (ISO) sizes and nonstandardized sizes such as extrafine, fine fine, fine, medium fine, fine medium, medium, large, and extralarge. Initially, it was fabricated in the standard 0.02 taper, but with the introduction of newer shaping instruments with larger tapers, GP points are manufactured to match those instruments. For instance, a fine medium cone means that the cone has a fine tip with a medium body.
Endodontic materials: from old materials to recent advances 277 GP can be sterilized by 5.25% NaOCl for 1 minute, then rinsed with ethyl alcohol to remove crystallized NaOCl, which, if not removed, would impair the obturation seal. It is important to note that it cannot be heat-sterilized. GP oxidizes to air and light and becomes brittle; consequently, storage in a cool dry place such as refrigerators is imperative. Moreover, it can be softened by chloroform, halothane, and to a lesser degree by turpentine and xylene. 12.5.1.3 Resilon Resilon is a polyurethane thermoplastic root canal filling material. It is used with certain sealers such as Epiphany, RealSeal, and Resinate. The introduction of this material aimed to achieve what is known as “monoblock” where the resilon sealer (epiphany) bonds to both the etched canal wall and the resilon cone, but whether this can be truly achieved is still controversial (Tay and Pashley, 2007; Raina et al., 2007). Moreover, the use of this bonding resin does not strengthen the tooth. Resilon consists of the following (Shipper et al., 2004): G G G Methacrylate resin (difunctional) bisphenol A glycidyl methacrylate (bis-GMA). Bioactive glass. Radiopaque fillers (bismuth oxychloride and barium sulfate) make up 65% of its content. Similar to GP, Resilon is manufactured in the same ISO sizes and shapes; it can also be softened using heat or chloroform and can be used with any obturation technique (lateral condensation, thermoplasticized, carrier based). Furthermore, the sealing ability of Resilon is comparable to GP (Baumgartner et al., 2007). In addition, Resilon is nontoxic, nonmutagenic, and biocompatible (Emre et al., 2008). It is also compatible to be used with cores and posts that are resin placed. The Resilon system consists of a primer, a sealer, and synthetic polymer points or pellets. Preparatory steps are needed to use this material, such as using a SelfEtch Primer. This primer consists of sulfonic acid-terminated functional monomer, 2-hydroxyethyl methacrylate, water, and polymerization initiator. The primer is applied (two to three drops) with a pipette or syringe or paper point. It is essential to know that NaOCl affects the bond of the primer so it is mandatory to use EDTA as a final rinse before rinsing the canal with sterile water, saline, or CHX. Afterward, the excess primer is removed, and the resin sealer is mixed and applied to a paper point, lentulo spiral, or Resilon point. Eventually, the canal is filled with any obturation technique. 12.5.2 Root canal sealers (cementing medium) Since root filling materials do not bond to the dentinal wall of the root canal, sealers are crucial for sealing the space between the two surfaces and fill the voids and irregularities in the root canal anatomy during obturation. In 1982 Grossman outlined the properties of an ideal sealer (Box 12.6). However, currently, no sealer satisfies all these characteristics. Different kinds of sealers are present, each with their advantages and disadvantages.
278 Advanced Dental Biomaterials Box 12.6 Properties of an ideal root canal sealer G G G G G G G G G G G Exhibits tackiness when mixed to provide good adhesion between it and the canal wall when set Can establish a hermetic seal Radiopaque Made of a very fine powder, so that it can mix easily with liquid Exhibits a slow set Does not undergo shrinkage on setting Does not stain the tooth structure Bacteriostatic, or at least does not encourage bacterial growth Insoluble in tissue fluids Tissue tolerant, that is, nonirritating to periradicular tissue Soluble in a common solvent if it is necessary to remove the root canal filling 12.5.2.1 Zinc oxide eugenol sealers It is the most commonly used sealer so far. There are various product brands that contain ZOE as the main ingredient and only differ in the other components added to the sealer. The advantages of ZOE sealers are its antimicrobial activity (Heling and Chandler, 1996), slow setting time (Allan et al., 2001), and it undergoes resorption when extruded into the periradicular tissues (Augsburger and Peters, 1990). Nevertheless, some formulas of ZOE sealers can stain the tooth structure and shrink on the setting (Krastl et al., 2013; Kazemi et al., 1993). Table 12.10 summarizes different ZOE sealers. One characteristic of ZOE sealers is the ability of the constituents of the material to enhance its properties. For instance, paraformaldehyde was added for antimicrobial and mummifying effects; however, it is not acceptable due to its severe adverse effects. When corticosteroids are added to the sealer, they can suppress inflammatory reactions. Canada Balsam can also be added for a more sticky and tacky properties that would improve the sealer’s adhesion to dentin. Rosins (colophony), which are derived from conifers, were initially added for better adhesive properties (Grossman, 1982), but it was later found that resins acids which form 90% of rosins might be cytotoxic, yet the addition of resin acids to zinc oxide may be beneficial. Recent research is trying to add nanoparticles such as chitosan and quaternary ammonium polyethylenimine to enhance the antibacterial effectiveness of the sealer (Shrestha and Kishen, 2016). Chitosan [poly(1,4-b-D-glucopyranosamine)] is a long polymer chain which gains its effectiveness from its chemical structure. By adding chitosan nanoparticles, the antibacterial effect of the sealer is increased without deteriorating its flow characteristics. However, more research is required for the incorporation of nanoparticles into dental materials. 12.5.2.2 Calcium hydroxide sealers This sealer was introduced to exploit the antibacterial and hard tissue formation properties of calcium hydroxide. Nevertheless, those properties could not be
Endodontic materials: from old materials to recent advances 279 Table 12.10 Zinc oxide eugenol (ZOE) sealers. Rickert’s formula Pulp Canal Sealer (SybronEndo) Pulp Canal Sealer EWT Proscol (Procosol, Inc., Philadelphia, PA) Roth’s Sealer (Roth’s Pharmacy, Chicago, IL) Wach’s sealer (Balas Dental, Chicago, IL) Tubliseal 1. The earliest formula of ZOE 2. It satisfied most of Grossman’s properties except that it caused staining to tooth tissues due to silver particles 1. It followed Rickert’s formula 2. It exhibited very rapid setting time 3. This sealer used extensively by clinicians in thermoplastic warm vertical obturation 1. It is a modification of Pulp Canal Sealer 2. It was marketed as having a 6 h working time 3. This sealer is used extensively by clinicians in thermoplastic warm vertical obturation 1. It is a modification of Rickert’s formula by removing the silver particles to make a nonstaining sealer 2. It contains zinc oxide, hydrogenated staybelite resin, bismuth subcarbonate, and barium sulfate 1. It is a modification of Rickert’s formula by removing the silver particles to make a nonstaining sealer 2. It contains zinc oxide, hydrogenated staybelite resin, bismuth subnitrate, and barium sulfate 1. Its powder contains zinc oxide, bismuth subnitrate, bismuth subiodide, magnesium oxide, and calcium phosphate, while the liquid consists of oil of cloves, eucalyptol, Canada balsam, and beechwood creosote 2. Canada balsam is what makes this sealer different than other sealers when used in lateral compaction due to its tacky property that softens the gutta-percha into a more homogeneous mass 1. Unlike the previous powder/liquid formulas, it is a catalyst/base sealer 2. It is known to be easily mixed with a two-paste system in two separate tubes. The paste contains zinc oxide-base, barium sulfate for radiopacity, and mineral oil, cornstarch, and lecithin. The catalyst tube has polypale resin, eugenol, and thymol iodide 3. Tubliseal has the disadvantage of having a rapid setting time and that it is modified in Tubliseal EWT EWT, Extended working time. demonstrated (Desai and Chandler, 2009). Furthermore, in order to release significant amounts of calcium hydroxide, the sealer should exhibit high solubility properties, which is inconsistent with the optimum properties of sealers. SealApex, Calciobiotic Root Canal Sealer, and Apexit have calcium hydroxide as one of their main ingredients. Table 12.11 describes the composition of the different calcium hydroxide sealers.
280 Advanced Dental Biomaterials Table 12.11 Composition of different calcium hydroxide Ca(OH)2 sealers. Type of Ca(OH)2 sealer SealApex Ingredients G G CRCS G G Apexit G G Base: Zinc oxide, calcium hydroxide, butyl benzene, sulfonamide, and zinc stearate Catalyst: Barium sulfate and titanium dioxide for radiopacity, a proprietary resin, isobutyl salicylate, and aerosil R792 Powder: Calcium hydroxide, zinc oxide, bismuth dioxide, and barium sulfate Liquid: Eugenol and eucalyptol Base: Calcium hydroxide, zinc oxide, calcium oxide, silicon dioxide, zinc stearate, hydrogenized colophony, tricalcium phosphate, and polydimethylsiloxane Activator: Trimethyl hexanediol diasalicylate, bismuth carbonate basic, bismuth oxide, silicon dioxide, 1,3-butanedioldisalicylate, hydrogenized colophony, tricalcium phosphate, and zinc stearate CRCS, Calciobiotic Root Canal Sealer. 12.5.2.3 Glass ionomer sealers Glass ionomers are capable of binding to dentin, and thus, they were incorporated in certain endodontic sealers, such as Ketac-Endo (3M, Minneapolis, MN) (Friedman et al., 1995). However, the disadvantage of this sealer is its limited antimicrobial effect (Heling and Chandler, 1996), the difficulty in conditioning the middle and apical parts of the root canal to receive the glass ionomer sealer, and the difficulty of removing the sealer in cases of retreatment. Glass ionomer sealers are usually used with Activ GP (Brasseler USA, Savannah, Georgia) which is composed of GP impregnated with glass ionomer. 12.5.2.4 Resin sealers Resin sealers are popular for their favorable adhesive properties and lack of eugenol. These sealers are either epoxy-resin sealers (e.g., AH-26 and AH-plus) or urethane methacrylate resin sealers (e.g., EndoREZ and Epiphany). AH-26 (Dentsply International Maillefer) is a bisphenol epoxy resin sealer that uses hexamethylenetetramine (methenamine) for polymerization. In addition to its adhesive properties, AH-26 demonstrates increased radiopacity, low solubility, low polymerization shrinkage, and tissue compatibility. However, the methenamine releases some formaldehyde as it sets, which is the major disadvantage of using this sealer (Spangberg et al., 1993). Moreover, tooth staining and extended working time are other disadvantages. AH-plus (Dentsply International) is a modification of AH-26 with comparable sealing properties without the release of formaldehyde (De Moor and De Bruyne, 2004). Compared to AH-26, the AH-plus sealer exhibits lower solubility, higher radiopacity, and shorter working time (B4 hours) and setting time compared to AH-26. As for methacrylate resin sealers, four generations of endodontic sealers have been marketed so far summarized in Table 12.12 (Kim et al., 2010).
Table 12.12 Methacrylate resin sealers. Methacrylate resin sealer generation Major components Main characteristics First generationHydron (Hydron Technologies, Inc., Boca Raton, FL) Poly(2-HEMA) G G Hydrophilic Injectable without the use of root filling core material Advantages G G G G G Second generationEndoREZ (Ultradent Products Inc, South Jordan, UT) G G Nonacidic diurethane dimethacrylate In addition to triethylene glycol dimethacrylate G G G G G G G Nonetching Radiopaque Dual cured Hydrophilic in nature Can be used with conventional gutta-percha or specific EndoREZ points which are resin coated gutta-percha cones Seals best when applied to moist intraradicular dentin No adjunctive use of a dentin adhesive. However, an accelerator has been introduced for rapid cure G G Injectable which makes it easy to use Nonirritating Highly adaptable to the canal walls Nonsupportive of bacterial growth Calcification ability in the event of inadvertent extrusion of the sealer into the periapical regions Retentive due to the formation of resin tags when it flows into the accessory canals and dentinal tubules especially after the removal of the smear layer with NaOCl and EDTA It might be used in the wet environment of the root canal system Disadvantages G G G G G G Smear layer removal Severe inflammatory reaction Absorption of the material Severe leakage Water sorption and swelling Low bond strength to the dentinal wall when used with conventional uncoated gutta-percha Studies that show its superiority lacked the use of additional sealers for comparison Smear layer removal by a final rinse of NaOCl and EDTA is very important for the formation of resin tags (Continued)
Table 12.12 (Continued) Methacrylate resin sealer generation Major components Main characteristics Third generation: selfetching sealers FibreFill R.C.S. root canal sealant (Pentron Clinical Technologies, Wallingford, CT) Epiphany (Pentron Clinical Technologies) RealSeal (SybronEndo, Orange, CA), Resinate (Obtura Spartan Corp, Fenton, MO) Smart (Discus Dental, Culver City, CA) Contains a selfetching primer and a dual-cured resin composite root canal sealer Using an acidic self-etching primer brought back the concept of incorporating the smear layer formed by hand/rotary instrumentation along the sealer dentin interface. The primer is applied to the surface of dentin to penetrate the smear layer and demineralize the superficial dentin and then it is air-dried to remove any volatile carrier 1. The dual cure flowable resin sealer is applied after the primer and polymerized G G G G Advantages G G Bonding procedure is simplified The self-etching primers in some sealers are even reduced from a two-bottle system to a single-bottle system. These self-etching primers contain AMPS as the functional acidic. This shorten the time needed before bonding Disadvantages Smear layer removal Bonding apically may be compromised if the primer was not aggressive enough to penetrate the smear layer in the apical third The primer is applied to the surface of dentin and able to penetrate through the smear layer and to demineralize the superficial dentin
Fourth generation: MetaSEAL, Parkell Inc; RealSeal SE, SybronEndo or Hybrid Bond SEAL (Sun Medical Co Ltd, Shiga, Japan) MetaSEAL or Hybrid Bond SEAL are composed of acidic resin monomers that are usually found in dentin adhesive primers are integrated into the resin-based sealer which makes them self-adhesive 1. 4-META which makes the sealer self-etching, hydrophilic and promotes monomer diffusion into the underlying intact dentin to produce a hybrid layer after polymerization G G Eliminated the separate etching/bonding step Those sealer creates a hybrid layer in both the thermoplastic root filling and the radicular dentin which promotes their bonding G G G G G The fact that the etchant, primer, and sealer form an all-inone self-etching, selfadhesive sealer reduces the application time and the errors that usually occur during the bonding stage Self-adhesive Dual-cured sealer Hydrophilic Self-etching G G Studies have shown similar or slightly inferior sealing properties as conventional nonbonding epoxy resinbased sealers Limited aggressiveness of hybridizing radicular dentin especially apically and inaccessible areas of the root call like isthmi It is recommended by the manufacturers to use EDTA as a final rinse when using methacrylate resin-based sealers in order to remove the smear layer, reduce the leakage and improve the seal of filled canals 4-META, 4-Methacryloyloxyethyl trimellitate anhydride; AMPS, 2-acrylamido-2-methyl-propanesulfonic acid; EDTA, ethylenediamine tetra-acetic acid; HEMA, hydroxyethyl methacrylate; NaOCl, sodium hypochlorite.
284 Advanced Dental Biomaterials 12.5.2.5 Bioceramic-based sealers These sealers are currently gaining a wide popularity due to their outgrowth from MTA and the extensive research and technology directed toward enhancing the use of bioceramics in endodontics. MTA and MTA-like materials are made of tricalcium silicate and their importance rises from their hydrophilic properties, excellent sealing ability, and their bioactivity (Al-Haddad and Che Ab Aziz, 2016). Table 12.13 summarizes the available bioceramic-based sealers and a few newer ones that are still in the experimental phase. 12.5.2.6 Other sealers Other types of sealers include noneugenol sealers, silicone sealers, chloropercha, and sealers containing formaldehyde. For more details about these sealers refer to Table 12.14 which summarizes the main sealers used in endodontics. Table 12.13 Bioceramic-based sealers. Bioceramic-based sealer MTA Fillapex (Angelus Industria de Produtos Odontologicos Ltda; Londrina, Parana, Brazil) iRoot SP (Innovative BioCeramix Inc., Vancouver, BC, Canada) aka Endosequence BC sealer (Brasseler USA, Savannah, GA) Endo CPM Sealer (EGEO SRL, Buenos Aires, Argentina) MTA Plus (Avalon Biomed, Bradenton, FL) and (Prevest Denpro, Jammu City, India) Description G G G G G G G G A resin MTAbased root canal sealer Available as a dual paste that are combined in a mixing tip A bioactive alkaline calcium silicate sealer Premixed readyto-use single paste An MTA-based root canal sealer Powder/gel A calcium silicate sealer Powder/gel Composition Natural resin, salicylate resin, diluting resin, bismuth trioxide, nanoparticulated silica, MTA, and pigments G G Zirconium oxide, calcium silicates, calcium phosphate monobasic CH and filling and thickening agents MTA in addition to calcium chloride, calcium carbonate, sodium citrate, propylene glycol alginate, and propylene glycol MTA powder with a finer particle size Two formulations are available either mixed with water or a hydrosoluble gel for decreasing washout property G G (Continued)
Endodontic materials: from old materials to recent advances 285 Table 12.13 (Continued) Bioceramic-based sealer Generex B ProRoot Endo Sealer (Dentsply Tulsa Dental Specialties, Tulsa, OK, United States) Description G G A calcium silicate sealer Mixed with a liquid-to-powder ratio of 1:2 Composition The powder is composed of tricalcium silicate, dicalcium silicate, calcium sulfate (as a setting retardant), bismuth oxide (as a radiopacifier), and a small amount of tricalcium aluminate The liquid is composed of water and a viscous watersoluble polymer Two formulas are available: The first one was introduced by Camilleri and associates consist of a mixture of 80% white Portland cement and 20% bismuth oxide The second one was introduced from Brazil and is composed of Portland cement, zirconium oxide (radiopacifier), calcium chloride, and a resinous vehicle Powder: α-tricalcium phosphate and hydroxySankin apatite in type I, iodoform added to powder in type II (30%) and type III (5%) Liquid: polyacrylic acid and water Powder: TTCP and DCPA, Portland cement (gray cement in type I and white cement in type II), zirconium oxide, and others Liquid: hydroxypropyl methyl cellulose in sodium phosphate solution G G Mineral Trioxide Aggregate Sealer (MTAS) Powder-to-liquid ratio of 5:3 by weight G G Sankin apatite root canal sealer (I III), Sankin Kogyo, Tokyo, Japan Calcium phosphate based root canal sealer G G G G Capseal (I and II) Experimental Calcium phosphate based root canal sealer G G CH, Calcium hydroxide; DCPA, dicalcium phosphate anhydrous; MTA, mineral trioxide aggregate; TTCP, tetracalcium phosphate.
Table 12.14 Root canal sealers. Description Zinc oxide eugenol cement sealer G G G G G G Calcium hydroxide sealers Most commonly used sealer Modification of the regular ZnOE The mixing vehicle is mostly eugenol Powder contains ZnO that has been finely sifted to enhance the flow Setting process is a chemical reaction with physical embedding of ZnO in a matrix of zinc eugenolate Free eugenol always remains in the mass as an irritant Advantages G G Absorbed if extruded into periradicular tissue Antimicrobial effect through ZnO or additives such as rosin and Canada Balsam or corticosteroids Disadvantages G G G Developed from periodontal dressing It is a catalyst/base sealer with a base of zinc oxide, barium sulfate, bismuth oxychloride, and vegetable oil, and the catalyst containing hydrogenated rosin, methyl abietate, lauric acid, chlorothymol, and salicylic acid G G G G G G G Noneugenol sealers G G It was developed for two main reasons: 1. Antimicrobial effect 2. Osteogenic cementogenic potential However, these two effects were not demonstrated G Shrinks on setting Stains the tooth (Roth’s Sealer was developed to be as a nonstaining ZnOE sealer) Examples No irritating effect of eugenol Needs to be soluble in order to leach out the hydroxyl group responsible for its activity but sealers shouldn’t exhibit solubility Poor cohesive strength Poor dentin adhesion G G G Pulp Canal Sealer (SybronEndo) Pulp Canal Sealer EWT Proscol (Procosol, Inc., Philadelphia, PA) Roth (Roth’s Pharmacy, Chicago, IL) Tubliseal (SybronEndo/ Kerr) Wach’s Sealer (Balas Dental, Chicago, IL) CRCS (Coltène/ Whaledent) Sealapex (SybronEndo) Apexit (Ivoclar Vivadent) Nogenol (GC America, Alsip, IL)
Glass ionomer sealers Developed to take advantage of their bond to dentin, fluoride release, antimicrobial activity, and biocompatibility G G G Dentin binding Minimal tissue irritation Low toxicity G G G G Resin sealers G G G Introduced because it provides good adhesion AH-26 and AH-plus are epoxy resins Methacrylate resin sealers were introduced in four generations G G Good adhesion ability Does not contain eugenol G G Solubility concern Inadequate bonding with guttapercha due to difficult conditioning of the middle and apical thirds of the root canal Minimal antibacterial effect Difficulty in case of retreatment AH-26 contains formaldehyde which is toxic when freshly mixed but this toxicity declines rapidly after 24 h. This was modified by the introduction of AH-plus that does not release formaldehyde Some of the disadvantages of methacrylate resin sealers are polymerization shrinkage, failure of bonding at the dentin sealer interface, presence of residual monomers, cytotoxicity in case extrusion outside the canal, difficulty of removal in cases of retreatment and its low cohesive, tensile, compressive strengths in comparison with dentin which indicates their inability to enhance the mechanical properties of root canal dentin G G G G G G G G G G G Ketac-Endo (3M ESPE, Minneapolis, MN) Activ GP (Brasseler USA, Savannah, GA) which is a gutta-percha cone impregnated with GI coating AH-26 (Dentsply International Maillefer) AH-Plus (Dentsply International) EndoREZ (Ultradent Products Inc, South Jordan, UT) FibreFill R.C.S. root canal sealant (Pentron Clinical Technologies, Wallingford, CT) Epiphany (Pentron Clinical Technologies) RealSeal (SybronEndo, Orange, CA) Resinate (Obtura Spartan Corp, Fenton, MO) Smart (Discus Dental, Culver City, CA) MetaSEAL (Parkell Inc; RealSeal SE, SybronEndo); Hybrid Bond SEAL (Sun Medical Co Ltd, Shiga, Japan) (Continued)
Table 12.14 (Continued) Description Advantages Disadvantages Examples Silicone-based sealers Two materials are available: 1. RoekoSeal (Coltène/Whaledent): which is a polyvinylsiloxane or polydimethylsiloxane root canal sealer and which exhibits slight expansion on setting (Orstavik et al., 2001) 2. GuttaFlow (Coltène/Whaledent) a. Cold, flowable, self-curing obturation material for root canals that combines gutta-percha and polydimethylsiloxane sealer (RoekoSeal) into one injectable system b. Fills canal irregularities (Zielinski et al., 2008) c. Inject the material in the canal and then place a single master cone d. Working time is 15 min and cures in 25 30 min e. Available in capsules to be triturated f. Delay of setting time with final irrigation of NaOCl (Bouillaguet et al., 2006) g. Biocompatible (Bouillaguet et al., 2006) Chloropercha White gutta-percha (Alba) mixed with chloroform Not used anymore Composed of zirconium oxide and calcium silicates. And some have calcium phosphate as well G G Bioceramicbased sealers The addition of chloroform allows the gutta-percha cone to fit better in the canal Hydrophilic Sets by utilizing moisture, the reaction of calcium silicates in the presence of the dentin’s moisture, will form hydroxyapatite and this will enhance the sealing ability of the sealer No shrinkage on setting Antimicrobial effect during setting Biocompatible G G No adhesive ability Toxicity of chloroform (Zakariasen et al., 1990) Hardens on setting so retreatment might be difficult to perform if required G G G G G G G G MTA Fillapex (Angelus Industria de Produtos Odontologicos Ltda; Londrina, Parana, Brazil) Generex B ProRoot Endo Sealer (Dentsply Tulsa Dental Specialties, Tulsa, OK, United States) iRoot SP (Innovative BioCeramix Inc., Vancouver, BC, Canada) aka
G G G G G Sealers containing formaldehyde Absolutely contraindicated G G Toxic Overextension causes irreversible neurotoxicity dysesthesia (Kleier and Averbach, 1988) G G G G CRCS, Calciobiotic Root Canal Sealer; EWT, extended working time; GI, glass ionomer; GP, gutta-percha; MTA, mineral trioxide aggregate; NaOCl, sodium hypochlorite. Endosequence BC sealer (Brasseler USA) Endo CPM Sealer (EGEO SRL, Buenos Aires, Argentina) MTA Plus (Avalon Biomed, Bradenton, FL) and (Prevest Denpro, Jammu City, India) MTAS Sankin apatite root canal sealer (I, II, and III) Sankin Kogyo, Tokyo, Japan Capseal I and II Endomethasone (Septodont, Paris, France) Rieber’s paste (Amubarut; Wera Karl, Biesingen, Germany) N2 (Indrag-Agsa, Losone, Switzerland) SPAD
290 Advanced Dental Biomaterials 12.6 Root-end filling materials Periradicular surgery is an indispensable part of endodontics. When orthograde retreatment does not eradicate the periradicular pathosis, root-end resection is required followed by applying a biocompatible root-end filling material for proper seal. Many root-end filling materials are available in the market, but none of them verify all the requirements of an ideal root-end filling as summarized by Chong and Pitt Ford (2005) (Box 12.7). 12.6.1 Amalgam Amalgam is a radiopaque material and can be manipulated easily. Nevertheless, there are many disadvantages that are associated with its use as a root-end filling, such as tissue biocompatibility, ineffective seal, amalgam tattoo appearance, corrosion, and dimensional changes (Dorn and Gartner, 1990). Hence, the use of amalgam as a root-end filling is decreasing significantly. 12.6.2 Zinc oxide eugenol cements These cements are made of the basic ZOE mixture with the addition of other materials that can counteract its shortcomings such as solubility and reduced strength (refer Table 12.15). 12.6.3 Composite resins (Retroplast) Retroplast is a radiopaque dentin-bonding composite resin. It is provided in a twopaste system that is mixed in equal parts to form a dual cure composite resin with a working time of 1.5 2 minutes. Paste A is composed of bis-GMA/TEGDMA 1:1, benzoyl peroxide N,N-di-(hydroxyethyl)-p-toluidine, and butylated hydroxytoluene Box 12.7 Requirements of an ideal root-end filling material G G G G G G G G G G G Should adheres or bond to tooth tissue and “seal” the root-end three dimensionally Should not promote, and preferably inhibit, the growth of pathogenic microorganisms Dimensionally stable and unaffected by moisture in either the set or unset state Well tolerated by periradicular tissues with no inflammatory reactions Stimulates the regeneration of normal periodontium Nontoxic both locally and systemically Should not corrode or be electrochemically active Should not stain the tooth or the periradicular tissues Easily distinguishable on radiographs Has a long shelf life Easy to handle
Endodontic materials: from old materials to recent advances 291 Table 12.15 Root-end filling materials—zinc oxide eugenol (ZOE) cements (Dorn and Gartner, 1990; Walivaara et al., 2009). Root-end filling material Chemical composition Description IRM (Caulk/ Dentsply, Milford, DE) Powder/liquid mixed in equal parts. However, a higher powder-to-liquid ratio of IRM is recommended, forming a consistency with easier placement, shorter setting time, decreased toxicity and solubility Powder: 75% zinc oxide and 20% 25% polymethacrylate Liquid: 99% eugenol and less than 1% acetic acid It is a reinforced ZOE cement with polymethacrylate added to it. Since eugenol is the major cytotoxic component in ZOE cement, polymethyl methacrylate will limit its release because of eugenol’s high affinity to it. It has better sealing ability than amalgam and it is well tolerated by the periapical tissue but without hard tissue regenerative capabilities (Crooks et al., 1994) Super EBA is another reinforced ZOE cement. It comes in two forms, regular set and fast set. Super EBA is well tolerated by the periradicular tissues yet with no capacity to regenerate cementum. It has much better physical properties than ZOE such as high compressive strength and better tensile strength. A good healing response was observed with super EBA with minimal chronic inflammation at the root apex and less apical leakage G G Super EBA (Bosworth, Skokie, IL) Powder/liquid mixed in equal parts Powder: 60% zinc oxide, 34% aluminum oxide, and 6% natural resin Liquid: 37.5% eugenol and 62.5% o-methoxybenzoic acid G G IRM, Intermediate restorative material; super EBA, super ethoxybenzoic acid. (BHT); whereas paste B is made of resin ytterbium trifluoride aerosil ferric oxide. To enable the adhesion of Retroplast to the root-end surface, a specifically designed dentin-bonding agent known as GLUMA is used to condition the root-end surface. Interestingly, it was shown that Retroplast-GLUMA system demonstrates better sealing properties than the other filling materials including amalgam, intermediate restorative material (IRM), super ethoxybenzoic acid (super EBA), and glass ionomer cements (GIC). Moreover, some studies showed that it promotes a favorable periradicular healing response and can promote periodontal regeneration with some evidence of cementum layer formation over the root-end filling (Yazdi et al., 2007; Rud et al., 1991a,b). However, similar to all composites, Retroplast is
292 Advanced Dental Biomaterials technique-sensitive and any blood contamination during bonding process will reduce its strength and increase leakage. 12.6.4 Glass ionomer cements GIC is made of aqueous polymeric acids such as polyacrylic acid plus basic glass powders such as calcium aluminosilicate. It can be either chemically cured or lightcured; however, the seal of light-cured GIC is superior to the chemically cured counterparts. GIC has similar sealability and tissue response to IRM and super EBA, and like Retroplast, contamination with blood and moisture affects its properties (Scherer and Dragoo, 1995; Zetterqvist et al., 1991). 12.6.5 Diaket (3M/ESPE, Seefeld, Germany) Diaket is a polyvinyl resin with an excellent sealing ability. When mixed at thicker consistencies, it demonstrated good biocompatibility with osseous tissues with a regenerative potential (Williams and Gutmann, 1996). This material is not available in the United States. 12.6.6 Resin ionomer suspension and compomer These materials combine the advantages of both composites and glass ionomers. Two common examples are Geristore and Dyract (Dentsply, Tulsa, Oklahoma). Both these materials are light-cured, require dentin-bonding agents, and exhibit less leakage than IRM, super EBA, and amalgam (Greer et al., 2001). Maintaining a dry environment during application results in a strong bond, though these materials are less sensitive to moisture than conventional GIC (Cho et al., 1995). Geristore has a paste/paste formula made of hydrophilic bis-GMA with long-term fluoride release. Geristore also demonstrates a potential to regenerate periradicular tissue which promoted its use as a root-end filling material and a restoration material (Scherer and Dragoo, 1995). 12.6.7 Other types of cement Other types of cements are also used as root filling materials including MTA, calcium aluminate, calcium phosphate, and calcium silicate cements. Refer to those materials where mentioned in this chapter. 12.7 Perforation repair materials Many materials have been used for repairing root perforations and Box 12.8 describes the ideal properties of those materials (Hartwell and England, 1993; Kakani et al., 2015). Examples of materials used include Indium foil, amalgam,
Endodontic materials: from old materials to recent advances 293 Box 12.8 Ideal properties of perforation repair materials G G G G G G G It should provide adequate seal It should be biocompatible It should have ability to produce osteogenesis and cementogenesis It should be bacteriostatic and radiopaque It should also be beneficial to use a resorbable matrix in which a sealing material can be condensed It should be relatively inexpensive It should be nontoxic, noncariogenic, and easy to place ZnO eugenol, IRM, super EBA, GIC, adhesives, calcium hydroxide, calcium phosphate cements, MTA, calcium silicate cements, etc. Refer to those materials where mentioned in this chapter. 12.8 Summary The principles of endodontic instrumentation and therapy are well established. Thorough cleaning of the canal system to render it as free as possible of pulpal tissue, bacterial presence, and by-products, followed by complete obturation to prevent apical or coronal leakage creates an environment that promotes periradicular healing. The philosophy of achieving a hermetic seal of the root apex has been abandoned in favor of development of a fluid-tight antibacterial seal as a requirement for modern root canal sealers. The continuum of research and development of materials and instruments will drive the continued development of new materials for endodontic therapy. References Abbott, P.V., 1992. Systemic release of corticosteroids following intra-dental use. Int. Endod. J. 25, 189 191. Al-Haddad, A., Che Ab Aziz, Z.A., 2016. Bioceramic-based root canal sealers: a review. Int. J. Biomater. 2016, 10. Al-Hezaimi, K., Al-Tayar, B.A., Bajuaifer, Y.S., Salameh, Z., Al-Fouzan, K., Tay, F.R., 2011. A hybrid approach to direct pulp capping by using emdogain with a capping material. J. Endod. 37, 667 672. Allan, N.A., Walton, R.C., Schaeffer, M.A., 2001. Setting times for endodontic sealers under clinical usage and in vitro conditions. J. Endod. 27, 421 423. Ambard, A.J., Mueninghoff, L., 2006. Calcium phosphate cement: review of mechanical and biological properties. J. Prosthodont. 15, 321 328. American Association of Endodontists, 2012. Glossary of Endodontic Terms. American Association of Endodontists, Chicago, IL.
294 Advanced Dental Biomaterials Andelin, W.E., Shabahang, S., Wright, K., Torabinejad, M., 2003. Identification of hard tissue after experimental pulp capping using dentin sialoprotein (DSP) as a marker. J. Endod. 29, 646 650. Asgary, S., Eghbal, M.J., Fazlyab, M., Baghban, A.A., Ghoddusi, J., 2015. Five-year results of vital pulp therapy in permanent molars with irreversible pulpitis: a non-inferiority multicenter randomized clinical trial. Clin. Oral Investig. 19, 335 341. Asgary, S., Eghbal, M.J., Parirokh, M., Torabzadeh, H., 2006. Sealing ability of three commercial mineral trioxide aggregates and an experimental root-end filling material. Iran Endod J 1, 101 105. Athanassiadis, B., Abbott, P.V., Walsh, L.J., 2007. The use of calcium hydroxide, antibiotics and biocides as antimicrobial medicaments in endodontics. Aust. Dent. J. 52, S64 S82. Augsburger, R.A., Peters, D.D., 1990. Radiographic evaluation of extruded obturation materials. J. Endod. 16, 492 497. Basrani, B., Lemonie, C., 2005. Chlorhexidine gluconate. Aust. Endod. J. 31, 48 52. Basrani, B., Santos, J.M., Tjaderhane, L., Grad, H., Gorduysus, O., Huang, J., et al., 2002. Substantive antimicrobial activity in chlorhexidine-treated human root dentin. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 94, 240 245. Basrani, B., Tjaderhane, L., Santos, J.M., Pascon, E., Grad, H., Lawrence, H.P., et al., 2003. Efficacy of chlorhexidine- and calcium hydroxide-containing medicaments against Enterococcus faecalis in vitro. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 96, 618 624. Basrani, B.R., Manek, S., Sodhi, R.N., Fillery, E., Manzur, A., 2007. Interaction between sodium hypochlorite and chlorhexidine gluconate. J. Endod. 33, 966 969. Baumgartner, G., Zehnder, M., Paque, F., 2007. Enterococcus faecalis type strain leakage through root canals filled with gutta-percha/AH plus or Resilon/Epiphany. J. Endod. 33, 45 47. Bettina, B., Markus, H., 2012. Update on endodontic irrigating solutions. Endod. Top. 27, 74 102. Block, S.S., 2001. Disinfection, Sterilization, and Preservation. Lippincott Williams & Wilkins. Borges, A.H., Pedro, F.L., Semanoff-Segundo, A., Miranda, C.E., Pecora, J.D., Cruz Filho, A.M., 2011. Radiopacity evaluation of Portland and MTA-based cements by digital radiographic system. J. Appl. Oral. Sci. 19, 228 232. Bouillaguet, S., Wataha, J.C., Tay, F.R., Brackett, M.G., Lockwood, P.E., 2006. Initial in vitro biological response to contemporary endodontic sealers. J. Endod. 32, 989 992. Brady, J.M., del Rio, C.E., 1975. Corrosion of endodontic silver cones in humans: a scanning electron microscope and X-ray microprobe study. J. Endod. 1, 205 210. Bystrom, A., Sundqvist, G., 1985. The antibacterial action of sodium hypochlorite and EDTA in 60 cases of endodontic therapy. Int. Endod. J. 18, 35 40. Camargo, S.E., Camargo, C.H., Hiller, K.A., Rode, S.M., Schweikl, H., Schmalz, G., 2009. Cytotoxicity and genotoxicity of pulp capping materials in two cell lines. Int. Endod. J. 42, 227 237. Camilleri, J., Montesin, F.E., Brady, K., Sweeney, R., Curtis, R.V., Ford, T.R., 2005. The constitution of mineral trioxide aggregate. Dent. Mater. 21, 297 303. Camilleri, J., Sorrentino, F., Damidot, D., 2015. Characterization of un-hydrated and hydrated BioAggregate and MTA Angelus. Clin. Oral Investig. 19, 689 698. Carrilho, M.R.O., Carvalho, R.M., de Goes, M.F., Di Hipólito, V., Geraldeli, S., Tay, F.R., et al., 2007. Chlorhexidine preserves dentin bond in vitro. J. Dent. Res. 86, 90 94.
Endodontic materials: from old materials to recent advances 295 Chen, H., Teixeira, F.B., Ritter, A.L., Levin, L., Trope, M., 2008. The effect of intracanal anti-inflammatory medicaments on external root resorption of replanted dog teeth after extended extra-oral dry time. Dent. Traumatol. 24, 74 78. Cho, E., Kopel, H., White, S.N., 1995. Moisture susceptibility of resin-modified glass-ionomer materials. Quintessence Int. 26, 351 358. Chong, B.S., Pitt Ford, T.R., 2005. Root-end filling materials: rationale and tissue response. Endod. Top. 11, 114 130. Crooks, W.G., Anderson, R.W., Powell, B.J., Kimbrough, W.F., 1994. Longitudinal evaluation of the seal of IRM root end fillings. J. Endod. 20, 250 252. Dai, L., Khechen, K., Khan, S., Gillen, B., Loushine, B.A., Wimmer, C.E., et al., 2011. The effect of QMix, an experimental antibacterial root canal irrigant, on removal of canal wall smear layer and debris. J. Endod. 37, 80 84. Davies, A., 1973. The mode of action of chlorhexidine. J. Periodontal Res. Suppl. 12, 68 75. De Moor, R.J., De Bruyne, M.A., 2004. The long-term sealing ability of AH 26 and AH plus used with three gutta-percha obturation techniques. Quintessence Int. 35, 326 331. Desai, S., Chandler, N., 2009. Calcium hydroxide-based root canal sealers: a review. J. Endod. 35, 475 480. Dick, H.M., Carmichael, D.J., 1980. Reconstituted antigen-poor collagen preparations as potential pulp-capping agents. J. Endod. 6, 641 644. Dorn, S.O., Gartner, A.H., 1990. Retrograde filling materials: a retrospective success-failure study of amalgam, EBA, and IRM. J. Endod. 16, 391 393. Duarte, M.A., Demarchi, A.C., Yamashita, J.C., Kuga, M.C., Fraga Sde, C., 2003. pH and calcium ion release of 2 root-end filling materials. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 95, 345 347. Emre, B., Mehtap, M., Mahmut, S., Tolga, G., 2008. The response of subcutaneous connective tissue to a new endodontic filling material. J. Biomed. Mater. Res. Part B Appl. Biomater. 84B, 463 467. Estrela, C., Estrela, C.R., Barbin, E.L., Spano, J.C., Marchesan, M.A., Pecora, J.D., 2002. Mechanism of action of sodium hypochlorite. Braz. Dent. J. 13, 113 117. Estrela, C., Estrada-Bernabe, P.F., Daniel de Almeida-Decurcio, D., Almeida-Silva, J., Rodrigues-Araujo-Estrela, C., Poli-Figueiredo, J.A., 2011. Microbial leakage of MTA, Portland cement, Sealapex and zinc oxide-eugenol as root-end filling materials. Med. Oral Patol. Oral Cir. Bucal. 16, e418 e424. Ford, T.R., Torabinejad, M., Abedi, H.R., Bakland, L.K., Kariyawasam, S.P., 1996. Using mineral trioxide aggregate as a pulp-capping material. J. Am. Dent. Assoc. 127, 1491 1494. Frank, A.L., 1966. Therapy for the divergent pulpless tooth by continued apical formation. J. Am. Dent. Assoc. 72, 87 93. Fridland, M., Rosado, R., 2003. Mineral trioxide aggregate (MTA) solubility and porosity with different water-to-powder ratios. J. Endod. 29, 814 817. Friedman, C.M., Sandrik, J.L., Heuer, M.A., Rapp, G.W., 1975. Composition and mechanical properties of gutta-percha endodontic points. J. Dent. Res. 54, 921 925. Friedman, S., Lost, C., Zarrabian, M., Trope, M., 1995. Evaluation of success and failure after endodontic therapy using a glass ionomer cement sealer. J. Endod. 21, 384 390. Giardino, L., Pecora, G., Ambu, E., Savoldi, E., 2005. A new irrigant in the treatment of apical periodontitis: from research to clinic. In: 12th Biennial Congress European Society of Endodontology, Dublin. pp. 15 17. Glass, R.L., Zander, H.A., 1949. Pulp healing. J. Dent. Res. 28, 97 107.
296 Advanced Dental Biomaterials Goldman, M., Kronman, J.H., Goldman, L.B., Clausen, H., Grady, J., 1976. New method of irrigation during endodontic treatment. J. Endod. 2, 257 260. Goodman, A., Schilder, H., Aldrich, W., 1974. The thermomechanical properties of guttapercha: II. The history and molecular chemistry of gutta-percha. Oral Surg. Oral Med. Oral Pathol. 37, 954 961. Grawehr, M., Sener, B., Waltimo, T., Zehnder, M., 2003. Interactions of ethylenediamine tetraacetic acid with sodium hypochlorite in aqueous solutions. Int. Endod. J. 36, 411 417. Grech, L., Mallia, B., Camilleri, J., 2013. Investigation of the physical properties of tricalcium silicate cement-based root-end filling materials. Dent. Mater. 29, e20 e28. Greer, B.D., West, L.A., Liewehr, F.R., Pashley, D.H., 2001. Sealing ability of Dyract, Geristore, IRM, and super-EBA as root-end filling materials. J. Endod. 27, 441 443. Grigoratos, D., Knowles, J., Ng, Y.L., Gulabivala, K., 2001. Effect of exposing dentine to sodium hypochlorite and calcium hydroxide on its flexural strength and elastic modulus. Int. Endod. J. 34, 113 119. Grossman, L.I., 1982. The effect of pH of rosin on setting time of root canal cements. J. Endod. 8, 326 327. Hand, R.E., Smith, M.L., Harrison, J.W., 1978. Analysis of the effect of dilution on the necrotic tissue dissolution property of sodium hypochlorite. J. Endod. 4, 60 64. Hartwell, G.R., England, M.C., 1993. Healing of furcation perforations in primate teeth after repair with decalcified freeze-dried bone: a longitudinal study. J. Endod. 19, 357 361. Hashem, A.A., Wanees Amin, S.A., 2012. The effect of acidity on dislodgment resistance of mineral trioxide aggregate and bioaggregate in furcation perforations: an in vitro comparative study. J. Endod. 38, 245 249. Hayashi, Y., Imai, M., Yanagiguchi, K., Viloria, I.L., Ikeda, T., 1999. Hydroxyapatite applied as direct pulp capping medicine substitutes for osteodentin. J. Endod. 25, 225 229. Heling, I., Chandler, N.P., 1996. The antimicrobial effect within dentinal tubules of four root canal sealers. J. Endod. 22, 257 259. Hulsmann, M., Heckendorff, M., Lennon, A., 2003. Chelating agents in root canal treatment: mode of action and indications for their use. Int. Endod. J. 36, 810 830. Jitaru, S., Hodisan, I., Timis, L., Lucian, A., Bud, M., 2016. The use of bioceramics in endodontics—literature review. Clujul Med. 89, 470 473. Kaida, H., Hamachi, T., Anan, H., Maeda, K., 2008. Wound healing process of injured pulp tissues with emdogain gel. J. Endod. 34, 26 30. Kakani, A.K., Veeramachaneni, C., Majeti, C., Tummala, M., Khiyani, L., 2015. A review on perforation repair materials. J. Clin. Diagn. Res. 9, ZE09 ZE13. Kazemi, R.B., Safavi, K.E., Spangberg, L.S., 1993. Dimensional changes of endodontic sealers. Oral Surg. Oral Med. Oral Pathol. 76, 766 771. Keskin, C., Demiryurek, E.O., Ozyurek, T., 2015. Color stabilities of calcium silicate-based materials in contact with different irrigation solutions. J. Endod. 41, 409 411. Kiatwateeratana, T., Kintarak, S., Piwat, S., Chankanka, O., Kamaolmatyakul, S., Thearmontree, A., 2009. Partial pulpotomy on caries-free teeth using enamel matrix derivative or calcium hydroxide: a randomized controlled trial. Int. Endod. J. 42, 584 592. Kim, Y.K., Grandini, S., Ames, J.M., Gu, L.S., Kim, S.K., Pashley, D.H., et al., 2010. Critical review on methacrylate resin-based root canal sealers. J. Endod. 36, 383 399. Kleier, D.J., Averbach, R.E., 1988. Painful dysesthesia of the inferior alveolar nerve following use of a paraformaldehyde-containing root canal sealer. Dent. Traumatol. 4, 46 48. Krastl, G., Allgayer, N., Lenherr, P., Filippi, A., Taneja, P., Weiger, R., 2013. Tooth discoloration induced by endodontic materials: a literature review. Dent. Traumatol. 29, 2 7.
Endodontic materials: from old materials to recent advances 297 Kratchman, S.I., 2004. Perforation repair and one-step apexification procedures. Dent. Clin. North Am. 48, 291 307. Lambrianidis, T., Margelos, J., Beltes, P., 1999. Removal efficiency of calcium hydroxide dressing from the root canal. J. Endod. 25, 85 88. Law, A.S., 2013. Considerations for regeneration procedures. J. Endod. 39, S44 S56. Lewis, B.B., Chestner, S.B., 1981. Formaldehyde in dentistry: a review of mutagenic and carcinogenic potential. J. Am. Dent. Assoc. 103, 429 434. Lianjia, Y., Yuhao, G., White, F.H., 1993. Bovine bone morphogenetic protein-induced dentinogenesis. Clin. Orthop. Relat. Res. 305 312. Loh, A., O’Hoy, P., Tran, X., Charles, R., Hughes, A., Kubo, K., et al., 2004. Evidencebased assessment: evaluation of the formocresol versus ferric sulfate primary molar pulpotomy. Pediatr. Dent. 26, 401 409. Lovato, K.F., Sedgley, C.M., 2011. Antibacterial activity of Endosequence Root Repair Material and proroot MTA against clinical isolates of Enterococcus faecalis. J. Endod. 37, 1542 1546. Lovschall, H., Fejerskov, O., Flyvbjerg, A., 2001. Pulp-capping with recombinant human insulin-like growth factor I (rhIGF-I) in rat molars. Adv. Dent. Res. 15, 108 112. Macedo, R., Verhaagen, B., Rivas, D.F., Versluis, M., Wesselink, P., van der Sluis, L., 2014. Cavitation measurement during sonic and ultrasonic activated irrigation. J. Endod. 40, 580 583. Margelos, J., Eliades, G., Verdelis, C., Palaghias, G., 1997. Interaction of calcium hydroxide with zinc oxide-eugenol type sealers: a potential clinical problem. J. Endod. 23, 43 48. McWalter, G.M., El-Kafrawy, A.H., Mitchell, D.F., 1976. Long-term study of pulp capping in monkeys with three agents. J. Am. Dent. Assoc. 93, 105 110. Min, K.S., Park, H.J., Lee, S.K., Park, S.H., Hong, C.U., Kim, H.W., et al., 2008. Effect of mineral trioxide aggregate on dentin bridge formation and expression of dentin sialoprotein and heme oxygenase-1 in human dental pulp. J. Endod. 34, 666 670. Mohammadi, Z., 2008. Sodium hypochlorite in endodontics: an update review. Int. Dent. J. 58, 329 341. Mohammadi, Z., Dummer, P.M., 2011. Properties and applications of calcium hydroxide in endodontics and dental traumatology. Int. Endod. J. 44, 697 730. Orstavik, D., Nordahl, I., Tibballs, J.E., 2001. Dimensional change following setting of root canal sealer materials. Dent. Mater. 17, 512 519. Parirokh, M., Torabinejad, M., 2010a. Mineral trioxide aggregate: a comprehensive literature review—Part I: Chemical, physical, and antibacterial properties. J. Endod. 36, 16 27. Parirokh, M., Torabinejad, M., 2010b. Mineral trioxide aggregate: a comprehensive literature review—Part III: Clinical applications, drawbacks, and mechanism of action. J. Endod. 36, 400 413. Parolia, A., Kundabala, M., Rao, N.N., Acharya, S.R., Agrawal, P., Mohan, M., et al., 2010. A comparative histological analysis of human pulp following direct pulp capping with propolis, mineral trioxide aggregate and Dycal. Aust. Dent. J. 55, 59 64. Qureshi, A., Soujanya, E., Nandakumar, Pratapkumar, Sambashivarao, 2014. Recent advances in pulp capping materials: an overview. J. Clin. Diagn. Res. 8, 316 321. Raina, R., Loushine, R.J., Weller, R.N., Tay, F.R., Pashley, D.H., 2007. Evaluation of the quality of the apical seal in Resilon/Epiphany and gutta-percha/AH Plus-filled root canals by using a fluid filtration approach. J. Endod. 33, 944 947. Rud, J., Munksgaard, E.C., Andreasen, J.O., Rud, V., 1991a. Retrograde root filling with composite and a dentin-bonding agent. 2. Endod. Dent. Traumatol. 7, 126 131.
298 Advanced Dental Biomaterials Rud, J., Munksgaard, E.C., Andreasen, J.O., Rud, V., Asmussen, E., 1991b. Retrograde root filling with composite and a dentin-bonding agent. 1. Endod. Dent. Traumatol. 7, 118 125. Safavi, K.E., Nichols, F.C., 1993. Effect of calcium hydroxide on bacterial lipopolysaccharide. J. Endod. 19, 76 78. Safavi, K.E., Spangberg, L.S., Langeland, K., 1990. Root canal dentinal tubule disinfection. J. Endod. 16, 207 210. Scherer, W., Dragoo, M.R., 1995. New subgingival restorative procedures with Geristore resin ionomer. Pract. Periodontics Aesthet. Dent. 7, 1 4. Schroeder, A., 1975. Ledermix 1962—Ledermix today. Evaluation after 13 years of experience. Zahnarztl Prax 26, 195 196. She, C.M., Cheung, G.S., Zhang, C.F., 2016. Long-term follow-up of a revascularized immature necrotic tooth evaluated by CBCT. Case Rep. Dent. 2016, 4982458. Shipper, G., Orstavik, D., Teixeira, F.B., Trope, M., 2004. An evaluation of microbial leakage in roots filled with a thermoplastic synthetic polymer-based root canal filling material (Resilon). J. Endod. 30, 342 347. Shrestha, A., Kishen, A., 2016. Antibacterial nanoparticles in endodontics: a review. J. Endod. 42, 1417 1426. Silva, G.A., Gava, E., Lanza, L.D., Estrela, C., Alves, J.B., 2013. Subclinical failures of direct pulp capping of human teeth by using a dentin bonding system. J. Endod. 39, 182 189. Siqueira JR., J.F., Batista, M.M., Fraga, R.C., de Uzeda, M., 1998. Antibacterial effects of endodontic irrigants on black-pigmented Gram-negative anaerobes and facultative bacteria. J. Endod. 24, 414 416. Siqueira Jr., J.F., Rocas, I.N., Favieri, A., Lima, K.C., 2000. Chemomechanical reduction of the bacterial population in the root canal after instrumentation and irrigation with 1%, 2.5%, and 5.25% sodium hypochlorite. J. Endod. 26, 331 334. Siren, E.K., Haapasalo, M.P., Waltimo, T.M., Orstavik, D., 2004. In vitro antibacterial effect of calcium hydroxide combined with chlorhexidine or iodine potassium iodide on Enterococcus faecalis. Eur. J. Oral Sci. 112, 326 331. Smith, J.J., Wayman, B.E., 1986. An evaluation of the antimicrobial effectiveness of citric acid as a root canal irrigant. J. Endod. 12, 54 58. Spangberg, L., Engstrom, B., Langeland, K., 1973. Biologic effects of dental materials. 3. Toxicity and antimicrobial effect of endodontic antiseptics in vitro. Oral Surg. Oral Med. Oral Pathol. 36, 856 871. Spangberg, L.S., Barbosa, S.V., Lavigne, G.D., 1993. AH 26 releases formaldehyde. J. Endod. 19, 596 598. Stojicic, S., Zivkovic, S., Qian, W., Zhang, H., Haapasalo, M., 2010. Tissue dissolution by sodium hypochlorite: effect of concentration, temperature, agitation, and surfactant. J. Endod. 36, 1558 1562. Stojicic, S., Shen, Y., Qian, W., Johnson, B., Haapasalo, M., 2012. Antibacterial and smear layer removal ability of a novel irrigant, QMiX. Int. Endod. J. 45, 363 371. Tarim, B., Hafez, A.A., Cox, C.F., 1998. Pulpal response to a resin-modified glass-ionomer material on nonexposed and exposed monkey pulps. Quintessence Int. 29, 535 542. Tay, F.R., Pashley, D.H., 2007. Monoblocks in root canals—a hypothetical or a tangible goal. J. Endod. 33, 391 398. Torabinejad, M., Hong, C.U., Mcdonald, F., Pitt Ford, T.R., 1995. Physical and chemical properties of a new root-end filling material. J. Endod. 21, 349 353.
Endodontic materials: from old materials to recent advances 299 Torabinejad, M., Khademi, A.A., Babagoli, J., Cho, Y., Johnson, W.B., Bozhilov, K., et al., 2003. A new solution for the removal of the smear layer. J. Endod. 29, 170 175. Walivaara, D.A., Abrahamsson, P., Samfors, K.A., Isaksson, S., 2009. Periapical surgery using ultrasonic preparation and thermoplasticized gutta-percha with AH Plus sealer or IRM as retrograde root-end fillings in 160 consecutive teeth: a prospective randomized clinical study. Oral Surg Oral Med Oral Pathol Oral Radiol Endod 108, 784 789. Williams, S.S., Gutmann, J.L., 1996. Periradicular healing in response to Diaket root-end filling material with and without tricalcium phosphate. Int. Endod. J. 29, 84 92. Windley III, W., Teixeira, F., Levin, L., Sigurdsson, A., Trope, M., 2005. Disinfection of immature teeth with a triple antibiotic paste. J. Endod. 31, 439 443. Yasuda, Y., Ogawa, M., Arakawa, T., Kadowaki, T., Saito, T., 2008. The effect of mineral trioxide aggregate on the mineralization ability of rat dental pulp cells: an in vitro study. J. Endod. 34, 1057 1060. Yavari, H.R., Shahi, S., Rahimi, S., Shakouie, S., Roshangar, L., Mesgari Abassi, M., et al., 2009. Connective tissue reaction to white and gray MTA mixed with distilled water or chlorhexidine in rats. Iran. Endod. J. 4, 25 30. Yazdi, P.M., Schou, S., Jensen, S.S., Stoltze, K., Kenrad, B., Sewerin, I., 2007. Dentinebonded resin composite (Retroplast) for root-end filling: a prospective clinical and radiographic study with a mean follow-up period of 8 years. Int. Endod. J. 40, 493 503. Zakariasen, K.L., Brayton, S.M., Collinson, D.M., 1990. Efficient and effective root canal retreatment without chloroform. J. Can. Dent. Assoc. 56, 509 512. Zehnder, M., 2006. Root canal irrigants. J. Endod. 32, 389 398. Zehnder, M., Soderling, E., Salonen, J., Waltimo, T., 2004. Preliminary evaluation of bioactive glass S53P4 as an endodontic medication in vitro. J. Endod. 30, 220 224. Zetterqvist, L., Hall, G., Holmlund, A., 1991. Apicectomy: a comparative clinical study of amalgam and glass ionomer cement as apical sealants. Oral Surg. Oral Med. Oral Pathol. 71, 489 491. Zielinski, T.M., Baumgartner, J.C., Marshall, J.G., 2008. An evaluation of Guttaflow and gutta-percha in the filling of lateral grooves and depressions. J. Endod. 34, 295 298. Further reading Guideline Pulp Therapy, 2016. Guideline on pulp therapy for primary and immature permanent teeth. Pediatr. Dent. 38, 280 288. Chong, B.S., Pitt Ford, T.R., Watson, T.F., Wilson, R.F., 1995. Sealing ability of potential retrograde root filling materials. Dent. Traumatol. 11, 264 269.
13 Fiber-reinforced composites Daniel Varley1, Safiyya Yousaf1, Mansour Youseffi1, Masoud Mozafari2,3, Zohaib Khurshid4 and Farshid Sefat5,6 1 Medical Engineering Department, Faculty of Engineering and Informatics, University of Bradford, Bradford, United Kingdom, 2Bioengineering Research Group, Nanotechnology and Advanced Materials Department, Materials and Energy Research Centre (MERC), Tehran, Iran, 3Department of Tissue Engineering & Regenerative Medicine, Faculty of Advanced Technologies in Medicine, Iran University of Medical Sciences (IUMS), Tehran, Iran, 4Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia, 5Interdisciplinary Research Centre in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom, 6 Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom Chapter Outline 13.1 Introduction 302 13.2 Anatomy and physiology of teeth 13.2.1 13.2.2 13.2.3 13.2.4 13.2.5 302 Enamel 303 Dentin 303 Dental pulp 304 Cementum 304 Tooth development 304 13.3 Mechanical properties of teeth 305 13.4 Biomaterials used in dentistry 305 13.4.1 Metals 305 13.4.2 Ceramics 306 13.4.3 Composites 306 13.5 Fiber-reinforced composites 306 13.5.1 Fiber-reinforced composite composition 306 13.5.2 Influencing factors on mechanical properties 308 13.6 Clinical applications of fiber-reinforced composites 13.6.1 13.6.2 13.6.3 13.6.4 13.6.5 13.6.6 Tooth restoration 311 Implants 311 Endodontics 312 Prosthodontics 312 Orthodontic 312 Periodontal 312 13.7 Conclusion 313 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00013-X Copyright © 2019 Elsevier Ltd. All rights reserved. 311
302 Advanced Dental Biomaterials References 313 Further reading 315 13.1 Introduction Dental application is constantly evolving as a result of innovative treatment solutions based on the development of new biomaterials, advancement in technologies, and more successful treatment techniques, with fiber-reinforced composites (FRCs) being a superb example of this. The use of fibers within composite materials has greatly enhanced their uses in the field of dentistry, giving dentists more options due to the many applications of this exciting group of materials. A FRC is a combination of conventional dental composite and most commonly glass fibers; this combination gives the same level of strength and flexibility that is also found in boats, light aircraft, and F1 racing cars. FRCs consist of three components: (1) the matrix, this is usually made of resin and becomes a polymer after curing or polymerization. Additionally, it provides a place for the fibers to reside. (2) The reinforcing constituent, which is fibers of high strength and modulus; these are usually glass, carbon, or polyethylene fibers. (3) The fine interphase region, also known as the interface; this is the key element of the composite as it transfers the load from the matrix to the fibers (Zhandarov and Mäder, 2005). For many decades, engineers have been using fibers as high aspect ratio fillers to construct materials/ devices with high levels of strength and fracture toughness. Hence, this is why these materials are desirable in the field of dentistry (He et al., 2017). FRCs have been the center of attention in dentistry for the last two decades because of their excellent adhesion and esthetic appearance. They have provided many clinical applications, such as fixed prosthodontics, restorative dentistry, periodontology orthodontics, and in repairs of prosthetic devices (Nagata et al., 2016). 13.2 Anatomy and physiology of teeth Human tooth can be broken down into two anatomical parts. The root is the part embedded within the jaw, as seen in Fig. 13.1. The purpose of the root is to anchor the tooth within the bony socket of the jaw. The root of the tooth is not normally visible as it is found below the cementoenamel junction and extends deep into the gingivae. Additionally, the majority of the root is composed of dentin and contains pulp canals. Figure 13.1 A diagram to show the structure of a fiber-reinforced composite.
Fiber-reinforced composites 303 Figure 13.2 An image showing the cross section of the tooth, including the various layers. Also depicted in Fig. 13.2 is the second anatomical part, which is known as the crown; this part of the tooth is above the cementoenamel junction and is visible within the mouth. The crown is completely covered with enamel, which is extremely hard and protects the tooth from the many stresses and strains faced on a daily basis (Berkovitz et al., 2009). A human tooth is very complex as it consists of various layers, each with different purposes and properties. 13.2.1 Enamel As depicted in Fig. 13.2, enamel is the outermost layer of the human tooth. Furthermore it is the hardest and most highly mineralized substance within the human body, which allows it to cope with impact forces and shearing very well. One very important property of enamel is that it has very high abrasion resistance; this constitutes to wear being very slow. Unfortunately, if enamel is destroyed, the body cannot repair or reproduce any more, and the reason behind this is that the enamel does not contain any living cells, so it cannot regenerate (Berkovitz et al., 2009). 13.2.2 Dentin Dentin, pale yellow in color, forms the bulk of the tooth and is a mineralized tissue; it is usually covered by enamel on the crown, cementum on the root, and surrounds the entire pulp. Dentin consists of a large quantity of small parallel tubules that lay within a collagen matrix, which is highly mineralized. This structure provides
304 Advanced Dental Biomaterials dentin with greater compressive, tensile, and flexural strength than enamel. However, dentin is not as hard as enamel but is harder than bone and cementum. Odontoblasts are responsible for forming the tissue, and the cell bodies of these odontoblasts play a key role in lining the deep surface of dentin. The combination of enamel and dentin provides a very hard and rigid structure that is perfect for tearing and masticating. 13.2.3 Dental pulp Dental pulp is a specialized, unmineralized oral tissue, which is composed of soft connective tissue, nervous, and vascular elements (Berkovitz et al., 2009). Pulp has a soft gelatinous consistency and occupies the central pulp cavity of each tooth; the main role is to form dentin of the tooth. In newly erupted teeth, the dental pulp is large, but as teeth age the pulp cavity becomes more constricted, meaning the pulp becomes progressively smaller (Nelson, 2014). The mesenchymal tissue is characterized by the presence of odontoblasts, which are the primary cells of dental pulp. Odontoblasts are extremely important as they continuously lay down secondary dentin throughout life. Furthermore, these are responsible for the production of tertiary reactionary dentin when the tooth comes under minor insult or experiences dental caries (Nor, 2006). 13.2.4 Cementum Cementum is the specialized thin layer of calcified tissue covering the dentin of the root; its role is to help support the tooth within the jaw by attaching to collagen fibers of the periodontal ligament. Unlike bone, cementum is not vascular and exhibits little turnover; it grows slowly by surface apposition, throughout life (Berkovitz et al., 2009). Cementum, such as dentin, is also pale yellow in color with a dull surface. However, when compared, cementum is a lot softer and far more permeable than dentin. Due to the softness and thinness of the tissue, it can be quite easily removed by abrasion, hence for this reason, it is well protected. 13.2.5 Tooth development There are three phases that sum up tooth development: initiation, morphogenesis, and histogenesis. During initiation, there is an appearance of tooth germs along the oral epithelium, and the sites of future teeth are established. It is during the morphogenesis phase a that combination of cell movement and proliferation determines the shape of the tooth. Finally, during the histogenesis phase, the differentiation of cells proceed into fully formed dental tissues, both mineralized and unmineralized. Additionally, tooth germs are categorized into three stages before actual tooth eruption, these are the bud stage, cap stage, and bell stage and are shown in Fig. 13.3.
Fiber-reinforced composites 305 Figure 13.3 An image showing the various stages of tooth development. (A) Thickening, (B) bud, (C) cap, (D) bell and (E) erupted tooth. 13.3 Mechanical properties of teeth The mechanical properties of teeth include hardness, elasticity, viscoelasticity, and fracture behavior. Hardness is a measure of the hardness or softness of a material and represents the ability of a solid material to resist plastic deformation, elastic deformation, and destruction. Elasticity is a term used to characterize how a material changes under an applied external force and resumes after the force is removed. The elastic property parameters of human teeth include the shear modulus, elastic modulus, and Poisson’s ratio. Viscoelasticity is the property of materials that exhibit both viscous and elastic characteristics when undergoing deformation. 13.4 Biomaterials used in dentistry The field of dentistry is blessed with an ever-expanding number of biomaterials that can be used for many different treatments and procedures. These biomaterials have a diverse range of physical and mechanical properties and can be categorized into three different sections: 13.4.1 Metals G G G G Amalgam—Also commonly named filling, amalgam is a sealant made of small particles of tin, copper, and silver allied with mercury. The mechanical properties and longevity of this material made it the first choice for many years. On the other hand, this material does have its flaws, it is not esthetically pleasing, and it contains mercury, which is poisonous to humans. Gold—In dentistry gold is found in the form of alloys. This material is ideal for posterior prosthetic realizations due to its harmlessness, rigidity, and precision. Titanium—This material is used most commonly in implantology due to its antiallergic qualities. Steel (chrome-cobalt)—Due to its high tensile strength, this material is used most commonly in removable prosthetics for frameworks and clasps.
306 Advanced Dental Biomaterials 13.4.2 Ceramics Dental ceramics are materials that are part of systems designed with the purpose of producing dental prostheses, which are used to replace damaged or missing dental structures. Ceramics have a high compressive strength but low tensile strength, meaning they are brittle and are likely to fracture under low strains (Shenoy and Shenoy, 2010). The different types of ceramics used in dentistry: G G G G Glass-ceramics—A glass ceramic is a multiphase solid containing a residual glass phase with a finely dispersed crystalline phase. Their size and growth rate are regulated by the time and temperature of heat treatment. Metal ceramics—A hybrid construction type where a porcelain covering encases a metal substructure. These crowns can provide both good esthetics and strength. However, many problems can occur due to the release of metallic ions into the gingival tissue. Infiltrated ceramics—Infiltrated ceramics are formed through a process called slipcasting. This process gives the material greater strength and toughness and allows fewer defects from processing. The three different infiltrated ceramics are spinel, alumina, and zirconia. Polycrystalline ceramics—Polycrystalline ceramics are formed from powders that can be packed only to 70% of their theoretical density. An example of a polycrystalline ceramic would be transformation-toughened zirconium oxide; this oxide can take three different crystalline structures depending on temperature (Shenoy and Shenoy, 2010). 13.4.3 Composites G G Composite resin—These are a type of synthetic material that consist of a resin matrix and quartz or silica particles known as fillers. The most common matrixes used in dental composites are: bisphenol A-glycidyl methacrylate, urethane dimethacrylate, and triethylene glycol dimethacrylate. The size of the filler determines the surface roughness of the composite, while the amount of filler increases the mechanical properties of the composite. Glass ionomer cements—Glass ionomer cements are tooth-colored materials that bond chemically to dental hard tissues and release fluoride for a relatively long period. These tend to be the material of choice when restoring primary teeth. 13.5 Fiber-reinforced composites 13.5.1 Fiber-reinforced composite composition 13.5.1.1 Fiber types The properties of a composite material are determined by the properties of the fiber used to reinforce it. There are three types of fibers mainly used to reinforce composites in dentistry, these are glass, carbon, and polyethylene, and each of these reinforcing fibers give very diverse physical and mechanical properties to the finished materials.
Fiber-reinforced composites 307 13.5.1.2 Glass Glass fibers are the most common of all reinforcing fibers used commercially for polymeric matrix composites (Ravindra et al., 1997). When toughness and tensile strength must be maximized, glass fibers are the reinforcement of choice. The recent innovations in glass fiber formulation allow strengths similar to that of carbon fibers (Stickel and Nagarajan, 2012). The main advantages of glass fibers are low cost, high tensile strength, good insulating properties, and high chemical resistance. While on the other hand, the disadvantages associated with glass fibers are their low tensile modulus, low fatigue resistance, as well as their high density and higher sensitivity to wear (Migliaresi and Alexander, 2004; Ravindra et al., 1997). Depending on the chemical composition of glass they are commercially available in different grades. 1. Glass A—or Alkali glass was a common basic material for glass fiber production. This glass is cheap and can be used as filler for plastics; however, with low chemical resistance to water and low strength, this type of glass is not as desirable in the field of dentistry. 2. Glass C—or Chemical-resistant glass, these fibers have good corrosion resistance and are used in the manufacturing of surfacing layers to provide additional chemical resistance over E glass. 3. Glass E—or Electric glass is the most common type of glass fiber used in dentistry due to its excellent electrical and mechanical properties. These fibers are characterized by their ability to strongly resist attack by water. The main disadvantage of this grade is unpredictable constituents such as fluorine. 4. Glass R—A reinforcement glass made with calcium aluminosilicates used where higher strength and acid corrosion resistance is needed. 5. Glass S—This glass has high elasticity and strength. However, it is very expensive due to its manufacturing processes. Glass S has very limited applications and is used mostly in the aerospace industry. 13.5.1.3 Carbon Carbon fibers were one of the early developments in composite reinforcement. In 1969, the first paper on the fabrication and characterization of an experimental composite with a pure silica matrix reinforced with carbon fibers was published (Boccaccini, 2005). Carbon fiber is a material consisting of fibers around 5 10 µm in diameter and composed mostly of carbon atoms. Carbon atoms are bonded together in crystals that are aligned parallel to the long axis of the fiber. It is this crystal alignment that gives the fiber high strength-to-volume ratio. Carbon fiber reinforced materials are lightweight, very high in strength, and have good impact resistance. However, on the other hand carbon fibers possess poor shear strength and when the material fractures fibers disintegrate making the results unpredictable. Carbon fibers did not produce the best results for in vivo tests; in fact, many patients that had been treated with carbon fiber reinforced biomaterial presented
308 Advanced Dental Biomaterials with osteolysis (Migliaresi and Alexander, 2004), resulting in a number of products being withdrawn from the market. 13.5.1.4 Polyethlene Polyethylene is a very versatile material and has many applications in the field of dentistry. The main use is to reinforce composites in a fiber form. Polyethylene fibers are used due to them being highly esthetic when compared with cast metal or metal mesh. Additionally, polyethylene fibers have high flexibility as well as being thin and strong making them suitable to strengthen composites in large restorations. Polyethylene fibers are recognized for having good biocompatibility and can be surface treated to enhance resin bonding. However, most clinical cases using polyethylene fibers were short term. To enhance credibility, further clinical trials with long-term results are required (Agrawal, 2014). 13.5.2 Influencing factors on mechanical properties 13.5.2.1 Fiber quantity The number of fibers used to reinforce a material can have a significant effect on the mechanical properties of that material. In general GFRCs have a high volume fraction of fiber at around 60 vol.%, meanwhile in dentistry the fiber fraction is kept relatively low, as the glass fiber should be covered with a layer of unfilled polymer (Abdulmajeed et al., 2011). Wear behavior of GFRC was investigated by Callaghan et al., what they found was the specimen with 7.6 wt.% glass fibers had very little matrix and a cluster of fibers due to possible overloading. High concentration of glass fibers can lead to premature fiber fracture. To achieve the best wear resistance, bonding, and decreased risk of fracture the amount of fibers should be between 2.0 and 7.6 wt.% for the matrix (Callaghan et al., 2006). 13.5.2.2 Fiber distribution The way fibers are distributed within a composite greatly influences the properties of that material (Khan et al., 2015). Fibers that are evenly distributed tend to have a positive effect on fatigue resistance, meaning it can withstand more cyclic stresses. However, on the other hand, if the fibers are situated in one place then they can cause the strength and Young’s modulus to increase (Narva, 2002). Fonseca et al. (2014) found that fracture toughness, flexural strength, and compressive strength were all higher when using randomly distributed short glass fibers. 13.5.2.3 Fiber orientation As shown in Fig. 13.4, fibers can orientate perpendicular to each other, parallel to each other, or randomly distributed. Unidirectional fibers are anisotropic, meaning reinforcement is in one direction only, while bidirectional fibers are orthotropic and
Fiber-reinforced composites 309 Figure 13.4 A diagram to show the various orientations of fibers within a matrix; the arrows represent the direction of force. offer support in two directions. Randomly orientated short fibers provide an isotropic reinforcement in all directions. Unidirectional longitudinal FRCs are a lot stronger when the force is applied along the direction of the fiber; when a stress is applied at an angle to the direction of the fibers, their strength reduces (Garoushi et al., 2006). Fiber orientation also plays a key role in adhesion to dentin and enamel. Tezvergil et al. showed that randomly orientated fibers gave the highest shear bond strength with enamel. However, on the other hand, it is bidirectional fibers that give the highest shear bond strength to dentin (Tezvergil-Mutluay, 2017). 13.5.2.4 Fiber length It is essential that stress transfers from the polymer matrix to the fibers, a key factor for this to happen is to determine the length of the reinforcing fibers (Petersen, 2005). To ensure the stress is transferred between the fibers, the reinforcing fibers length must be equal to or greater than the critical fiber length (Lc). When the length of the fiber is much higher than Lc, it is known as continuous, and shorter fibers are termed discontinuous. Fibers that are significantly shorter than Lc cause the matrix to deform around the fibers, producing very little transference of stress and virtually no reinforcement (Garoushi et al., 2007).
310 Advanced Dental Biomaterials 13.5.2.5 Adhesion of fibers to polymer matrix Adhesion is defined as a molecular or atomic attraction between two contacting surfaces promoted by the interfacial force or attraction between the molecules or atoms. Adhesion between matrix and fiber is of great importance as it affects the strength of the material. If the adhesion was inadequate, the fibers would act as an inclusion in the matrix, therefore weakening the composite (Khan et al., 2015). Typically all glass fibers contain hydroxyl groups ( OH) on their surface, meaning that they potentially have good adhesion to resinous materials (Vallittu, 2017). To help the adhesion process, fibers are chemically treated on their surface, this is known as sizing. The process of sizing consists of fibers being coated with antistatic compounds; there is a diverse blend of ingredients used in sizing, such as polymeric components, a coupling agent, and a range of additives (Vallittu, 2017). Silane coupling agents are synthetic compounds and are mainly for surface modification between two different materials. These bonds are created artificially and so not occur naturally (Vallittu, 2017). 13.5.2.6 Impregnation of fibers with polymer matrix Good contact of reinforcing fibers to the resin matrix is an essential requirement for adhesion of fibers with the resin; therefore wetting of the surface of the fiber is required (Vallittu, 1998). Resin impregnation relates to the surface wetting properties and viscosity of resin material. Poor impregnation creates voids between the matrix and fiber resulting in decreased load-bearing capacity (Abdulmajeed et al., 2011). Another additional problem with poor impregnation is it can lead to water sorption, this is where water is allowed to enter through cracks and voids reducing bond strength between matrix and fiber, leading to hydrolytic degradation (Khan et al., 2015). 13.5.2.7 Water sorption The effect of sorption and solubility may lead to adverse consequences in a dental FRC. Over time, the absorption of water can lead to the decrease in mechanical properties. The reason for this is a hydrolytic breakdown of the bond between fibers and matrix or hydrolytic degradation of the fibers (Soderholm et al., 1984). It has been found that the mechanical properties, such as wear resistance, flexural strength, tensile strength, and modulus of elasticity, are all deleteriously affected by water absorption (Soderholm and Roberts, 1990; Sideridou et al., 2003). The negative effect that water sorption has on mechanical properties of FRCs has raised concern about the shortened service life of these materials (Ferracane, 2006). Although water sorption has many negative effects on the properties of FRC materials, it also can have some positive effects. McCabe and Rusby (2004) found expansion caused by water sorption can help to overcome stresses from polymerization shrinkage on the restoration interphase.
Fiber-reinforced composites 311 13.5.2.8 Polymerization shrinkage When polymerized, dental FRCs shrink due to a decrease in the distance between the atoms as the monomers react to establish a covalent bond (Rueggeberg, 2002; Schneider et al., 2010). This polymerization shrinkage creates contraction and internal stresses resulting in deformation in the surrounding tooth structure and, furthermore, a still poor marginal seal (Kleverlaan and Feilzer, 2005). The curing reaction of dental resin composites often involves visible-light initiation of dimethacrylate monomers to form a highly cross-linked polymer. Higher fiber volume fraction leads to a decrease in polymerization shrinkage (Puckett et al., 2007). The problem of polymerization shrinkage and shrinkage stress is still yet to be solved; however, alternative curing methods may be the answer. 13.6 Clinical applications of fiber-reinforced composites FRCs are a group of relatively new materials, and due to this factor, their clinical experience is limited. However, with fibers and matrix systems that are biocompatible, they have become biomaterials within dentistry. Additionally, these materials possess great esthetics, corrosion resistance, no metal allergy response, and the ability to be handled chair-side (Khan et al., 2015). The various clinical (dental) applications of FRCs are listed and explained here. 13.6.1 Tooth restoration Someday it will likely be possible to culture cells from a patient and use them to grow a new tooth to replace one they have lost. However, until then dentists are required to use a number of artificial replacements for lost and missing teeth. FRCs provide an amazing blend of strength, esthetics, and the least amount of drilling to sound tooth tissue, making these materials an excellent candidate for tooth restoration. FRCs can be adapted by hand to connect teeth together, reinforce cracked teeth, support crowns or bridges, and can be used as a direct filling (Butterworth et al., 2003). 13.6.2 Implants Implants are routinely restored with overdentures, fixed partial dentures, or hybrid prostheses (Tayab and Shetty, 2015). However, standard metallic abutment has proven unacceptable for hybrid prostheses due to limitations in bonding; hence, FRC has become the common material of choice, as they can be bonded and retained successfully due to excellent bonding properties (Bergendal et al., 1995).
312 Advanced Dental Biomaterials 13.6.3 Endodontics FRC endodontic posts have been introduced to replace ceramics and metal alloys, due to them having high tensile strength and similar rigidity and elasticity to that of root dentin. This minimizes the risk of root fractures caused by tension peaks induced by loading forces (Khan et al., 2015). Catastrophic collapse, which leads to a fracture in the tooth root, is less likely in contrast to posts made from metal (Mannocci et al., 1999). 13.6.4 Prosthodontics Metal ceramic fixed partial dentures are a standard restoration for single-tooth replacement in dental practice. For many years metal ceramic prostheses have provided an excellent record due to their strong and durable substructure; however, a number of problems continue to arise regarding esthetics (Freilich et al., 1998). FRCs have combatted the problems with esthetics, as the substructure is translucent and requires no opaque masking. Various types of fibers may be used to strengthen provisional partial dentures; however, a number of studies have shown glass fibers are the most effective (Khan et al., 2015). 13.6.5 Orthodontic As FRC’s success grows in other areas of dentistry, they are also gaining popularity in the orthodontic area due to their use a space-maintaining device. Both glass and polyethylene fibers have been used to produce unilateral fixed space maintainers (Tayab and Shetty, 2015). In this application, bonding and fracture characteristics are of great importance. Subramaniam et al. found that FRC space maintainers demonstrated superior retention when compared with stainless steel band and loop maintainers after a 1-year follow-up period of children aged 6 8 years (Tayab and Shetty, 2015; Subramaniam et al., 2008). 13.6.6 Periodontal A splint is defined as a rigid or flexible device that maintains a displaced or movable part in a position. Splints are used in the cases of periodontal disease; where periodontal fibers are disorganized, resorption of bone and destruction of epithelial cell attachment occur (Davies et al., 2001), resulting in teeth drift, and gingival margins start to recede. Due to excellent bonding properties, FRC materials are extremely common for splinting periodontally mobile teeth. FRCs provide stabilization with no need for extra tooth preparation as retention is accomplished by resin adhesion to etched enamel (Freilich et al., 2002).
Fiber-reinforced composites 13.7 313 Conclusion This chapter demonstrates the wide range of clinical applications that FRCs have in the field of dentistry. As more research is carried out on FRCs, there will be a greater understanding of these fascinating materials. I can be sure that the clinical applications in dentistry will expand further. In a generation where esthetics is paramount, FRCs are becoming extremely popular and sought-after by patients. However, there are still many questions to be answered about FRCs as well as barriers to overcome. The literature has shown that FRCs have good mechanical properties; the problem is that many factors can influence these properties, making it difficult for dental technicians to carry out high-quality treatments with long-term positive outcomes. It is agreed by many that more long-term clinical trials are a necessity to answer the questions about the longevity of dental FRCs. These long-term trials would increase the credibility of these new materials massively, as the literature has very few longitudinal cases that monitor the performance of FRC restorations. References Abdulmajeed, A.A., Narhi, T.O., Vallitu, P.K., Lassila, L.V., 2011. The effect of high fiber fraction on some mechanical properties of unidirectional glass fibre-reinforced composite. Dent. Mater. 27, 313 321. Agrawal, M., 2014. Applications of ultrahigh molecular weight polyethylene fibres in dentistry: a review article. J. Adv. Med. Dent. Sci. 2 (2), 95 99. Bergendal, T., Ekstrand, K., Karlsson, U., 1995. Evaluation of implant supported carbon/ graphite fiber-reinforced poly(methyl methacrylate) prostheses. A longitudinal multicentre study. Clin. Oral Implants Res. 6, 246 253. Berkovitz, B.K.B., Holland, G.R., Moxham, B.J., 2009. Oral Anatomy, Histology and Embryology, fourth ed. Mosby Elsevier, London. Boccaccini, A.R., 2005. Continuous fibre reinforced glass and glass-ceramic matrix composites. Handbook of Ceramic Composites. Springer US, pp. 461 484. Butterworth, C., Ellakwa, A.E., Shortall, A., 2003. Fibre-reinforced composites in restorative dentistry. Dent. Update 30 (2), 300 306. Callaghan, J.D., Vaziri, A., Nayeb-Hashemi, H., 2006. Effect of fiber volume fraction and length on the wear characteristics of glass fiber-reinforced dental composites. Dent. Mater. 22, 84 93. Davies, S.J., Gray, R.J., Linden, G.J., James, J.A., 2001. Occlusal considerations in periodontics. Br. Dent. J. 191, 597 604. Ferracane, J.L., 2006. Hygroscopic and hydrolytic effects in dental polymer networks. Dent. Mater. 22, 211 222. Fonseca, R.B., de Paula, M.S., Favarao, I.N., Kasuya, A.V.B., de Almeida, L.N., Mendes, G. A.M., et al., 2014. Reinforcement of dental methacrylate with glass fibre after heated silane application. Biomed. Res. Int. 2014, 364398. Freilich, M.A., Karmaker, A.C., Burstone, C.J., Goldberg, A.J., 1998. Development and clinical applications of a light-polymerized fibre-reinforced composite. J. Prosthet. Dent. 80 (3), 311 318.
314 Advanced Dental Biomaterials Freilich, M.A., Meiers, J.C., Duncan, J.P., 2002. Clinical evaluation of fibre-reinforced fixed bridges. J. Am. Dent. Assoc. 133, 1524 1534. Garoushi, S.K., Lassila, L.V., Vallittu, P.K., 2006. Short fibre reinforced composite: the effect of fibre length and volume fraction. J. Contemp. Dent. Pract. 7, 10 17. He, J., Vallittu, P., Lassila, L.V., 2017. Preparation and characterization of high radio-opaque E-glass fibre-reinforced composite with iodine containing methacrylate monomer. Dent. Mater. 33 (2), 218 225. Khan, A.S., Azam, M.T., Khan, M., Mia, S.A., Rheman, I.U., 2015. An update on glass fibre dental restorative composites: a systematic review. Mater. Sci. Eng. C. 47 (1), 26 39. Kleverlaan, C.J., Feilzer, A.J., 2005. Polymerization shrinkage and contraction stress of dental resin composites. Dent. Mater. 21 (12), 1150 1157. Mannocci, F., Ferrari, M., Watson, T.F., 1999. Intermittent loading of teeth restored using quartz fibre, carbon quartz fibre and zirconium dioxide ceramic root canal posts. J. Adhes. Dent. 1, 153 158. McCabe, J.F., Rusby, S., 2004. Water absorption, dimensional change and radial pressure in resin matrix dental restorative materials. Biomaterials 25, 4001 4007. Migliaresi, C., Alexander, H., 2004. Biomaterials Science, second ed. Academic Press, San Diego, CA, pp. 181 197. Nagata, K., Garoushi, S.K., Vallittu, P.K., Wakabayashi, N., Takahashi, H., Lassila, L.V.J., 2016. Fracture behaviour of single-structure fibre-reinforced composite restorations. Acta Biomater. Odontol. Scand. 2 (1), 118 124. Narva, K., 2002. Clinical and laboratory findings reinforcing denture base acrylic. In: The Third International Symposium on Fibre-Reinforced Plastics in Dentistry. pp. 113 124. Nelson DDS, MS, S.J., 2014. Wheeler’s Dental Anatomy, Physiology, & Occlusion, 10th ed. Elsevier, 31, 203. Nor, J.E., 2006. Tooth regeneration in operative dentistry. Oper. Dent. 31, 633 642. Petersen, R.C., 2005. Discontinuous fibre-reinforced composites above critical length. J. Dent. Restor. 84, 365 370. Ravindra, R., Mcilroy, P.J., Patel, S.A., 1997. Drugs that influence tubulin polymerization modulate thyrotropin-releasing hormone receptor number in at T-20 cells. Pharmacol. Toxicol. 80, 24 29. Rueggeberg, F.A., 2002. From vulcanite to vinyl, a history of resins in restorative dentistry. J. Prosthet. Dent. 87, 364 379. Schneider, L.F., Cavalcante, L.M., Silikas, N., 2010. Shrinkage stresses generated during resin-composite applications: a review. J. Dent. Biomech. 2010, 1 14. Shenoy, A., Shenoy, N., 2010. Dental ceramics: an update. J. Conserv. Dent. 13 (4), 195 203. Sideridou, I., Tserki, V., Papanastasiou, G., 2003. Study of water sorption, solubility and modulus of elasticity of light-cured dimethacrylate-based dental resins. Biomaterials 23, 655 665. Soderholm, K.J., Roberts, M.J., 1990. Influence of water exposure on the tensile strength of composites. J. Dent. Res. 69, 1812 1816. Soderholm, K.J., Zigan, M., Ragan, M., Fischlsweiger, W., Bergman, M., 1984. Hydrolytic degradation of dental composites. J. Dent. Res. 63, 1248 1254. Stickel, J.M., Nagarajan, M., 2012. Glass fiber-reinforced composites: from formulation to application. Int. J. Appl. Glass Sci. 3 (2), 122 136. Subramaniam, P., Babu, G.K., Sunny, R., 2008. Glass fibre reinforced composite resin as a space maintainer. A clinical study. J. Indian Soc. Pedod. Prevent. Dent. 26 (Suppl 3), S98 S103.
Fiber-reinforced composites 315 Tayab, T., Shetty, A., 2015. The clinical applications of fibber reinforced composites in all specialties of dentistry an overview. Int. J. Compos. Mater. 5 (1), 18 24. Tezvergil-Mutluay, A., 2017. Tooth as an adhesive substrate for fibre reinforced composites. In: Clinical Guide to Principles of Fibre-Reinforced Composites in Dentistry, vol. 1. Woodhead Publishing Series in Biomaterials, pp. 79 96. Vallittu, P., 2017. Types of FRCs used in dentistry. In: Vallittu, P., Özcan, M. (Eds.), Clinical Guide to Principles of Fiber-Reinforced Composites in Dentistry. Elsevier Science & Technology, Cambridge, pp. 11 32. Zhandarov, S., Mäder, E., 2005. Characterization of fibre/matrix interface strength: applicability of different tests, approaches and parameters. Compos. Sci. Technol. 65 (1), 149 160. Further reading Altair Enlighten, 2016. Fibre reinforced composites. ,https://altairenlighten.com/in-depth/ fiber-reinforced-composites/. (accessed 20.02.18.). Dyer, S.R., Lassila, L.V., Jokinen, M., Vallittu, P.K., 2004. Effect of fibre position and orientation on fracture load of fibre-reinforced composite. Dent. Mater. 20, 947 955. Tucker, A., Sharpe, P., 2004. The cutting-edge of mammalian development; how the embryo makes teeth. Nat. Rev. Genet. 5 (7), 499 508. Zhang, Y.R., Du, W., Zhou, X.D., Yu, H.Y., 2014. Review of research on the mechanical properties of the human tooth. Int. J. Oral Sci. 6 (2), 61 69.
Zirconium in dentistry 14 Rafael Pino Vitti1,2, Anderson Catelan3, Marina Amaral1 and Rafael Rocha Pacheco4 1 Department of Dentistry, University of Taubaté, Taubaté, Brazil, 2Department of Dentistry, Araras Dental School (FHO|UNIARARAS), Araras, Brazil, 3Faculty of Health Sciences, University of Western São Paulo, Presidente Prudente, Brazil, 4 School of Dentistry, University of Detroit Mercy, Detroit, MI, United States Chapter Outline 14.1 Introduction 317 14.2 Classification 321 14.2.1 14.2.2 14.2.3 14.2.4 14.2.5 Feldspathic ceramics 323 Leucite-based ceramics 323 Lithium disilicate based ceramics 323 Alumina-based ceramics 323 Zirconia-based ceramics 324 14.3 14.4 14.5 14.6 Zirconia in dentistry 324 Yttrium-stabilized tetragonal zirconia 326 Zirconia-toughened alumina 328 Surface topography, clinical treatments of zirconia surface, and adhesion to zirconia in dental restorations 329 14.7 Failure and fractographic analysis of zirconia restorations 333 14.8 Mechanical testing of zirconia ceramics 335 14.9 Limitations and challenges 337 Further reading 338 14.1 Introduction Restorative dentistry has always been an essential area in the dental field and has grown recently due to higher esthetic demands from society. The term “restorative” implies the act of performing a “restoration,” which can be defined by the action of returning something to a former owner, place, or condition. In the dental field, a “restoration” can be defined as a structure provided to replace or repair dental tissue to restore its form and function. Each tooth is one individual organ formed by different tissues, with different properties. Hard tissues present in the dental structure, such as enamel and dentin, also differ regarding properties such as mechanical, physical. Clinicians have used many restorative materials as an attempt at “restoring” these original tooth properties, with extensive research and development. Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00014-1 Copyright © 2019 Elsevier Ltd. All rights reserved.
318 Advanced Dental Biomaterials Restorative dental materials are usually classified as ceramics, polymers, composites, or metals. Ceramics are inorganic, nonmetallic materials widely used in different applications. Glass is a subset of ceramics, known by its elevated hardness and translucency, properties that are similar to dental enamel. Due to these desired properties, one of the first ceramics used in dentistry was the feldspathic ceramic, a glass obtained from feldspar, one of the most common minerals present on earth. These ceramics can be processed through traditional laboratorial condensation methods followed by sintering. The sintering process is used to produce densitycontrolled materials by applying thermal energy, resulting, in this particular case, in a glassy matrix containing leucite crystals. The amount and distribution of these crystals can modify the mechanical and physical properties of the ceramic. An increased content of leucite crystals can lead to increased flexural strength, increased hardness, and increased thermal expansion. However, it also reduces translucency. Due to its low flexural strength, the use of feldspathic ceramic is limited to areas with lower stress incidences. Furthermore, these materials can be processed through pressing or milling methods. Due to the low flexural strength (approximately 65 MPa) observed for feldspathic ceramics, the indications for these materials are limited to laminates, and veneering of infrastructure/frameworks. Contrarily to the feldspathic ceramic, metals are known to present increased ductility/malleability, resulting in increased plastic strain and, consequently, increased toughness. The association of feldspathic ceramics to different alloys results in fullcoverage restorations with increased flexural strength and indications for higher load bearing regions. The porcelain fused to metal (PFM) restorations represented an important moment in dentistry since the range of procedures where ceramics could be used were significantly increased. A metallic core (made with specific alloys) is created by using the well-established lost-wax technique with room for the feldspathic ceramic to be applied later. The compatibility of physical properties between alloy and ceramic is essential for a successful restoration. The fusion temperature for the ceramic cannot be higher than the melting temperature of the alloy. In addition, differences in the coefficient of thermal expansion could lead to stress concentrations and future failure of the restoration. The crystalline portion of the ceramic determines the coefficient of thermal expansion and mechanical properties, allowing the properties to be manipulated in such a way that these materials are compatible with the selected alloy. The feldspathic ceramic is applied to the pretreated surface of the alloy and, after sintering, both materials are intimately bonded by the mechanical retention caused by the interpenetration of the ceramic in the micromechanical retentions of the alloy surface, chemical bonds between ceramic and metallic oxides on the surface of the alloy, and stresses caused by the shrinkage of the ceramic during cooling. This intimate contact allows increased stress distribution from the ceramic to the metallic core. PFMs can be indicated for full crown procedures, fixed dental prosthesis (FPD), and are still extensively used in dentistry, especially for the restoration of implants. However, the metallic core of PFMs can lead to reduced esthetic properties due to the reduced translucency of the alloy (when compared to the ceramic and dental structure) and limitations caused by the metallic appearance of the core.
Zirconium in dentistry 319 Furthermore, with time, there is the possibility of exposure of cementing line and corrosion of the alloy which could result in increased concerns regarding esthetics. Due to the previously mentioned constant growing esthetic demand, there was the need for restorations that would match the shade of the dental element regarding all color parameters (hue, chroma, and value). Value is significantly affected by the translucency of the restorative material; thus the metallic core can negatively affect the final restoration aspect. As an attempt to remove the metallic core without impacting the mechanical properties, dental ceramics were reinforced by an increased crystalline content, resulting in metal-free restorations. An increased leucite content leads to higher flexural strength (160 MPa) increasing the range of indications when compared to feldspathic ceramic: veneers, laminates, full anterior crowns, anterior full crown infrastructure, inlays, and onlays. Leucite-reinforced ceramics are considered glass-ceramics and can be processed by conventional condensation methods (also used for PFMs), pressed, or milled. There are limitations for the incorporation of leucite crystals, which impacts the strength and esthetics of the ceramic. For a ceramic to be considered as leucite reinforced, it needs to contain leucite at a volume of at least 35%. The incorporation of different crystals, such as lithium disilicate, allowed a new perspective on the research and development of dental ceramics. This crystal, in particular, resulted in a type of reinforced glass-ceramic with increased flexural strength (350 MPa), when compared to other glass-ceramics, with excellent esthetic properties and adhesive potential. These properties increased the number of indications for this material including, but not limited to veneers, laminates, full anterior crowns, anterior full crown infrastructure, posterior full crowns, posterior full crown infrastructure, crowns over implants, abutments, inlays, onlays, and up to three-element FPD. Lithium disilicate reinforced ceramics contain, approximately, 70% of lithium disilicate crystals in their composition and can be processed through pressing methods and are also available as CAD/CAM (computer-assisted design/computerassisted manufacturing) blocks. These blocks can be found presintered or fully sintered. A great advantage for glass-ceramics over other restorative materials is the possibility of being subjected to surface treatments, allowing the modification of the surface regarding surface area and energy, and the possibility of silanization. Silane is a bifunctional molecule (silanol and methacrylate groups) that creates a link between silica and the polymeric matrix of resin-based materials, such as adhesive systems and resin cements. The modern practice of dentistry is based on a conservative approach, usually referred to as minimally invasive dentistry, which tends to preserve as much sound dental structure as possible. The introduction of adhesive dentistry allowed clinicians to perform minimally invasive procedures by bonding these glass-ceramics to the dental tissues. However, due to the brittle nature of glass-ceramics (lack of deformation before fracture), preparations for these restorations can result in increased amount of reduction of the dental structure, in order to provide adequate thickness for the restorative material and, consequently, increased mechanical strength. The selection of a glass-ceramic as the restorative material for a full-coverage crown would require a minimal thickness for the restoration of, approximately,
320 Advanced Dental Biomaterials 1.5 mm, whereas for metallic restorations the minimal thickness is of 0.5 mm. The option for a metallic crown could result in minimal reduction of the dental structure, however, without the esthetics required for anterior procedures and the disadvantage of not bonding these materials to the tooth structure. Polycrystalline ceramics have been used as an attempt to replace the metallic structures without drastically reducing the flexural strength. White metallic oxides such as alumina (aluminum oxide) and zirconia (zirconium oxide) have been used as crystals in ceramics that are fabricated through processing methods of slipcasting and/or milling. These materials present a reduced (to none) volume of glass, resulting in elevated flexural strength. However, due to its nature and morphology, these materials demonstrate reduced translucency and increased hardness, when compared to any glass-ceramic. Alumina (aluminum oxide) is commonly used in dentistry as an abrasive due to its high elastic modulus (300 GPa). Alumina-based ceramics contain approximately 99.5% of alumina crystals. The reduced volume of glass (or the absence of) leads to a high elastic modulus and reduced translucency. Due to that fact, these ceramics require veneering and are not acid sensitive, which represents a limitation regarding adhesion. The indication of this material is limited to infrastructures for the crown and three elements anterior bridge (requires veneering). The traditional processing method used for this ceramic is through slipcasting, which is a time-consuming and technique-sensitive procedure, yet some blocks are available for CAD/CAM. The development of new materials and technologies, added to the limitations of alumina-based materials, has reduced the use of these materials. One particular polycrystalline ceramic has drawn interest due to the potential of combining the minimal reduction with acceptable esthetic properties. Due to its composition and morphological structure, which will be discussed further in this chapter, zirconia tends to demonstrate increased mechanical properties when compared to any other ceramic. The increased flexural strength and, consequently, fracture toughness allow the restorations to be fabricated with reduced thickness. Thus preparations do not require as much reduction as the ones for glass-ceramics. Higher flexural strength also allows this material to replace the metallic core of PFM crowns, creating a new type of restoration named as porcelain fused to zirconia (PFZ). The mechanical properties of zirconia lead to a wide range of indications including, but not limited to PFZ core, monolithic crown, FPD frameworks, implants, implant abutments, and dentures. However, the properties observed for zirconia are not necessarily similar to the ones of dental tissues (flexural strength, translucency, and hardness). The absence of glass on the composition also makes bonding to these materials a challenge. Modifications have been made to obtain optimal properties, such as the association of zirconia to other types of ceramics, such as alumina-toughened zirconia (ATZ) or zirconia-toughened alumina (ZTA). One example of ATZ ceramic ( . 50% of zirconia crystals) is the combination of 67.9% of ZrO2, 21.5% Al2O3, 19.6% CeO2, 0.06% MgO, and 0.03% TiO. Recently, promising new types of glass-ceramics have been released to the dental market as an attempt at overcoming the limitations of lithium disilicate based ceramics and reducing milling/processing time. Both classes of materials include
Zirconium in dentistry 321 Figure 14.1 Fully crystallized glass-ceramic milling block. Table 14.1 Classification of dental ceramics according to fusion temperature and indications. Classification Fusion temperature ( C) Indications High fusing .1300 Medium fusing Low fusing Ultralow fusing 1101 1300 Denture teeth, and fully sintered alumina and zirconia core ceramics Denture teeth and presintered zirconia 850 1100 ,850 Crown and bridge veneering Crown and bridge veneering addition zirconia or alumina to the crystalline portion of lithium silicate-based ceramics (Fig. 14.1). The constant development and advancements in CAD/CAM technology and its applications in dentistry have increased the need for new restorative materials. Resin-based CAD/CAM blocks with increased inorganic content have been introduced in the dental field as an alternative to the ceramic blocks, named as resin-matrix ceramics (nanoceramics). 14.2 Classification Dental ceramics can be classified according to the volume and composition of crystalline content, and fabrication methods, including fusion temperature (Table 14.1). In terms of crystalline content (Fig. 14.2), there are, essentially, predominantly glass or polycrystalline ceramics. The third type of ceramic in that classification would be the so-called glass-infiltrated ceramics which are processed through the slip-casting technique. As the name already describes, these ceramics (usually
322 Advanced Dental Biomaterials Figure 14.2 Classification of dental ceramics according to crystalline content. Figure 14.3 Classification of dental ceramics according to the composition. alumina-based) are, basically, a porous crystalline structure interpenetrated by the glass. Predominantly glass-ceramics are considered acid-sensitive since its surface can be modified by etching procedures using hydrofluoric acid, while polycrystalline ceramics are considered acid resistant. The acid interacts with the silica on the surface of the ceramic, promoting an increase in the surface area and, consequently, surface energy. Furthermore, the silanol group present on the silane molecule bonds chemically to the silica present on the ceramic surface, while the methacrylate group is free to bond to other dental polymeric materials, creating a link between the ceramic and adhesive system/resin cement. Dental ceramics vary in terms of composition (Fig. 14.3), especially regarding their crystalline content. Feldspathic ceramic is a predominantly glass-ceramic (K2O Al2O3 6SiO2) used mostly for restoring enamel tissue, due to its optical properties, and veneering of infrastructures and frameworks. For processing metalfree restorations, these ceramics needed to be reinforced by an increased crystalline content, resulting in higher flexural strength. The reinforced ceramics can be divided by the type of crystal used in their composition which can be, mainly, leucite (K2O Al2O3 4SiO2), lithium disilicate (Li2O 2SiO2), alumina (Al2O3), or zirconia (ZrO2).
Zirconium in dentistry 323 14.2.1 Feldspathic ceramics 14.2.1.1 Commercial examples G G G Condensation: VITAVM 7 (Vita Zahnfabrik); VITAVM 9 (Vita Zahnfabrik); VITAVM 11 (Vita Zahnfabrik); VITAVM 13 (Vita Zahnfabrik); VITAVM 15 (Vita Zahnfabrik); VITA VMK Master (Vita Zahnfabrik); IPS e.max Ceram (Ivoclar Vivadent); Ceramco (Dentsply); CERABIEN ZR (Noritake); Super Porcelain EX-3 (Noritake); Cercon Ceram Kiss (Dentsply); Lava Ceram (3M ESPE) Pressing: IPS e.max ZirPress (Ivoclar Vivadent); VITAPM 9 (Vita Zahnfabrik); CERABIEN ZR PRESS (Noritake); Super Porcelain EX-3 PRESS (Noritake); Cercon Ceram Press (Dentsply) Milling: VITABLOCS TriLuxe forte (Vita Zahnfabrik); VITABLOCS RealLife (Vita Zahnfabrik); VITABLOCS Mark II (Vita Zahnfabrik); VITABLOCS Esthetic Line (Vita Zahnfabrik); CEREC Blocs C (Dentsply Sirona) 14.2.2 Leucite-based ceramics 14.2.2.1 Commercial examples G G G Condensation: IPS Empress Esthetic Veneer (Ivoclar Vivadent); IPS Empress In-Line (Ivoclar Vivadent); Fortress (Mirage Dental Systems); IPS d.SIGN (Ivoclar Vivadent), Avanté Micro Crystal (Pentron) Pressing: IPS Empress Esthetic (Ivoclar Vivadent); Cergo Kiss (DeguDent); Ceramco iC (Dentsply) Milling: IPS Empress CAD (Ivoclar Vivadent); Paradigm C (3M ESPE); Kavo Everest G-Blank (KaVo); Rosetta BM (HASS) 14.2.3 Lithium disilicate based ceramics 14.2.3.1 Commercial examples G G Pressing: IPS e.max PRESS (Ivoclar Vivadent); Suprême.press (Axsys); Amber Press (HASS); Rosetta SP (HASS); Amber Li-Si-POZ (HASS) Milling: IPS e.max CAD (Ivoclar Vivadent); Suprême.cad (Axsys); Amber Mill (HASS); Rosetta SM (HASS). Lithium aluminosilicate reinforced with lithium disilicate: n!ce (Straumann) Zirconia-reinforced lithium silicate: Suprinity (VITA Zahnfabrik); Celtra Duo (Dentsply Sirona) G G 14.2.4 Alumina-based ceramics 14.2.4.1 Commercial examples G G Slip-cast technique: VITA In-Ceram ALUMINA (Vita Zahnfabrik); VITA In-Ceram SPINELL (Vita Zahnfabrik) Milling: VITA In-Ceram ALUMINA for inLab (Vita Zahnfabrik); Procera AllCeram (Nobel Biocare)
324 Advanced Dental Biomaterials 14.2.5 Zirconia-based ceramics 14.2.5.1 Commercial examples G Milling: VITA YZ XT (Vita Zahnfabrik); VITA YZ ST (Vita Zahnfabrik); VITA YZ HT (Vita Zahnfabrik); VITA YZ T (Vita Zahnfabrik); AxZir HT Zr (Axsys); AxZir ST Zr (Axsys); AxZir TT Zr (Axsys); Zirtooth (HASS); NobelProcera (Nobel Biocare); Cercon (Dentsply Sirona); Celtra Duo (Dentsply Sirona); Zerion (Straumann); Zolid (Straumann); IPS e.max ZirCAD (Ivoclar Vivadent) 14.3 Zirconia in dentistry The zirconia is a biomaterial used in biomedical applications since the 1960s. Among the dental ceramics, zirconia or zirconium dioxide (ZrO2) has approximately 96% 99% crystalline content with no glassy phase, presenting high physicochemical properties, such as flexural strength, fracture toughness, and hardness, increased mechanical properties, excellent biocompatibility, and satisfactory esthetic, thus expanding its application considerably in dentistry. Due to the increasing interest in esthetics and concerns about cytotoxicity and allergic reactions to certain metals, patients and dentists have been looking for metal-free restorations. Although there are many types of zirconia-based ceramic systems available, currently only two types are commonly used in dentistry: yttrium-stabilized tetragonal zirconia (Y-TZP) and ZTA. Zirconia structures used in dentistry treatments are fabricated using CAD/CAM in two different methods. In one method, the fully sintered block of zirconia is milled, and no distortion (shrinkage) occurred in the final structure. The main disadvantage is the wear promoted by burrs and, consequently, flaws produced during the machining that can compromise the mechanical property of the zirconia structure. The other method consists of milling the zirconia structure from a presintered block. In this method, the final mechanical property is reached after the sintering process, and the structural shrinkage can be partly compensated at the drawing phase improving the fit of the zirconia restoration. The zirconia can be used as base material for feldspathic ceramic (porcelain) restorations where the zirconia core structure is layered with porcelain and is mainly used for anterior teeth. Full zirconia restorations are used in posterior teeth where no esthetic porcelain layered on top is necessary. The bond strength of layered PFZ is critical. Thus crowns and bridges may be entirely constructed of the zirconia ceramic in situations where esthetic demands are not extremely high. Zirconium (Zr) is a solid metal with similar physicochemical properties to titanium. Thus zirconia also can be used in dental implants as abutments attached to the implants, and it provides some advantages over alumina in order to solve the problem of brittleness and the potential failure of implants. It is important to point out that the fatigue performance of zirconia abutments is dependent on its abutment diameter.
Zirconium in dentistry 325 However, the clinical success of zirconia implants is also dependent of the feldspathic ceramic survival. The long-term stability of the zirconia implant is questionable because the implant failures are related with zirconia fractures, which can occur through laboratory technical complications, and/or feldspathic veneering ceramics fractures. Recently, a single body implant in zirconia (monotype zirconia implant) was also developed. In this way, an entire implant reconstruction will be metal free. Titanium implants are widely used in implants treatment, but their disadvantages (unwanted chemical biological interactions with the surrounding soft and hard tissues and discoloration) can promote unsatisfying esthetical results, mainly in patients with thinner mucosal biotype, gingival recession, or high smile line. However, there are limited studies about zirconia in dental implants. Thus the clinical success rate of these devices is not entirely known. Table 14.2 shows all dental applications where the zirconia can be utilized. For restorations of endodontically treated anterior teeth, metallic posts may result in unfavorable esthetic results, such as a gray discoloration of ceramic crowns and the surrounding gingival margin. Further, corrosive reactions with prefabricated posts may cause complications involving the surrounding tissues. These concerns promoted the development of esthetic posts made of zirconia and other ceramic materials. Besides esthetic properties, some studies demonstrated that the zirconia posts have higher mechanical strength than those reported for other types of ceramic posts. In addition, zirconia posts also possess some advantages with respect to biocompatibility but present some limitations, such as stiffness and low ductility. Thus dentists found difficulties in roots with small sizes and in root canal retreatment procedures. The esthetic orthodontic brackets can be manufactured using zirconia. As mentioned above, zirconia brackets are also another option to alumina ceramic brackets, because they have the greatest toughness among all polycrystalline ceramics. Moreover, in general zirconia brackets are cheaper than alumina ceramic brackets. However, zirconia has high opacity making the orthodontic brackets less esthetically pleasing. Studies show no differences regarding frictional properties between zirconia and alumina brackets, but higher sliding properties have been reported for zirconia brackets over both stainless steel and nickel titanium archwires with reduced plaque adhesion, besides of the clinically acceptable bracket-tooth bond strength. Table 14.2 Zirconia in dentistry. Materials/procedures Abutments Composite resins Cores Crows (full and partial) Implants Orthodontic brackets Posts Veneers
326 14.4 Advanced Dental Biomaterials Yttrium-stabilized tetragonal zirconia Zirconia is obtained by minerals such as baddeleyite and zircon (ZrSIO4), and it has both basic and acidic chemical properties. Unalloyed zirconia is a polymorph that assumes three crystallographic forms depending on the temperature: monoclinic (warped prism with parallelepiped borders), tetragonal (straight prism with rectangular borders), and cubic (straight prism with squares borders). The zirconia has a monoclinic crystalline structure at room temperature and is stable up to 1170 C. In this form, the zirconia shows reduced mechanical properties as well as cohesive strength (low density). Between 1170 C and 2370 C the zirconia has a tetragonal structure and allows a ceramic with improved mechanical properties. Above 2370 C, the zirconia crystals have a cubic shape with intermediate mechanical properties. The zirconium oxide can be stabilized with the addition of approximately 2 3 mol.% of yttrium oxide (Y2O3). After its processing, during the cooling, the tetragonal phase again becomes monoclinic, and this transformation results in a substantial volume expansion of approximately 3% 4.5%. Zirconia exposure to water and humidity also triggers spontaneous phase transformation, creating residual compression stress even without mechanical surface change or load application. However, the tetragonal phase at room temperature may be stabilized by oxides such as calcium, magnesium, cerium, and yttrium, controlling crystal phase transformation. Thus yttrium oxide is added to zirconia in order to maintain a metastable phase at temperatures below the tetragonal-to-monoclinic transformation temperature, creating a fully tetragonal zirconia ceramic made of small metastable tetragonal grains (Y-TZP). This procedure depends on the processing temperature, the yttrium oxide content, and the grain size, and it improves the mechanical strength of the zirconia, because when the Y-TZP is induced by tensions that initiate the propagation of a crack, the tetragonal crystals close to the crack become the stable monoclinic phase (Fig. 14.4). The metastable phase exhibits good crack propagation resistance since it is associated with a volume expansion of 3% 5% of the particles that cause stresses of compression that oppose the growth and propagation of the crack. Thus extra energy is required in order to propagate the crack. Biomedical applications are focused in Y-TZP characterized by the addition of approximately 2 3 mol.% of yttrium oxide tetragonal fine-grained microstructures. For this reason, Y-TZP has physicochemical properties of great interest to dentistry, such as high strength, fracture toughness, hardness, high frictional behavior, wear and corrosion resistance, and low thermal conductivity (Table 14.3). Besides the physicochemical properties, Y-TZP is a chemically inert material and no negative reactions (cytotoxic, oncogenic, or mutagenic effects) in biological tissue have been reported. Several studies demonstrated the high biocompatibility of Y-TZP with cytotoxicity results similar to alumina and lower than titanium dioxide. The improved biocompatibility of Y-TZP implants is due to strong bone response and low bacterial colonization. Furthermore, the Y-TZP implants present long-term osseointegration stability without biomechanical problems as well as titanium ones.
Zirconium in dentistry 327 Figure 14.4 Tetragonal monoclinic transformation phase during crack propagation. Table 14.3 Physicochemical, electrical, and thermal properties of yttrium-stabilized tetragonal zirconia. Properties Mean values Properties Mean values Compressive strength Density Dielectric constant Elastic modulus Flexural strength Fracture toughness Latent heat of fusion Melting point 4900 MPa 6.05 g/cm3 16.5 200 GPa 900 1600 MPa 9.5 MPa/m 760 kJ/kg 2625 C 0.27 3.16 3 1027 Ω m 600 MPa 2.1 2.2 70 GPa 413 MPa 2 W/m K 11 3 1026 Porosity ,0.05% Poisson’s ratio Resistivity Rupture modulus Refractive index Shear strength Tensile strength Thermal conductivity Thermal expansion coefficient Vickers hardness 1200 kgf/mm2 The optical property is another important factor to be considered in dental restoration manufacturing. Y-TZP has a high refractive index, high opacity, and low absorption coefficient (Table 14.3). The Y-TZP transmission coefficient (400 700 nm wavelengths) is higher than alumina and ZTA ceramics, but lower than densely sintered alumina and lithium disilicate ceramics. Y-TZP contains a highly crystalline phase; this fact confers opacity of the indirect restoration; since being a white material it needs to be veneered most often with porcelain-based esthetic ceramics, serving as prosthesis infrastructure. The high opacity is useful in
328 Advanced Dental Biomaterials some clinical situations, such as masking abutment teeth or metallic posts/cores. The Y-TZP opaque effect is predicated on dispersed particles slightly greater in size than the wavelength of light. Thus the size of the crystalline grain is the microstructural characteristic more closely related to the adjustment of opacity of Y-TZP ceramics. Lager grains lead to a smaller number of grain borders, decreasing light scattering, but larger grains are harmful to mechanical properties and the stability of the tetragonal phase of Y-TZP. Therefore the adjustment of the Y-TZP opacity could not be obtained by increasing its grain size. Moreover, the Y-TZP tetragonal phase has different refractive indexes according to its crystallographic orientation in the microstructure. This anisotropic behavior is related to different refractive indexes and causes significant light scattering (birefringence phenomenon). Increased esthetic properties of zirconia by sintering temperature change and fabricating processes were previously reported, mainly its translucency. In this same way, monolithic zirconia appears as a good alternative for clinical use, showing excellent mechanical properties, even at minimum thicknesses, and adequate translucency, and avoiding delamination and chipping failures of the porcelain layer due to poor bonding between the zirconia framework and the veneering porcelain. Some manufacturers introduce different zirconia colors to better esthetic outcomes of the white shade zirconia frameworks, which can be further esthetically customized by veneering with feldspathic ceramic using layering or pressing techniques. Both methods require some sort of feldspathic ceramic for zirconia bonding. Several studies show some chemical elements which may contribute to this bonding. Recently, among polycrystalline ceramics, Y-TZP for monolithic restorations has been developed to overcome problems related to chipping of feldspathic ceramic layers applied over zirconia. The Y-TZP high density and low residual porosity (Table 14.3) provide a satisfactory level of opacity, which can mask underlying discolored abutments and allow a controlled depth of translucency after veneering even with restricted core thickness (0.5 mm). Moreover, Y-TZP restorations do not affect visible color changes with clinical situations with 2 3 mm thick mucosa, with or without the application of veneering ceramic. The Y-TZP is exposed to several surface flaws as a result of the pressing or milling methods and other laboratory procedures. Besides its microstructure, the mechanical properties of Y-TZP ceramic also depend on the defects produced during the manufacturing methods. Impurities in zirconium dioxide (ZrO2) particles, as well as differences in Y-TZP processing, can affect the stability of the tetragonal phase. There is a correlation between flexural strength and any surface damage as shown by surface roughness. Thus the mechanical properties and long-term stability of Y-TZP restorations can be quite different than those of starting ceramic. 14.5 Zirconia-toughened alumina The introduction of zirconia up to 10 vol.% into an alumina matrix results in a class of ceramic material named as ZTA. The ZTA ceramics are used in several fields and especially in dentistry (dental restorations) because it exhibits transformation
Zirconium in dentistry 329 toughening at high temperature, which improves the fracture toughness, mechanical strength, and stability. Alumina is one of the hardest ceramics and, thus, has low surface wear. However, with a high hardness, the chance of brittle fracture is increased. The addition of Y-TZP to the alumina matrix to form ZTA improves the mechanical properties of fracture toughness (6 12 MPa/m) and flexural strength of alumina, decreasing the risk of brittle fracture and improving the efficiency and reliability of the ceramic restoration. Thus the combination of alumina and zirconia allows compensating for the moderate toughness of alumina and the aging effect of zirconia. ZTA ceramic exhibits not only a high toughness value but also, more importantly, a greater threshold for the stress coefficient, decreasing the crack propagation. Furthermore, alumina has a lower susceptibility to assisted corrosion by water or body fluid. Then, ZTA has higher hydrothermal stability than Y-TZP ceramic. The Y-TZP particles are associated with tetragonal monoclinic transformation, increasing the fracture toughness of the ceramic. For alumina, toughening is due to grain bridging in the direction of the propagating crack, which shields the ceramic from the applied load. Moreover, there is the difference in elastic modulus between the alumina matrix and the zirconia particles, so the cracks will tend to move across the zirconia grains, inducing their tetragonal-to-monoclinic transformation and dissipating the crack energy. Thus ZTA ceramics can be used as infrastructure materials for crowns and fixed-partial dentures, improving the lifetime and reliability of ceramic prostheses. 14.6 Surface topography, clinical treatments of zirconia surface, and adhesion to zirconia in dental restorations The glass-free dense polycrystalline microstructure of zirconia ceramic makes the material chemically inert, and due to the absence of silica in its composition, conventional adhesive luting procedures using hydrofluoric acid etching and silanization are not recommended for this material. Therefore high bond strength and bonding durability to zirconia in the long term is still a challenge. Although conventional cements, such as zinc phosphate and glass ionomer, had been initially recommended for zirconia restoration luting, resin-based luting agents provide higher retention with better physicochemical properties than conventional ones, with increased fracture resistance, fatigue resistance, and long-term durability of ceramic restorations. Thus several mechanical and chemical treatments on zirconia surface are proposed. In addition, after clinical evaluation of ceramic restoration for occlusion adjustment and contact with adjacent teeth, saliva, blood, try-in paste, silicone indicators, and other contaminants may be present on the internal surface of indirect restoration and lead to a reduction in the bond strength. Water-spray rinsing alone is not sufficient to remove saliva contaminants; so chemical cleaning using solvent, emulsion, acidic, alkaline, and neutral products has been recommended for adequate cleaning.
330 Advanced Dental Biomaterials Among the main cleaning agents, alcohol, acetone, sodium hypochlorite, suspension of zirconium dioxide particles, and phosphoric acid gel are the most used for zirconia decontamination. However, alcohol was not effective in removing salivary organic contaminants. Already acetone was effective in removing silicone contamination but was not effective in removing salivary contaminants. Although sodium hypochlorite is a nonspecific proteolytic solution capable of removing organic material, resin polymerization may be impaired due to oxygen generation. A commercial alkaline paste of zirconium dioxide suspension removes salivary phosphates by adsorption, but residual agglomerated zirconium dioxide particles on the zirconia surface may be associated with slight bond strength reduction or bonding effectiveness of uncontaminated zirconia. Commonly 30% 40% phosphoric acid gel is recommended to remove contaminants of the bonding surface of prosthetic restoration. Phosphoric acid was an effective method to remove salivary contamination, but it reduces surface free energy and residual phosphate groups could compete with the binding sites of 10methacryloyloxydecyl dihydrogen phosphate (MDP) monomer on the hydroxyl groups of zirconia surface, reducing the bond strength. Therefore additional airborne-particle abrasion after try-in procedures removes contaminants successfully and improves bond strength durability in the long term. Previous studies showed that aluminum oxide airborne-particle abrasion caused an increase in the surface roughness, and surface free energy, increasing the bonding area available for mechanical interlocking and surface wettability of zirconia ceramic. Air abrasion removes mechanically the superficial ceramic and exposes a fresh bonding area, improving the bond between rougher zirconia and resin-based materials and between zirconia and veneering porcelain by micromechanical retention. Alumina grain size used in the air-abrasion particles method varies from 25 to 250 μm applied under 0.05 0.45 MPa for 5 30 seconds at a distance of 5 20 mm between the nozzle of the sandblaster and ceramic restoration. Unfortunately, it has been reported that alumina airborne-particle abrasion could cause surface microcracks and activate the transformation of zirconia crystals from tetragonal-to-monoclinic phase at the surface and subsurface of zirconia and should be used with caution since that excessive abrasion may weaken and decrease the mechanical properties of zirconia material. For this reason, manufacturers suggest the use of alumina abrasion before final sintering and/or the heat treatment after this procedure to reverse the tetragonal-to-monoclinic phase. In addition, air-abrasion particles of 50 μm, air-abrasion using the regular pressure of 0.25 MPa, or the combination of low pressure of 0.05 MPa and priming with MDP-containing primers, at 10 mm of distance from the bonding surface, seem to be useful to achieve durable bond strength, which could reduce the chance of phase transformation. Surface grinding methods using disks and diamond rotatory instruments were tested in order to increase roughness and optimize bond strength. However, due to the high hardness of zirconia coarse diamond burrs are necessary to produce a rougher surface and adequate bonding of luting cements to zirconium dioxide ceramics, but this procedure may induce microcracks, decrease mechanical properties,
Zirconium in dentistry 331 increase crystalline phase transformation, and cause damage to zirconia surface. So proximal and occlusal adjustments should be performed with caution and under constant water spray to avoid heating. Tribochemical silica coating consists of airborne-particle abrasion of aluminum oxide particles modified by silica, promoting a roughened surface that is coated with silica and receptive to silanization treatment. It is reported that this technique is better compared to conventional air-abrasion particle. Silica coating improves bond strength due to micromechanical interlocking by the increased bonding area and surface free energy, but ultrasonic cleaning destroys this effect. Moreover, silicatized alumina followed by application of silane coupling agent promotes chemical adhesion by the formation of silanol groups, which react with the silica layer on the inner surface of the zirconia restoration. Silica coating and silanization improved long-term resin bonding to zirconia ceramics and water resistance. In dentistry, lasers are used to cut dental hard tissues, in the same way, the laser may modify the material surface, including zirconia ceramic. Laser irradiation represents an effective technique for conditioning zirconia surfaces, enhancing micromechanical retention, and improving the bond strength of resin cement on zirconia ceramic and porcelain zirconia bonding, due to the increase in the surface roughness and optimizing wettability. Carbon dioxide (CO2) laser treatment is reported to be capable of increased bond strength for irradiated zirconia. Erbiumdoped yttrium aluminum garnet and neodymium-doped yttrium aluminum garnet etching also improved the bond strength. However, the use of these lasers may induce microcracks formation, cause silver spot at application point, and increase the amount of monoclinic phase. The titanium/sapphire-based laser with ultrashort femtosecond pulses seems to be a promising technique to provide necessary surface roughness, increase adhesive efficiency of zirconia resin interface, remove contaminants with minimum thermal and mechanical damage, and without phase transformation. Another nondestructive method to improve zirconia bonding is the use of plasma technology by cleaning, etching, surface functionalization, and deposition of thin functional films. For dental purposes, the application of nonthermal atmospheric pressure plasma has been commonly used for surface modification of biomaterials by the emission of plasma-containing ions, electrons, and molecules in an unbalanced nature. Silica vapor nanofilm is deposited on zirconia surface by reactive magnetron sputtering, making it more chemically reactive. Film deposition of silica precursors, such as hexamethyldisiloxane and tetramethylsilane, is also responsible for improving adhesive bonding. Plasmatized elements and hydrophilic properties of zirconia are enhanced after decomposition and vaporization processes, resulting in the reduction of contact angle and improvement of wettability of argon and methane plasma. Also, plasma application increases the surface energy and removes contaminants, once C C and C H bonds can be broken; as well as enhancing chemical bonding by van der Waals secondary bonds. After atmospheric pressure plasma treatment, the polycrystalline structure of zirconia ceramic remains unchanged, with the adhesive interface having waterresistant behavior.
332 Advanced Dental Biomaterials Selective infiltration etching on the structural integrity of the zirconia ceramic consists of heating-induced maturation and grain boundary diffusion, which in contact with a thin layer of low temperature melting glass, transforms the zirconia surface from a relatively nonretentive into a highly retentive surface, establishing a strong and durable bonding to zirconia. Application of a thin layer of vitreous porcelain glaze, composed basically of high silica content, amorphous matrix, or silicon dioxide, results in the vitrification of the ceramic surface that promotes adequate bond strength. Glazed zirconia makes this surface etchable by hydrofluoric acid and subject to silanization, a condition similar to silica-based ceramics. Therefore low-fusing porcelain glaze application on zirconia surface seems to be a promising method for improving the bond strength to the resin cement and optimizes zirconia porcelain adhesion and fracture resistance. However, coated zirconia surfaces with porcelain glaze should not be treated with hydrofluoric acid for too long, because the etching may completely remove the glassy film from the surface, to approximately 12 μm of thickness. Moreover, glass coating in the inner surface of a metal-free prosthesis forms a layer thick enough to interfere in the seating of the zirconia restorations, but within clinically acceptable marginal discrepancies. Hydrofluoric acid is considered an efficient agent capable of dissolving the glassy matrix, but zirconium dioxide ceramics are acid resistant to typically etching using low concentration (4% 10%) hydrofluoric acid. Therefore it is unable to create microretention on the crystalline surface of zirconia, without glass content. In this way, several studies have evaluated different concentrations, application times, ultrasonic and hot etching. Highly concentrated (above 40%) hydrofluoric acid applied on zirconia surface under medium-to-long time at room temperature has shown increased wettability, surface roughness, morphologic irregularities, and bond strength to the resin-luting materials. Hot hydrofluoric acid etching enhances surface roughening to improve resin zirconia bonding; even at low concentration (9.5%) for a short duration time (60 seconds), it may create roughness, making the etching safer due to the hazardous nature of hydrofluoric acid. Unfortunately, zirconia surfaces etched with hydrofluoric acid induce tetragonal-to-monoclinic phase transformation due to low-temperature degradation (LTD), but this was lower compared to the airborne-particle abrasion. Other solutions include hot hydrochloric acid, hydrochloric acid and ferric selenide solution, sulfuric acid in solution with hydrogen peroxide, a mixture of nitric acid and hydrofluoric acid, hot sulfuric acid, and hot sulfuric acid mixed with hydrofluoric acid that also improved bond strength and promoted a roughened surface. The use of coupling agents, such as zirconia primers, metal primers, and silanes, have been recommended to provide chemical bonding to zirconia and improve bond strength, preferentially after surface treatment. After tribochemical silica coating, glass-ceramic coating, and deposition of silica by plasma technique on the zirconia surface, application of bifunctional silane coupling agent promotes bonding to both inorganic and organic surfaces, increasing the adhesion between resin cement and zirconia, because zirconium dioxide ceramic is not readily reactive to conventional silane agents. The silane-coupling agent contains 1% 2% of prehydrolyzed γ-methacryloxypropyltrimethoxy-silane in ethanol water solution, under
Zirconium in dentistry 333 a pH of 4 5 adjusted using acetic acid. Silanes chemically bond to the silicacoated zirconia through silanol functional groups and may react to zirconia forming Si O Zr bonds, sufficient to stabilize tetragonal-phase and copolymerize with the resin cement matrix, increasing surface energy and wettability of substrate. An important component of primers is the adhesive functional monomer phosphate MDP, which improves the bond strength, due to the formation of P O Zr covalent chemical bonds between MDP and zirconia, and is resistant to water storage. Some universal primers also include phosphoric acid and sulfide acrylate monomers that enhance chemical bonding to oxide ceramics and other prosthetic materials. These adhesive monomers have the ability to form secondary bonds as van der Waals forces and hydrogen bonds with metal oxides, interfacial forces that improve the surface wettability of zirconia. Heat treatment of metal/zirconia primer improved the initial bond strength. In addition, resin cements containing phosphoric acid group monomers are recommended to improve the bonding to zirconium dioxide ceramic, but these cements alone are not sufficient for creating a stable adhesion and primer should be applied previously. Some universal adhesives contain both silane and phosphoric monomers and seem to be a good alternative strategy for bonding to zirconia, by simplifying the practical handling. Therefore the combination of a surface treatment followed by the application of phosphatebased primer and resin-luting agents containing phosphate monomers shows promising results. 14.7 Failure and fractographic analysis of zirconia restorations Clinical studies report the higher failure of PFZ restorations than PFM restorations, mainly related to chipping or delamination of veneering porcelain and deboning of the restoration. Core fractures of zirconia-based restorations may also be found in literature but are not usually reported due to the high strength of zirconia. For multiunit FPDs, clinical failures have their origin located at the gingival area of connectors; the origin is reported to be at the cervical margin of the crowns, where the anatomical shape of tooth crown presents a curvature toward the occlusal surface and where stress concentrates during chewing. Chipping or delamination of porcelain are usually reported as material dependent, resulting from thermal incompatibilities between porcelain and zirconia, adhesion/wet problems or pores at this interface, inappropriate support provided by zirconia infrastructure for the veneering porcelain, contamination during veneering, or irregular porcelain thickness. The fracture origin may be located at the interface between the framework and porcelain or related to occlusal contact wear. For implant supported crowns, zirconia presented seven times more complications, most related to screw loosening, loss of retention, and porcelain fracture. The areas of contact between the zirconia abutment and the implant platform and/or the titanium screw are also sites of fracture origin. The difference between elastic
334 Advanced Dental Biomaterials modulus of involved materials (zirconia: 269 GPa; titanium: 110 GPa) promotes unequal stress distribution at the system, favoring fractures at the interfaces. The criteria established for clinical survival or failure varies between the clinical trials. Cohesive fractures of veneering ceramic usually do not impair function, they may be polished/repaired and restoration is not removed. For this reason, chipping may be classified as “survival.” The variation in classifying survival/failure criteria explain studies that show clinical success rates of PFZ crowns similar to PFM crowns, both tooth or implant supported, or no difference in esthetic or function of metal-ceramic or all-ceramic three-unit FPD, according to patients and dentist evaluation after 3 years. The adequate approach by the thought of the authors would be to consider failure “the restorations that require replacement”—those which failed in reestablishing esthetics and function. It is also difficult to reproduce the clinical failure of crowns in in vitro experiments. The occlusal loading is complex to simulate in laboratory experiments, usually leading to damage at the load application site, with the origin of fractures related to contact damage but not from the cervical area, as in the clinical scenario. The use of soft pistons for load application and abutments with elastic properties similar to dentin may be useful for simulating clinical failures of all-ceramic restorations. The load used in laboratory experiments is usually greater than the clinical chewing loads, also a reason for fracture origins at the load application site. Higher loads are used for accelerating the experiments since clinically failures take place after 3 5 years of service, which is not a feasible laboratory time. The study of fractured restorations started around the 1990s. It aims to identify signs of crack propagation at the fractured surface, for determination of the fracture origin. Furthermore, quantitative analysis may be performed after in vitro tests for calculation of fracture toughness. The fracture origin may help researchers to identify the causes of failure, and where the efforts of improvement must be focused on increasing clinical survival of restorations. The fractographic analysis of ceramics and glasses has a guide recommended by the National Institute for Standards and Technology, written by George Quinn, where researchers and clinicians may find information on how to proceed with the analysis of both laboratory samples and clinically fractured restorations and also the equipment that are useful/necessary for the analysis. It is available online and free of charge. All materials have an inherent population of defects. Defects in zirconia may be originated during manufacturing, machining (presintering), or during laboratory and clinical adjustments (after sintering). A defect with a critical size will determine the strength and fatigue life of the prosthetic piece, and the material will fail when the nominal strength overcome the stress field created around this defect. Laboratory failed specimens may be directly observed under the microscope for determination of the fracture origin. In the case of evaluating fractured clinical pieces, the restoration may be removed. In cases where replacement is required, or when the restoration may be repaired and keep performing esthetic and function, it is possible to make an impression of the fractured restoration/surfaces with ultralight-bodied impression material followed by an epoxy-poured replica, which can be observed under the microscope.
Zirconium in dentistry 335 The first evaluation under a stereomicroscope with lateral illumination will guide the observer to the fracture origin site. Afterward, scanning electron microscopy images will allow more detailed observation of the critical defect which led the restoration to fracture. The identification of the critical defect, together with information, such as the stress field, the load applied/tooth arc position, the history of fracture reported by the patient, cementation procedures adopted, number of cycles/ time in service, will help the identification of the most likely cause of fracture. In the case of zirconia, the image of round-shaped grains at the failure origin indicates a defect introduced in the presintering stage, while angulated grains may indicate a defect introduced after sintering, as grinding cracks for example. The observation with transillumination of frameworks prior to veneer application will help to identify failures occurred in the manufacturing or machining stages. These rough failures, and also failures introduced during veneer application, are responsible for premature clinical failures. The reproduction of clinical failures in the laboratorial environment is of high relevance since in vitro data have relevance when simulating the oral situation. The fractography is the tool that allows the clinician and the researcher to investigate the possible reason for the failure of the restoration and to verify if the laboratorial pieces failed for the same reasons. 14.8 Mechanical testing of zirconia ceramics Every new technique, composition or structural design, proposed for a material/restoration must be tested before large appliance. In the field of dental materials, not all materials launched or the alterations proposed may be tested clinically due to the long time expended in clinical researches. Thus mechanical tests are applied. These tests should be performed with a focus on investigating factors responsible for the clinical failures, such as chipping, delamination, and fractures caused by introduced defects in different stages of material production and the restoration’s manufacturing. Besides that, fatigue tests are of high relevance in clinical experimentation of ceramics, since restorations will fail after being subjected to a period of clinical service, at loads below the fracture strength of the material or prosthetic piece. Fracture occurs after the rupture of the bonds in a material, and in ceramics is determined by the largest defect present in the loaded piece. Usually bending and/or flexural tests (three- and four-point bending and biaxial flexural strength tests) are used to measure the mechanical performance, structural integrity, strength, and fatigue of zirconia ceramics, by employing bar or disk specimens. The biaxial flexural strength test is advocated to present a more clinical relevant stress scenario and failure mode than uniaxial tests, presenting delamination failures in layered zirconia. Several standards are applied for testing the properties of dental materials. The ISO 6872 describes the tests to be applied for dental ceramic materials indicated for fixed all-ceramic and metal-ceramic restorations and prostheses. Zirconia is
336 Advanced Dental Biomaterials classified as type II ceramics, which are provided in other forms than powders, pastes, or aerosols. The dimensions of samples and test devices, such as the calculation of strengths, are provided in detail by the standard. Other standards are available for specific tests, such as American Society for Testing and Materials standards, Deutsches Institut für Normung, and others. The martensitic phase transformation of zirconia grains when the surface is subjected to stimuli, such as air abrasion or grinding, increases strength and fatigue of zirconia. The mechanical stimuli trigger the transformation on the zirconia surface, and the volume expansion creates a compression layer of about 0.5 7 μm, which acts against the tensile forces generated during strength tests. This characteristic, together with the high initial mechanical strength of zirconia, will require test equipment with high strength capacity. One of the main clinical failures of zirconia is the chipping or delamination of the veneering porcelain. For testing of layered structures, the variation in elastic properties of the different materials and the thickness of the layers must be considered. The use of nonplanar specimens, such as curved disks or dome structures, may also lead to relevant failure modes like fractures with the origin located at the margin of the structures. However, the clinical geometries are far more complex, mainly in occlusal and cervical proximal tooth faces. Even though, load-to-failure tests of cemented crowns lead to very high load results and do not present a significant clinical outcome. In the clinical situation, failures occur under the application of a load below that registered in strength tests. However, for zirconia, the fatigue limit of monolithic zirconia bridges was reported to be around 1600 N even after 106 mechanical cycles in water, a load much higher than the occlusal loads registered for the posterior region (around 300 400 N). The high loads and high cyclic frequencies are applied with the aim of decreasing the time expended in laboratory tests. Fatigue is defined as the degradation of the material under mechanical, chemical, and biological stress. Defects in the bulk or surface of the ceramic present a slow growth under fatigue, until reaching a critical size, which will lead to fracture. Fatigue tests may represent the clinical failure mode, based on time in service. The sliding-contact fatigue test presented failures as porcelain chipping similar to the clinical pattern. The aim is to reproduce the clinical failure mechanics in laboratory fatigue tests, with similar origins, flaw type, and damage caused over time. A complete description of characteristics and indications of fatigue tests for ceramics can be found in ADM guidance, published in 2017. Zirconia ceramics present a specific degradation caused by water, extensively evaluated and exposed by Chevalier et al. (1999), called LTD. It is represented by the phase transformation of zirconia crystals, from tetragonal to monoclinic when in the presence of water at low temperatures. The transformation initiates in isolated surface grains and spreads through the entire surface of the material. This transformation is accompanied by a volume expansion of around 4%, which leads to nucleation of grains and enhanced surface roughness. Also, cracks are created between gains and allow the penetration of water, which brings the degradation process to the bulk of the material. Properties such as strength, toughness, and density are
Zirconium in dentistry 337 affected. Besides that, the transformation of the crystals that also occurs during heating and cooling may lead to the development of residual stresses at grain boundaries, which can lead to nucleation of microcracking under external stresses. The distribution of flaws has a great effect on the strength of ceramics: ceramics with large flaws have low strength and high scatter in the experimental data. For this reason, the Weibull statistical analysis is recommended for analysis of data generated during strength and fatigue tests. 14.9 Limitations and challenges The zirconia-based ceramics have been the target of many studies, in vitro, in vivo, or in silico (numerical method with computer simulation). Like any other material, it has many advantages, as discussed already in this chapter, but also presents some disadvantages and aspects to be improved. The initial obstacle to be overcome is to develop a material which can replace dental tissues—one single material is not able, at the moment, to replace both enamel and dentin, due to the mechanical and structural discrepancies between both tissues. In the clinical situation, zirconia presents other limitations, like the chipping/ delamination of veneering porcelain, and the difficulty in adhesion, which requires mechanically retentive abutments. The LTD was already reported in a clinical situation, but for hip prosthesis, and demonstrated in vitro. In laboratory tests, the origin of failures, the high loads, and time required for fracture hamper the simulation of the clinical situation. The high incidence of chipping and delamination of veneering porcelain of zirconia restorations in clinical studies has led the manufacturers to develop zirconia for monolithic restorations, which may be stained for a better esthetic match. Polychromatic CAD/CAM blocks are also available in the market, which imitate the color variations from cervical to incisal thirds of teeth. Translucency has also been increased in zirconia blocks, aiming to mimic the esthetic effect of glass containing ceramics. However, clinically, they are more suitable for the posterior region. The absence of the glassy phase or silica content in zirconia ceramics leads to a material with improved mechanical characteristics, but also a chemically inert material. The noninteraction of the ceramic surface hinders the adhesive procedures. Sandblasting may help adhesion, but the damage caused to the material is something to consider in the long-term clinical service of restorations. The application of very complex surface treatments, like plasma deposition, does not help in overcoming this limitation either. At the moment, zirconia restorations are somehow dependent on mechanical retention, limiting the indications of the material. The development of glass-infiltrated zirconia is a promising solution for adhesion. When the glass is added to the cementation surface, the adhesive procedures are facilitated, and the esthetic appearance and the veneer bonding benefit. Even strength was reported to be the highest in glass-infiltrated zirconia. Different
338 Advanced Dental Biomaterials techniques for glass infiltration into zirconia grains were already presented in literature, but no version is available on the market. Regarding adhesion tests, the gold standard microtensile bond strength test is hard to perform due to the very high strength and hardness of sintered zirconia, which impairs the cutting of extremely small samples. To evaluate bonding to zirconia surface, the most employed tests are the shear and microshear bond strength tests. The limitations offered by this methodology, such as unequal stress distribution, the emergence of compression and tensile stresses at the bonded interface, and cohesive failures of the substrate, may influence the result of the researches but are well accepted in literature. The simulation of fatigue of zirconia pieces in laboratory conditions is the most important approach for the prediction of the clinical performance of the material. Due to the excellent good performance of the material, very high loads and very long time are necessary to obtain a failure by fatigue. Time-consuming laboratorial tests are usually the first option to be discarded. Thus very high loads are used, a fact that may underestimate the lifetime/survival of zirconia samples. Very high frequencies (above 10 Hz) for load application may also distance the laboratory test from the clinical situation. Moreover, when high loads and/or frequencies of load application are employed, they require specific and more sophisticated equipment. Thus obtaining fatigue failures of zirconia samples is not a simple task. The high loads applied, for fatigue or even for load-to-fracture tests, to zirconia pieces may cause fractures with the origin located at the load application site, which is not found clinically. Since laboratorial tests aim to generate clinically relevant data, the origin located in a different site decreases the significance of the research. Despite the presented limitations, zirconia is still the highest strength ceramic available for dental application, allowing restorations that were only possible in metal, to be fabricated in an all-ceramic setup. The adhesive cementation does not have an established protocol and must be further investigated. The limitations regarding veneering are supposed to be solved with the use of monolithic zirconia. The laboratorial simulation of clinical services must also be improved. However, following the indications, zirconia is a beneficial and predictable material, with high clinical success rates. Further reading Abbas, S., Maleksaeedi, S., Kolos, E., Ruys, A.J., 2015. Processing and properties of zirconia-toughened alumina prepared by gel casting. Materials (Basel) 8 (7), 4344 4362. Aboushelib, M.N., Kleverlaan, C.J., Feilzer, A.J., 2007. Selective infiltration-etching technique for a strong and durable bond of resin cements to zirconia-based materials. J. Prosthet. Dent. 98 (5), 379 388. Aboushelib, M.N., Feilzer, A.J., Kleverlaan, C.J., 2010. Bonding to zirconia using a new surface treatment. J. Prosthodont. 19 (5), 340 346.
Zirconium in dentistry 339 Akın, H., Ozkurt, Z., Kırmalı, O., Kazazoglu, E., Ozdemir, A.K., 2011. The shear bond strength of resin cement to zirconia ceramic after aluminum oxide sandblasting and various laser treatments. Photomed. Laser. Surg. 29 (12), 797 802. Akpinar, Y.Z., Kepceoglu, A., Yavuz, T., Aslan, M.A., Demirtag, Z., Kılıc, H.S., et al., 2015. Effect of femtosecond laser beam angle on bond strength of zirconia-resin cement. Lasers Med. Sci. 30 (8), 2123 2128. Al Jeaidi, Z.A., Alqahtani, M.A., Awad, M.M., Rodrigues, F.P., Alrahlah, A.A., 2017. Bond strength of universal adhesives to air-abraded zirconia ceramics. J. Oral Sci. 59 (4), 565 570. Amaral, M., Valandro, L.F., Bottino, M.A., Souza, R.O., 2013. Low-temperature degradation of a Y-TZP ceramic after surface treatments. J. Biomed. Mater. Res. B Appl. Biomater. 101 (8), 1387 1392. Amaral, M., Belli, R., Cesar, P.F., Valandro, L.F., Petschelt, A., Lohbauer, U., 2014. The potential of novel primers and universal adhesives to bond to zirconia. J. Dent. 42 (1), 90 98. Amaral, M., Villefort, R.F., Melo, R.M., Pereira, G.K.R., Zhang, Y., Valandro, L.F., et al., 2017. Fatigue limit of monolithic Y-TZP three-unit-fixed dental prostheses: effect of grinding at the gingival zone of the connector. J. Mech. Behav. Biomed. Mater. 72, 159 162. Amaral, M., Weitzel, I.S.S.L., Silvestri, T., Guilardi, L.F., Pereira, G.K.R., Valandro, L.F., 2018. Effect of grinding and aging on subcritical crack growth of a Y-TZP ceramic. Braz. Oral Res. 32, e32. Antunes, M.C.F., Miranda, J.S., Carvalho, R.L.A., Carvalho, R.F., Kimpara, E.T., Assunção, E., et al., 2018. Can low-fusing glass application affect the marginal misfit and bond strength of Y-TZP crowns?. Braz. Oral Res. 32, e34. Anusavice, K.J., Shen, C., Rawls, H.R., 2013. Phillips’ Science of Dental Materials. Elsevier Health Sciences. Att, W., Grigoriadou, M., Strub, J.R., 2007. ZrO2 three-unit fixed partial dentures: comparison of failure load before and after exposure to a mastication simulator. J. Oral Rehabil. 34 (4), 282 390. Attia, A., Kern, M., 2011. Long-term resin bonding to zirconia ceramic with a new universal primer. J. Prosthet. Dent. 106 (5), 319 327. Bacchi, A., Consani, R.L.X., Mesquita, M.F., dos Santos, M.B., 2013. Stress distribution in fixed-partial prosthesis and peri-implant bone tissue with different framework materials and vertical misfit levels: a three-dimensional finite element analysis. J. Oral Sci. 55 (3), 239 244. Borba, M., de Araújo, M.D., de Lima, E., Yoshimura, H.N., Cesar, P.F., Griggs, J.A., et al., 2011. Flexural strength and failure modes of layered ceramic structures. Dent. Mater. 27 (12), 1259 1266. Bottino, M.A., Bergoli, C., Lima, E.G., Marocho, S.M., Souza, R.O., Valandro, L.F., 2014. Bonding of Y-TZP to dentin: effects of Y-TZP surface conditioning, resin cement type, and aging. Oper. Dent. 39 (3), 291 300. Campos, T.M., Ramos, N.C., Machado, J.P., Bottino, M.A., Souza, R.O., Melo, R.M., 2016. A new silica-infiltrated Y-TZP obtained by the sol-gel method. J. Dent. 48, 55 61. Cavalcanti, A.N., Foxton, R.M., Watson, T.F., Oliveira, M.T., Giannini, M., Marchi, G.M., 2009. Bond strength of resin cements to a zirconia ceramic with different surface treatments. Oper. Dent. 34 (3), 280 287. Cesar, P.F., Della Bona, A., Scherrer, S.S., Tholey, M., van Noort, R., Vichi, A., et al., 2017. ADM guidance-ceramics: fracture toughness testing and method selection. Dent. Mater. 33 (6), 575 584.
340 Advanced Dental Biomaterials Chen, L., Suh, B.I., Brown, D., Chen, X., 2012. Bonding of primed zirconia ceramics: evidence of chemical bonding and improved bond strengths. Am. J. Dent. 25 (2), 103 108. Cheng, C.W., Chien, C.H., Chen, C.J., Papaspyridakos, P., 2018. Randomized controlled clinical trial to compare posterior implant-supported modified monolithic zirconia and metal-ceramic single crowns: one-year results. J. Prosthodont. 28, 15 21. Chevalier, J., Cales, B., Drouin, J.M., 1999. Low-temperature aging of Y-TZP ceramics. J. Am. Ceram. Soc. 82 (8), 2150 2154. Chevalier, J., Deville, S., Münch, E., Jullian, R., Lair, F., 2004. Critical effect of cubic phase on aging in 3mol% yttria-stabilized zirconia ceramics for hip replacement prosthesis. Biomaterials 25 (24), 5539 5545. Cho, J.H., Kim, S.J., Shim, J.S., Lee, K.W., 2017. Effect of zirconia surface treatment using nitric acid-hydrofluoric acid on the shear bond strengths of resin cements. J. Adv. Prosthodont. 9 (2), 77 84. Coelho, P.G., Silva, N.R., Bonfante, E.A., Guess, P.C., Rekow, E.D., Thompson, V.P., 2009. Fatigue testing of two porcelain-zirconia all-ceramic crown systems. Dent. Mater. 25 (9), 1122 1127. Conrad, H.J., Seong, W.J., Pesun, I.J., 2007. Current ceramic materials and systems with clinical recommendations: a systematic review. J. Prosthet. Dent. 98 (5), 389 404. Curtis, A.R., Wright, A.J., Fleming, G.J., 2006. The influence of surface modification techniques on the performance of a Y-TZP dental ceramic. J. Dent. 34 (3), 195 206. Dal Piva, A., Contreras, L., Ribeiro, F.C., Anami, L.C., Camargo, S., Jorge, A., et al., 2018. Monolithic ceramics: effect of finishing techniques on surface properties, bacterial adhesion and cell viability. Oper. Dent. 43 (3), 315 325. Delgado-Ruı́z, R.A., Calvo-Guirado, J.L., Moreno, P., Guardia, J., Gomez-Moreno, G., MateSánchez, J.E., et al., 2011. Femtosecond laser microstructuring of zirconia dental implants. J. Biomed. Mater. Res. B Appl. Biomater. 96 (1), 91 100. Della Bona, A., Kelly, J.R., 2008. The clinical success of all-ceramic restorations. J. Am. Dent. Assoc. 139 (Suppl.), 8S 13S. Della Bona, A., Donassollo, T.A., Demarco, F.F., Barrett, A.A., Mecholsky Jr., J.J., 2007. Characterization and surface treatment effects on topography of a glass-infiltrated alumina/zirconia-reinforced ceramic. Dent. Mater. 23 (6), 769 775. Denry, I., 1996. Recent advances in ceramics for dentistry. Crit. Rev. Oral. Biol. Med. 7 (2), 134 143. Denry, I., Kelly, J.R., 2008. State of the art of zirconia for dental applications. Dent. Mater. 24 (3), 299 307. Denry, I., Kelly, J.R., 2014. Emerging ceramic-based materials for dentistry. J. Dent. Res. 93 (12), 1235 1242. Dérand, P., Dérand, T., 2000. Bond strength of luting cements to zirconium oxide ceramics. Int. J. Prosthodont. 13 (2), 131 135. Druck, C.C., Pozzobon, J.L., Callegari, G.L., Dorneles, L.S., Valandro, L.F., 2015. Adhesion to Y-TZP ceramic: study of silica nanofilm coating on the surface of Y-TZP. J. Biomed. Mater. Res. B Appl. Biomater. 103 (1), 143 150. Ernst, C.P., Cohnen, U., Stender, E., Willershausen, B., 2005. In vitro retentive strength of zirconium oxide ceramic crowns using different luting agents. J. Prosthet. Dent. 93 (6), 551 558. Everson, P., Addison, O., Palin, W.M., Burke, F.J., 2012. Improved bonding of zirconia substructures to resin using a “glaze-on” technique. J. Dent. 40 (4), 347 351. Feitosa, S.A., Patel, D., Borges, A.L., Alshehri, E.Z., Bottino, M.A., Özcan, M., et al., 2015. Effect of cleansing methods on saliva-contaminated zirconia—an evaluation of resin bond durability. Oper. Dent. 40 (2), 163 171.
Zirconium in dentistry 341 Goyatá, F., Galvão, Y., Simões, T.R., Goyatá, L.F., Arruda, J.A., Moreno, A., 2018. Effect of surface treatments with acid solutions on the surface roughness of an yttrium-tetragonal zirconia polycrystal. J. Clin. Exp. Dent. 10 (4), e367 e370. Gracis, S., Thompson, V.P., Ferencz, J.L., Silva, N.R., Bonfante, E.A., 2015. A new classification system for all-ceramic and ceramic-like restorative materials. Int. J. Prosthodont. 28 (3), 227 235. Grasel, R., Santos, M.J., Rêgo, H.C., Rippe, M.P., Valandro, L.F., 2018. Effect of resin luting systems and alumina particle air abrasion on bond strength to zirconia. Oper. Dent. 43 (3), 282 290. Griffith, A.A., 1921. The phenomena of rupture and flow in solids. Phil. Trans. R. Soc. Lond. A 221, 163 198. Han, G.J., Kim, J.H., Kim, C.K., Chung, S.N., Chun, B.H., Cho, B.H., 2013. Sequential deposition of hexamethyldisiloxane and benzene in non-thermal plasma adhesion to dental ceramic. Macromol. Res. 21 (10), 1118 1126. Helmer, J.O., Driskell, R.D., 1969. Research on bioceramics. Symposium on Use of Ceramics as Surgical Implants. Clemson University, Clemson, SC. Henriques, B., Fabris, D., Souza, J.C.M., Silva, F.S., Carvalho, Ó., Fredel, M.C., et al., 2018. Bond strength enhancement of zirconia-porcelain interfaces via Nd:YAG laser surface structuring. J. Mech. Behav. Biomed. Mater. 81, 161 167. Hirano, M., 1992. Inhibition of low-temperature degradation of tetragonal zirconia ceramics—a review. Br. Ceram. Trans. J. 91, 139 147. Inokoshi, M., De Munck, J., Minakuchi, S., Van Meerbeek, B., 2014. Meta-analysis of bonding effectiveness to zirconia ceramics. J. Dent. Res. 93 (4), 329 334. Ishii, R., Tsujimoto, A., Takamizawa, T., Tsubota, K., Suzuki, T., Shimamura, Y., et al., 2015. Influence of surface treatment of contaminated zirconia on surface free energy and resin cement bonding. Dent. Mater. J. 34 (1), 91 97. Ito, Y., Okawa, T., Fukumoto, T., Tsurumi, A., Tatsuta, M., Fujii, T., et al., 2016. Influence of atmospheric pressure low-temperature plasma treatment on the shear bond strength between zirconia and resin cement. J. Prosthodont. Res. 60 (4), 289 293. Kaimal, A., Ramdev, P., Shruthi, C.S., 2017. Evaluation of effect of zirconia surface treatment, using plasma of argon and silane, on the shear bond strength of two composite resin cements. J. Clin. Diagn. Res. 11 (8), ZC39 ZC43. Karakoca, S., Yilmaz, H., 2009. Influence of surface treatments on surface roughness, phase transformation, and biaxial flexural strength of Y-TZP ceramics. J. Biomed. Mater. Res. B Appl. Biomater. 91 (2), 930 937. Kelly, J.R., 2004. Dental ceramics: current thinking and trends. Dent. Clin. North Am. 48 (2), 513 530. Kelly, J.R., Cesar, P.F., Scherrer, S.S., Della Bona, A., van Noort, R., Tholey, M., et al., 2017. ADM guidance-ceramics: fatigue principles and testing. Dent. Mater. 33 (11), 1192 1204. Kim, J.W., Covel, N.S., Guess, P.C., Rekow, E.D., Zhang, Y., 2010. Concerns of hydrothermal degradation in CAD/CAM zirconia. J. Dent. Res. 89 (1), 91 95. Larsson, C., Wennerberg, A., 2014. The clinical success of zirconia-based crowns: a systematic review. Int. J. Prosthodont. 27 (1), 33 43. Lee, J.H., Lee, C.H., 2017. Effect of the surface treatment method using airborne-particle abrasion and hydrofluoric acid on the shear bond strength of resin cement to zirconia. Dent. J. (Basel) 5 (3), E23. Lee, M.H., Son, J.S., Kim, K.H., Kwon, T.Y., 2015. Improved resin-zirconia bonding by room temperature hydrofluoric acid etching. Materials (Basel) 8 (3), 850 866.
342 Advanced Dental Biomaterials Lekesiz, H., 2014. Reliability estimation for single-unit ceramic crown restorations. J. Dent. Res. 93 (9), 923 928. Liu, D., Tsoi, J.K., Matinlinna, J.P., Wong, H.M., 2015. Effects of some chemical surface modifications on resin zirconia adhesion. J. Mech. Behav. Biomed. Mater. 46, 23 30. Lohbauer, U., Amberger, G., Quinn, G.D., Scherrer, S.S., 2010. Fractographic analysis of a dental zirconia framework: a case study on design issues. J. Mech. Behav. Biomed. Mater. 3 (8), 623 629. Lohbauer, U., Scherrer, S.S., Della Bona, A., Tholey, M., van Noort, R., Vichi, A., et al., 2017. ADM guidance-ceramics: all-ceramic multilayer interfaces in dentistry. Dent. Mater. 33 (6), 585 598. Malkondu, O., Tinastepe, N., Kazazoglu, E., 2016. Influence of type of cement on the color and translucency of monolithic zirconia. J. Prosthet. Dent. 116 (6), 902 908. McLean, J.W., Hughes, T.H., 1965. The reinforcement of dental porcelain with ceramic oxides. Br. Dent. J. 119 (6), 251 267. Menani, L.R., Farhat, I.A., Tiossi, R., Ribeiro, R.F., Guastaldi, A.C., 2014. Effect of surface treatment on the bond strength between yttria partially stabilized zirconia ceramics and resin cement. J. Prosthet. Dent. 112 (2), 357 364. Negreiros, W.M., Ambrosano, G.M.B., Giannini, M., 2017. Effect of cleaning agent, primer application and their combination on the bond strength of a resin cement to two yttriumtetragonal zirconia polycrystal zirconia ceramics. Eur. J. Dent. 11 (1), 6 11. Nevarez-Rascon, A., Aguilar-Elguezabal, A., Orrantia, E., Bocanegra-Bernal, M.H., 2009. On the wide range of mechanical properties of ZTA and ATZ based dental ceramic composites by varying the Al2O3 and ZrO2 content. Int. J. Refract. Met. Hard Mater. 27, 962 970. Nicolaisen, M.H., Bahrami, G., Schropp, L., Isidor, F., 2016. Functional and esthetic comparison of metal-ceramic and all-ceramic posterior three-unit fixed dental prostheses. Int. J. Prosthodont. 29 (5), 473 481. Øilo, M., Arola, D., 2018. Fractographic analyses of failed one-piece zirconia implant restorations. Dent. Mater. 34, 922 931. Øilo, M., Kvam, K., Tibballs, J.E., Gjerdet, N.R., 2013. Clinically relevant fracture testing of all-ceramic crowns. Dent. Mater. 29 (8), 815 823. Øilo, M., Hardang, A.D., Ulsund, A.H., Gjerdet, N.R., 2014. Fractographic features of glassceramic and zirconia-based dental restorations fractured during clinical function. Eur. J. Oral Sci. 122 (3), 238 244. Özcan, M., Bernasconi, M., 2015. Adhesion to zirconia used for dental restorations: a systematic review and meta-analysis. J. Adhes. Dent. 17 (1), 7 26. Özcan, M., Cura, C., Valandro, L.F., 2011. Early bond strength of two resin cements to YTZP ceramic using MPS or MPS/4-META silanes. Odontology 99 (1), 62 67. Paes, P.N.G., Bastian, F.L., Jardim, P.M., 2017. The influence of Y-TZP surface treatment on topography and ceramic/resin cement interfacial fracture toughness. Dent. Mater. 33 (9), 976 989. Pandoleon, P., Kontonasaki, E., Kantiranis, N., Pliatsikas, N., Patsalas, P., Papadopoulou, L., et al., 2017. Aging of 3Y-TZP dental zirconia and yttrium depletion. Dent. Mater. 33 (11), e385 e392. Park, C., Yoo, S.H., Park, S.W., Yun, K.D., Ji, M.K., Shin, J.H., et al., 2017. The effect of plasma on shear bond strength between resin cement and colored zirconia. J. Adv. Prosthodont. 9 (2), 118 123. Peampring, C., Aksornmuang, J., Sanohkan, S., 2017. In vitro fracture resistance of composite-resin-veneered zirconia crowns. J. Conserv. Dent. 20 (4), 225 229.
Zirconium in dentistry 343 Pereira, G.K., Silvestri, T., Camargo, R., Rippe, M.P., Amaral, M., Kleverlaan, C.J., et al., 2016. Mechanical behavior of a Y-TZP ceramic for monolithic restorations: effect of grinding and low-temperature aging. Mater. Sci. Eng. C Mater. Biol. Appl. 63, 70 77. Piconi, C., Maccauro, G., 1999. Zirconia as a ceramic biomaterial. Biomaterials 20 (1), 1 25. Pitta, J., Branco, T.C., Portugal, J., 2018. Effect of saliva contamination and artificial aging on different primer/cement systems bonded to zirconia. J. Prosthet. Dent. 119 (5), 833 839. Qasim, T., Ford, C., Bush, M.B., Hu, X.Z., Malament, K.A., Lawn, B.R., 2007. Margin failures in brittle dome structures: relevance to failure of dental crowns. J. Biomed. Mater. Res. B Appl. Biomater. 80 (1), 78 85. Qualtrough, A.J., Piddock, V., 1997. Ceramics update. J. Dent. 25 (2), 91 95. Quinn, G.D., 2016. NIST recommended practice guide: fractography of ceramics and glasses. Special Publication 960-16e2. National Institute of Standards and Technology, Washington, DC. Ramos-Tonello, C.M., Trevizo, B.F., Rodrigues, R.F., Magalhães, A.P.R., Furuse, A.Y., Lisboa-Filho, P.N., et al., 2017. Pre-sintered Y-TZP sandblasting: effect on surface roughness, phase transformation, and Y-TZP/veneer bond strength. J. Appl. Oral Sci. 25 (6), 666 673. Ritter, A.V., 2017. Sturdevant’s Art & Science of Operative Dentistry. Elsevier Health Sciences. Sailer, I., Balmer, M., Hüsler, J., Hämmerle, C.H.F., Känel, S., Thoma, D.S., 2017. Comparison of fixed dental prostheses with zirconia and metal frameworks: five-year results of a randomized controlled clinical trial. Int. J. Prosthodont. 30 (5), 426 428. Scherrer, S.S., Cattani-Lorente, M., Vittecoq, E., de Mestral, F., Griggs, J.A., Wiskott, H.W., 2011. Fatigue behavior in water of Y-TZP zirconia ceramics after abrasion with 30 μm silica-coated alumina particles. Dent. Mater. 27 (2), e28 e42. Scherrer, S.S., Lohbauer, U., Della Bona, A., Vichi, A., Tholey, M.J., Kelly, J.R., et al., 2017. ADM guidance-ceramics: guidance to the use of fractography in failure analysis of brittle materials. Dent. Mater. 33 (6), 599 620. Silva, L.H., Costa, A.K., Queiroz, J.R., Bottino, M.A., Valandro, L.F., 2012. Ceramic primer heat-treatment effect on resin cement/Y-TZP bond strength. Oper. Dent. 37 (6), 634 640. Silva, L.H.D., Miranda, R.B.D.P., Favero, S.S., Lohbauer, U., Cesar, P.F., 2017. Dental ceramics: a review of new materials and processing methods. Braz. Oral Res. 31 (Suppl.), e58. Skovgaard, M., Almdal, K., van Lelieveld, A., 2011. Stabilization of metastable tetragonal zirconia nanocrystallites by surface modification. J. Mater. Sci. 46 (6), 1824 1829. Available from: https://doi.org/10.1007/s10853-010-5007-1. Smielak, B., Klimek, L., 2015. Effect of hydrofluoric acid concentration and etching duration on select surface roughness parameters for zirconia. J. Prosthet. Dent. 113 (6), 596 602. Sriamporn, T., Thamrongananskul, N., Busabok, C., Poolthong, S., Uo, M., Tagami, J., 2014. Dental zirconia can be etched by hydrofluoric acid. Dent. Mater. J. 33 (1), 79 85. Tabatabaian, F., 2018. Color in zirconia-based restorations and related factors: a literature review. J. Prosthodont. 27 (2), 201 211. Tholey, M.J., Swain, M.V., Thiel, N., 2011. Thermal gradients and residual stresses in veneered Y-TZP frameworks. Dent. Mater. 27 (11), 1102 1110. Thompson, J.Y., Stoner, B.R., Piascik, J.R., Smith, R., 2011. Adhesion/cementation to zirconia and other non-silicate ceramics: where are we now?. Dent. Mater. 27 (1), 71 82.
344 Advanced Dental Biomaterials Tsuo, Y., Yoshida, K., Atsuta, M., 2006. Effects of alumina-blasting and adhesive primers on bonding between resin luting agent and zirconia ceramics. Dent. Mater. J. 25 (4), 669 674. Tzanakakis, E.G., Tzoutzas, I.G., Koidis, P.T., 2016. Is there a potential for durable adhesion to zirconia restorations? A systematic review. J. Prosthet. Dent. 115 (1), 9 19. Ural, Ç., Külünk, T., Külünk, Ş., Kurt, M., 2010. The effect of laser treatment on bonding between zirconia ceramic surface and resin cement. Acta Odontol. Scand. 68 (6), 354 359. Vagkopoulou, T., Koutayas, S.O., Koidis, P., Strub, J.R., 2009. Zirconia in dentistry: Part 1. Discovering the nature of an upcoming bioceramic. Eur. J. Esthet. Dent. 4 (2), 130 151. Vanderlei, A., Bottino, M.A., Valandro, L.F., 2014. Evaluation of resin bond strength to yttria-stabilized tetragonal zirconia and framework marginal fit: comparison of different surface conditionings. Oper. Dent. 39 (1), 50 63. Vechiato-Filho, A.J., Matos, A.O., Landers, R., Goiato, M.C., Rangel, E.C., De Souza, G.M., et al., 2017. Surface analysis and shear bond strength of zirconia on resin cements after non-thermal plasma treatment and/or primer application for metallic alloys. Mater. Sci. Eng. C Mater. Biol. Appl. 72, 284 292. Vicente, M., Gomes, A.L., Montero, J., Rosel, E., Seoane, V., Albaladejo, A., 2016. Influence of cyclic loading on the adhesive effectiveness of resin-zirconia interface after femtosecond laser irradiation and conventional surface treatments. Lasers Surg. Med. 48 (1), 36 44. Wandscher, V.F., Prochnow, C., Rippe, M.P., Dorneles, L.S., Callegari, G.L., Baldissara, P., et al., 2017. Retentive strength of Y-TZP crowns: comparison of different silica coating methods on the intaglio surfaces. Oper. Dent. 42 (5), E121 E133. Weber, K.K., 2018. Biomaterials, Zirconia. NCBI Bookshelf. Wendler, M., Belli, R., Petschelt, A., Mevec, D., Harrer, W., Lube, T., et al., 2017. Chairside CAD/CAM materials. Part 2: Flexural strength testing. Dent. Mater. 33 (1), 99 109. Xie, H., Chen, C., Dai, W., Chen, G., Zhang, F., 2013. In vitro short-term bonding performance of zirconia treated with hot acid etching and primer conditioning etching and primer conditioning. Dent. Mater. J. 32 (6), 928 938. Xie, H., Li, Q., Zhang, F., Lu, Y., Tay, F.R., Qian, M., et al., 2016. Comparison of resin bonding improvements to zirconia between one-bottle universal adhesives and tribochemical silica coating, which is better?. Dent. Mater. 32 (3), 403 411. Xie, H., Cheng, Y., Chen, Y., Qian, M., Xia, Y., Chen, C., 2017. Improvement in the bonding of Y-TZP by room-temperature ultrasonic HF etching. J. Adhes. Dent. 19 (5), 425 433. Yamamoto, L.T., Rodrigues, V.A., Dornelles, L.S., Bottino, M.A., Valandro, L.F., Melo, R. M., 2016. Low-fusing porcelain glaze application on 3Y-TZP surfaces can enhance zirconia-porcelain adhesion. Braz. Dent. J. 27 (5), 543 547. Yang, B., Wolfart, S., Scharnberg, M., Ludwig, K., Adelung, R., Kern, M., 2007. Influence of contamination on zirconia ceramic bonding. J. Dent. Res. 86 (8), 749 753. Yang, B., Lange-Jansen, H.C., Scharnberg, M., Wolfart, S., Ludwig, K., Adelung, R., et al., 2008. Influence of saliva contamination on zirconia ceramic bonding. Dent. Mater. 24 (4), 508 513. Yang, B., Barloi, A., Kern, M., 2010. Influence of air-abrasion on zirconia ceramic bonding using an adhesive composite resin. Dent. Mater. 26 (1), 44 50. Yoshida, K., 2018. Influence of cleaning methods on resin bonding to saliva-contaminated zirconia. J. Esthet. Restor. Dent. 30, 259 264.
Zirconium in dentistry 345 Zhang, Y., Kim, J.W., 2010. Graded zirconia glass for resistance to veneer fracture. J. Dent. Res. 89 (10), 1057 1062. Zhang, Y., Pajares, A., Lawn, B.R., 2004. Fatigue and damage tolerance of Y-TZP ceramics in layered biomechanical systems. J. Biomed. Mater. Res. B Appl. Biomater. 71 (1), 166 171. Zhao, L., Jian, Y.T., Wang, X.D., Zhao, K., 2016. Bond strength of primer/cement systems to zirconia subjected to artificial aging. J. Prosthet. Dent. 116 (5), 790 796.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 15 Zeeshan Sheikh1,2,3, Nader Hamdan4, Mohamed-Nur Abdallah1, Michael Glogauer1,5 and Marc Grynpas2,3,6 1 Faculty of Dentistry, University of Toronto, Toronto, ON, Canada, 2Lunenfeld-Tanenbaum Research Institute, Mount Sinai Hospital, Toronto, ON, Canada, 3Department of Laboratory Medicine and Pathobiology, University of Toronto, Toronto, ON, Canada, 4 Department of Dental Clinical Sciences, Faculty of Dentistry, Dalhousie University, Halifax, NS, Canada, 5Institute of Biomaterials and Biomedical Engineering, University of Toronto, Toronto, ON, Canada, 6Princess Margaret Cancer Centre, Department of Dental Oncology and Maxillofacial Prosthetics, Toronto, ON, Canada Chapter Outline 15.1 Introduction 347 15.2 Rationale behind use of bone replacement graft materials 348 15.3 Natural tissues and synthetic biomaterials used for bone grafting 15.3.1 15.3.2 15.3.3 15.3.4 Autografts Allografts Xenografts Alloplasts 350 352 353 354 355 15.4 Biocompatibility of bone replacement graft materials and their degradation products 363 15.5 Biodegradation of implanted graft materials and bone formation 363 15.6 Future of bone tissue graft materials 365 References 366 15.1 Introduction Bone replacement is a complex phenomenon and a challenging task to accomplish. Bone grafts are the second most common transplanted tissue after blood and more than 2.2 million grafting procedures take place annually worldwide in order to repair bone defects in orthopedics, dentistry, and neurosurgery (Lewandrowski et al., 2000). The treatments of posttraumatic skeletal Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00015-3 Copyright © 2019 Elsevier Ltd. All rights reserved.
348 Advanced Dental Biomaterials complications, such as delayed unions, nonunions, malunions along with spinal fusions, filling defects following removal of bone tumors, and several congenital diseases, are challenges that can be met successfully by grafting procedures (Aludden et al., 2017). An understanding of bone biology and its constitutional elements is imperative in comprehending the challenges that must be overcome when successful bone grafting is the desired goal. Bone formation starts with the secretion of collagen molecules and ground substance by osteoblasts. The collagen molecules polymerize in order to create collagen fibers. Calcium salts then precipitate in the ground substance to form osteoid along the collagen fibers. The osteoblasts become trapped in the osteoid, which are referred to as osteocytes (Hall, 2015). Mature compact bone is composed of approximately 30% organic matrix and 70% calcium salts. Of note, 90% 95% of the organic matrix is collagen fibers, and the remainder is the gelatinous medium named as ground substance, which comprises chondroitin sulfate and hyaluronic acid (Hall, 2015). The collagen fibers are oriented along the lines of tensional force. The main crystalline salt, composed of calcium and phosphate, is hydroxyapatite (HA). Compact bone has HA crystals that are adjacent to and bound to the collagen fibers. The collagen fibers provide tensile strength, and the HA crystals are responsible for imparting compressional strength (Hall, 2015). Hence this natural composite we commonly refer to as bone comes into existence, comprised organic and inorganic components. Bone formation is an elaborate process and it is beyond the scope of this chapter to cover all aspects in detail. 15.2 Rationale behind use of bone replacement graft materials Bone replacement grafts serve as a structural framework for clot development, maturation, and remodeling that ultimately leads to bone formation in osseous defects (Sheikh, 2013). After implantation of the graft material, the hard and soft tissue host bed surrounding the graft must be viable and have adequate blood supply. It is rare for grafts to survive when placed within nonvital host bone. In the weeks following the implantation, new blood vessels, interstitial cells and materials, and woven bone (by new osteoblasts) are produced by the host bed. All of these embed the graft material to create the graft-woven bone complex. The graft material and the woven bone are fused together by cement lines to achieve the desired mechanical strength. Bone formation in grafting procedures involves one or more of the following biologic mechanisms: 1. Osteogenesis: The formation of new bone by osteoblasts derived from the graft material itself. 2. Osteoinduction: The ability of a material to induce the formation of osteoblasts from the surrounding tissue at the graft host site, resulting in bone growth.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 349 3. Osteoconduction: The ability of a material to support the growth of bone over a surface. Osteogenesis is achieved with autogenous grafts and the presence of cellular elements or progenitor cells of the autogenous grafts is thought to be responsible for this to occur (Schallhorn et al., 1970). As explained later in this chapter, other types of bone replacement grafts do not provide any cellular elements. The best case scenario for these nonautogenous grafts would be osteoinduction, which would stimulate phenotypic conversion of progenitor cells within the healing wound to those that can form osseous tissue. Osteoconductive materials provide a scaffold to allow bone ingrowth and deposition. Most bone replacement grafts are osteoconductive, relatively inert filling materials, and integrate with new bone without providing much histological evidence of extensive regeneration (Sheikh et al., 2017a). Although osteointegration, the ability to chemically bind to surrounding bone is not directly responsible for bone formation, is a desirable property as it aids in the incorporation of the graft at the host site. The most important aspect for the success of a bone graft is that it can be completely resorbed and remodeled. This allows the graft to not interfere with physiologic bone adaptation. The presence and recruitment of osteoblast precursors and growth factors at sites of augmentation are essential for bone regeneration to occur. Some graft materials (cancellous autogenous grafts) and the recipient bed can provide the osteoblast precursors required (Finkemeier, 2002), whereas the growth factors come from the vasculature and recipient bed. Active bone resorption and formation throughout the graft dominates the early phase of bone regeneration at grafted sites (Sheikh et al., 2015a,b). The latter phase is mainly known to be characterized by the osteoconductive processes (Urist, 1980). Osteoconduction is a function of a bone graft substrate providing a three-dimensional (3D) scaffold area promoting ingrowth of host capillaries and osteoprogenitor cells (Cornell and Lane, 1998). Biomaterials that imitate natural bone chemistry and structure closely are considered ideal for cellular osteogenic differentiation. Graft macroporosity and pore interconnection have a major impact on osteoinduction potential as higher levels of porosity, appropriate pore-shape, and sufficient interconnectivity are essential for ingrowth of blood vessels and bone matrix deposition (Mastrogiacomo et al., 2006). During the initial first few weeks, new bone is synthesized by mature osteoblasts that are differentiated from osteoblast precursors under the influence of osteoinductors. The growth factors involved in formation of new bone act directly on osteoblast and fibroblast proliferation, mesenchymal cell differentiation, extracellular matrix deposition, and vascular proliferation (Boyne, 1988). Early stages of induction are influenced by the fibroblast growth factor and platelet-derived growth factor (PDGF) by stimulating fibroblast and osteoblast proliferation. On the other hand, bone morphogenetic proteins (BMPs) affect later stages of osteoinduction such as vascular proliferation and mesenchymal cell differentiation. Whereas transforming growth factor-beta does not affect mesenchymal cell differentiation but acts on cellular proliferation, matrix deposition, and vascularization (Sheikh et al., 2015e).
350 15.3 Advanced Dental Biomaterials Natural tissues and synthetic biomaterials used for bone grafting There are various graft options available and used for alveolar bone grafting and divided into natural transplants (autografts, allografts, and xenografts) and synthetic materials (alloplasts) (Tables 15.1 and 15.2) (Sheikh et al., 2015e). These graft materials are used because they possess either osteogenic, osteoinductive, and/or osteoconductive properties (Sheikh, 2013). These grafts should ideally be biocompatible, easily molded and/or carved, integrate well with the native bone, and have adequate mechanical properties with an ideal replacement rate, and be predictable with a good level of patient acceptance (Sheikh et al., 2015d). Hard tissue substitute graft materials that have the ability to be resorbed undergo a replacement process during which they are partially or completely resorbed by macrophages/osteoclasts before native bone is deposited by osteoblasts Table 15.1 Available tissue and biomaterial bone replacement graft options. Bone replacement graft materials Human bone graft tissues 1. Autografts (cancellous and/or cortical) a. Extraoral b. Intraoral 2. Allografts (cancellous and/or cortical) c. Fresh or frozen bone d. FDBA e. DFDBA Nonhuman source materials 1. Xenografts a. Bovine hydroxyapatite b. Porcine bone c. Equine bone d. Coralline calcium carbonate Synthetic materials (alloplasts) 1. Bioactive glasses 2. Calcium phosphates a. Hydroxyapatite b. Tricalcium phosphate c. Other calcium phosphates (brushite, monetite, CPP) 3. Calcium sulfate CPP, Calcium polyphosphates; DFDBA, demineralized freeze-dried bone allograft; FDBA, freeze-dried bone allograft.
Table 15.2 Examples of common commercially available bone grafts for dental and maxillofacial applications. Category Brand name Company Generic name/composition Source Allograft Accell DBX Putty DynaBlast Citagenix DENTSPLY Keystone Dental Human bone Human bone Human bone Grafton MTF— DFDBA Puros Raptos BioHorizons Musculoskeletal Transplant Foundation Musculoskeletal Transplant Foundation Bacterin International Bacterin International Zimmer Biomet Citagenix Demineralized bone matrix Demineralized bone matrix Demineralized and mineralized bone matrix Demineralized bone matrix Demineralized freeze-dried bone Biocoral Bio-Oss Inoteb Geistlich Cerabone Botiss Dental Endobon Zimmer Biomet Gen-Os Interpore-200 Tecnoss Dental Interpore International Osteomed Dentsply Dentsply MTF— FDBA OsteoSelect OsteoSponge Xenograft Xenograft/ synthetic Alloplast Osseograft Osteograf/N PepGen P-15 Ceros Cerasorb IngeniOs β-TCP Macrobone Vitoss Guidor easygraft Biogran PerioGlass Capset Osteoset Calciresorb 35 Ceraform Eurobone Cerapatite IngeniOs HA TransOssatite Thommen Medical Curasan Zimmer Dental Euroteknika Groupe Stryker Sunstar Atek Medical Group NovaBone Lifecore Biomedicals Wright Medical Ceraver Teknimed Kasios Ceraver Zimmer Dental Transysteme JMT implant Human bone Human bone Freeze-dried bone Human bone Demineralized bone matrix Human bone Demineralized bone matrix Human bone Mineralized bone Mineralized/demineralized bone allograft Corraline calcium carbonate Deproteinized bovine bone mineral Decalcified freeze-dried bovine bone Deproteinized bovine bone mineral Anorganic porcine bone mineral Porous HA Human bone Human bone Demineralized bone matrix Anorganic bovine bone mineral Anorganic bovine bone mineral with a synthetic biomimetic of the 15 amino acid sequence of Type-I collagen β-TCP β-TCP β-TCP Bovine bone Bovine bone Bovine bone/ tissue engineering β-TCP β-TCP In situ hardening β-TCP granules coated with PLGA Bioactive glass Bioactive glass CS CS DCP (65% HA/35% β-TCP) Synthetic Synthetic Synthetic DCP (65% HA/35% β-TCP) DCP/dihydrate (Brushite)— DCPD HA HA HA Synthetic Synthetic marine corals Bovine bone Bovine bone Bovine bone Porcine bone Marine corals Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic Synthetic β-TCP, Beta-tricalcium phosphate; CS, calcium sulfate; DCP, dicalcium phosphate; DFDBA, demineralized freeze-dried bone allograft; FDBA, freeze-dried bone allograft; HA, hydroxyapatite; PLGA, poly(lactic-co-glycolic acid).
352 Advanced Dental Biomaterials (Sheikh et al., 2015a,b). This section discusses the various graft tissues and biomaterial alternatives used for alveolar bone grafting and periodontal defect fill applications. 15.3.1 Autografts Autografts are harvested from a donor site in the same individual and transplanted to another site. Autografts are a source of most osteogenic organic material for grafting, however, donor site morbidity and limited graft volume that can be obtained are disadvantages (Table 15.3) (McAllister and Haghighat, 2007; Younger and Chapman, 1989). Autografts used for dental grafting applications may be of extraoral or intraoral origin. Intraoral autograft harvest sites are the spina nasalis, the tuberosity and crista zygomatico-alveolaris from the maxilla, the ramus, retromolar region and the symphysis region in the mandible, as well as bony exostoses and bone harvested from different sites utilizing bone scrapers (Draenert et al., 2014). Mandibular autografts are commonly used as bone chips, blocks, and milled particles (Misch, 1997; Simion et al., 2001). Autografts obtained from extraoral sites such as the iliac crest provide osteoinductive, osteoconductive, and osteogenic potential (Cypher and Grossman, 1996). The calvaria is another extraoral site that can be used to potentially obtain bone tissue for surgical applications (Simion and Fontana, 2004; Jackson et al., 1986). However, there is less morbidity associated with intraoral donor sites and that is the reason they are preferred (Misch, 1997). The common extraoral harvest site that provides large amounts of autologous cortical-cancellous bone is the iliac crest (Nkenke et al., 2004). Cortical autografts have high initial strength that after about 6 months of implantation is about 50% weaker than the physiologically normal bone tissue (Wilk, 2004). Conversely, cancellous bone autografts are initially weaker because of their porous structure and gain strength over time (Sheikh, 2013). The cancellous autografts revascularize earlier than the cortical grafts around the fifth day after implantation due to their Table 15.3 Advantages and disadvantages of using autologous (autografts) bone graft. Advantages Disadvantages Biocompatible Additional surgical phase to obtain the autograft Increased operative time and cost Donor site morbidity, increased postsurgical pain Increased risk fracture Dead bone is at a potential for infection Variability in quality of harvested bone tissue Limitation of the amount of bone tissue that can be harvested Osteoinductive Osteoinductive High osteogenic potential Sufficient mechanical properties to provide adequate structural support Good and easier integration into the transplant host site (no immune response) Availability in both cortical and cancellous forms for use
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 353 spongy architecture (Sheikh, 2013). Vertical and horizontal alveolar ridge augmentation using particulate autografts with GBR has been shown to be successful for placing dental implants (Rocchietta et al., 2015; Merli et al., 2010). However, block grafts outperform particulate grafts with regards to revascularization, bone-toimplant contact, and bone remodeling (Rocchietta et al., 2015). 15.3.2 Allografts Graft tissues obtained from genetically nonidentical members of the same species are known as allografts. These grafts are available in larger quantities for use and do not have the usual drawbacks of autografts (Table 15.3). Allografts (cortical and cancellous) of various particle size ranges are used routinely for bone augmentation procedures with minimal risk of disease transmission (Table 15.4) (Block and Degen, 2004; Araujo et al., 2013; Sterio et al., 2013). Allografts are available for periodontal applications as cortical wedges, cortical chips, cortical granules, and cancellous powders prepared as frozen, freeze-dried, mineralized, and demineralized bone (Al Ruhaimi, 2001). 15.3.2.1 Fresh or frozen allografts Fresh and/or frozen cancellous bone and marrow tissues demonstrate the highest osteoconductive and osteoinductive potential among all allografts (Dias et al., 2014; Macedo et al., 2012). Atrophic maxillary ridges when grafted with human block grafts of tibia and fresh-frozen chips show features representative of mature and compact osseous tissue surrounded by marrow spaces (Contar et al., 2009, 2011). However, due to the risk of disease transmission, use of fresh or frozen iliac allografts is now obsolete. Table 15.4 Comparison between autografts and allografts. Bone graft Strength Osteogenesis Osteoconduction Osteoinduction 3 0 2 3 2 3 2 3 3 1 1 1 0 0 0 0 0 0 2 2 1 1 0 0 Autografts Cortical Cancellous Allografts Cortical Frozen Freeze-dried Cancellous Frozen Freeze-dried Scores: 0 (none), 1 (moderate), 2 (good), and 3 (excellent).
354 Advanced Dental Biomaterials 15.3.2.2 Mineralized freeze-dried bone allografts The freeze-drying to process these grafts for use distorts the 3D presentation of the human leukocyte antigens on the surface of graft particles that affects the immune recognition (Quattlebaum et al., 1988; Friedlaender et al., 1976). Freeze-dried bone allograft (FDBA) is known to be osteoconductive and can be combined with autografts to enhance the osteogenic potential (Committee on Research, 2001; Mellonig, 2000). These graft tissues are mineralized and used for the treatment of periodontal defects (Mellonig, 1991; Kukreja et al., 2014; Blaggana et al., 2014; Markou et al., 2009). Cortical FDBA demonstrate greater osteoinductive potential due to the growth factors stored in the matrix (Sunitha Raja and Naidu, 2008). FDBA used in combination with resorbable barrier membranes have been used as a replacement for autograft blocks for ridge augmentation (Lyford et al., 2003). The use of FDBA blocks for alveolar ridge grafting has shown the presence of vital bone with a lamellar organization (Jacotti et al., 2012; Wallace and Gellin, 2010). 15.3.2.3 Demineralized freeze-dried bone allogeneic grafts These allografts that have been demineralized are used alone or in combination with FDBA and autografts. Demineralized freeze-dried bone allograft (DFDBA) undergoes resorption at a quick rate (Russell et al., 1997; Hopp et al., 1989) and often have osteoinductive potential due to the BMPs and growth factors present in the graft matrix (Mellonig et al., 1981). DFDBA has been shown to produce a lesser amount of vital new bone in comparison to autografts (Scarano et al., 2006). DFDBA acquired from younger cadavers has higher osteogenic potential in comparison with grafts from older individuals, resulting in variation in BMP levels in different batches of DFDBA (Dodson et al., 1996; Jergesen et al., 1991). 15.3.3 Xenografts Xenografts are graft tissues obtained from nonhuman species, that is, animals, and are usually osteoconductive with limited resorptive potential (Thaller et al., 1993; McAllister et al., 1999). The xenograft most commonly used in periodontal regeneration procedures is the deproteinized bovine bone mineral, commercially known as Bio-Oss, which is a commercially available bone of bovine origin processed to yield natural bone mineral without the organic elements (Liu et al., 2016). After heat and chemical treatments, the inorganic phase of bovine bone consists mainly of HA that retains the porous architecture (Jarcho, 1981). Bovine-derived bone graft particles and blocks have been used for alveolar ridge augmentation procedures and intrabony defect filling (Zitzmann et al., 1997; Yildirim et al., 2000). Bio-Oss remains the most researched xenogeneic bone grafting material. Several research papers were published on the use of Bio-Oss in different surgical scenarios (Proussaefs et al., 2003; Valentini and Abensur, 1997; Wallace et al., 2005; Proussaefs et al., 2002; Berglundh and Lindhe, 1997). Of particular interest was the use of Bio-Oss as a graft material during direct subantral augmentation (sinus lift)
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 355 procedures where dental implants placed in Bio-Oss grafts had survival rates at least similar if not better than autogenous grafts (Wallace and Froum, 2003). However, although bovine-derived bone block grafts have high osteoconductive potential, these grafts are inherently brittle and lack toughness. This makes them prone to failure during the screw fixation procedures and/or after implantation (Yildirim et al., 2000; Felice et al., 2009). In addition to bovine-derived bone mineral, bone mineral can be obtained from other animal sources, such as equine or porcine sources. Porcine bone graft tissue is a porous inorganic bone graft material consisting predominantly of calcium phosphate. These are supplied in granular form with a particle size of 0.25 1 and 1 2 mm (Gen-Os) and are produced by removal of the organic components from porcine bone (Nannmark and Sennerby, 2008; Pearce et al., 2007). The anorganic bone mineral matrix is biocompatible, having an interconnecting macro- and microscopic porous structure that supports the formation and ingrowth of new bone at the implantation site (Nannmark and Sennerby, 2008). A clinical study in humans in which porcine-derived graft was investigated for implant site development showed it reduced the hard tissue resorption after tooth extraction (Festa et al., 2013). The porous microstructure of marine coral has also been used as a template to fabricate porous coralline HA materials such as interpore-200 (White and Shors, 1986). These materials are fabricated by coral being subjected to high temperature under pressurized treatment in the presence of aqueous phosphate solutions (Roy and Linnehan, 1974). This converts the coral to calcium HA, while conserving the highly organized, permeable, and interconnecting pore structure (Roy and Linnehan, 1974; White and Shors, 1986). These graft materials have an average pore diameter of 200 μm, and consist of about 60% porosity/void spaces (Weber and White, 1973; Weber et al., 1971). Some other xenograft materials that are being researched include chitosan, red algae, and gusuibu (Wong and Rabie, 2006; Cho et al., 2005). Chitosan is a product of the exoskeleton of crustaceans, and has shown the ability to stimulate mesenchymal stem cell differentiation into osteoblasts. When chitosan is combined with HA, the osteoconductivity is markedly increased (Cho et al., 2005). Red algae are one of the oldest groups of eukaryotic algae, and also one of the largest, with about 5000 6000 species of mostly multicellular, marine algae, including many notable seaweeds. Red algae can be chemically converted to HA, which then is utilized for grafting bone defects (Ewers, 2005). Gusuibu is a rhizome of perennial pteridophyte Drynaria fortunei. This herb has been shown to have good osteoconductive properties and an increase in alkaline phosphatase expression that has a positive effect of promoting calcification (Wong and Rabie, 2006) (Fig. 15.1). 15.3.4 Alloplasts Alloplastic synthetic biomaterials were developed to overcome the disadvantages of autografts and are fabricated in various forms with varying physicochemical properties and can be both degradable and nondegradable (Sheikh et al., 2015a,c,d,e, 2017a; AlGhamdi et al., 2010; Shetty and Han, 1991). Alloplasts are usually
356 Advanced Dental Biomaterials Figure 15.1 Examples of different bone grafting materials used in different clinical scenarios as per their indication. (A) Block autograft cut from chin. (B) Block autograft being stabilized using fixation screws in posterior mandible. (C) Allograft (FDBA) being placed for socket preservation covered with absorbable collagen membrane for placing dental implant in future. (D) Xenograft (Bio-Oss) used to augment bone around dental implants. (E and H) Allograft (FDBA) and xenograft (Bio-Oss) used together for sandwich technique to augment bone in the mandibular anterior region for placing dental implants. (F) An allograft (FDBA) used to augment anterior maxilla after ridge splitting for future implant placement. (G) Xenograft (Bio-Oss) used for performing sinus lifting. (I) Allograft (FDBA) covered with absorbable collagen membrane used for site preparation of partially edentulous maxillary ridge. (J) Xenograft used for ridge augmentation as part of pontic site development for three-unit fixed partial denture. (K) Alloplast (FDBA and Enamel Matrix Derivative/ EMD) for GTR procedure. FDBA, Freeze-dried bone allograft; GTR, guided tissue regeneration. osteoconductive without any osteoinductive or osteogenic potential on their own and have been used extensively for periodontal regeneration (Shetty and Han, 1991). The most routinely used alloplastic materials are HA, tricalcium phosphates (TCP), and bioactive glasses. Calcium phosphate biomaterials are of great interest to be used as bone replacement graft materials in periodontal regeneration as they have a similar composition to bone mineral, are osteoconductive, form bone apatite-like material or carbonated HA, and form a very strong bone calcium phosphate biomaterial interface (Sheikh et al., 2015a,d). Alloplastic bone grafting materials are sought after because they provide an abundant amount without the problems associated with autografts (Hench, 1998). These are fabricated in various forms and with varying physicochemical properties and can be both resorbable and nonresorbable (Sheikh et al., 2015c,d,e; AlGhamdi et al., 2010; Shetty and Han, 1991). Alloplastic materials are usually osteoconductive without having any osteogenic and osteoinductive potential and have been used successfully in periodontal reconstructive applications (Shetty and Han, 1991).
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 357 15.3.4.1 Polymers Polymers can be classified based on origin: natural and synthetic. Natural polymers that have been utilized in the fabrication of bone grafts include polysaccharides (e.g., alginate, argose, chitosan, and hyaluronic acid) and polypeptides (e.g., gelatin and collagen). Natural polymers possess weak mechanical properties and variable or negligible rates of degradation, hence their use is limited as bone grafting materials (Asghari et al., 2017). However, natural polymers play an important role in composite grafts by serving as polymeric shell capsules, which incorporate particles of allografts. Synthetic polymers (e.g., polyglycolic acid, polylactic acid, polyorthoester, and polyanhydride) provide a platform for regulating and controlling the biomechanical properties of scaffolds and they also serve as drug delivery carriers in tissue engineering applications (Sokolsky-Papkov et al., 2007). HTR Synthetic Bone (Bioplant, Norwalk, CT) is a biocompatible microporous composite made up of polymethylmethacrylate (PMMA), polyhydroxylethylmethacrylate, and calcium hydroxide. The acronym HTR stands for “Hard Tissue Replacement” and acceptable clinical results have been achieved in treatment of intrabony and furcation defects with its use (Yukna, 1990, 1994). Histologically new bone growth has been observed on the particles of this material (Froum, 1996; Yukna and Greer, 1992; Stahl et al., 1990). This material has hydrophilicity that enhances clotting, and the negative particle charges on the surface allows adherence to bone. Clinically acceptable defect fill and resolution can be achieved and support the use of HTR as a biocompatible alloplastic bone substitute (Buck et al., 1989; Murray, 1988). Furthermore, polymers have also traditionally been used as barrier membranes in the periodontal guided tissue regeneration procedures (Murphy and Gunsolley, 2003). 15.3.4.2 Calcium phosphates Calcium phosphate biomaterials have outstanding properties: a similar composition to bone mineral, bioactivity, ability to promote cellular function and expression leading to formation of a uniquely strong bone biomaterial interface and osteoconductivity. In addition, calcium phosphates with appropriate 3D geometry are able to bind and concentrate endogenous BMPs in circulation, may become osteoinductive, and can be effective carriers of bone cells. Therefore calcium phosphates are potentially useful in bone tissue engineering. Hydroxyapatite This is a commonly used calcium phosphate biomaterial for bone regeneration applications due to having a composition and structure similar to natural bone mineral (Table 15.5) (Wang et al., 2007). HA-based grafts form a chemical bond directly to bone once implanted (Bagambisa et al., 1993). Synthetic HA is available and used in various forms: (1) porous nonresorbable, (2) solid nonresorbable, and (3) resorbable (nonceramic, porous) (Tevlin et al., 2014). HA is nonosteogenic and mainly functions as an osteoconductive graft material. HA grafts show slow and limited resorptive potential and generally are dependent on passive dissolution in
Table 15.5 List of existing calcium phosphate compounds (Fernandez et al., 1999; Dorozhkin, 2009; Dorozhkin and Epple, 2002). Ca/P ionic ratio Density (g/cm3) Solubility at 37 C, 2 log(ksp) MCPM 0.5 2.23 Not available Ca(H2PO4)2 MCPA 0.5 2.58 Not available CaHPO4.2H2O DCPD Brushite 1.0 2.38 6.63 CaHPO4 DCPA Monetite 1.0 2.89 7.02 Ca8(HPO4)2(PO4)4 5H2O α-Ca3(PO4)2 β-Ca3(PO4)2 CaxHy(PO4)z nH2O, n 5 3 4.5; 15% 20% H2O Ca82x(HPO4)x(PO4)62x(OH)22x OCP α-TCP B-TCP ACP 1.33 1.5 1.5 1.2 2.2 2.61 2.86 3.07 3.01 95.9 25.5 29.5 25 32 PHA 1.33 1.67 3.16 Not available Ca102x(HPO4)x(PO4)62x (OH)22x (0 , x , 1) Ca10(PO4)6(OH)2 Ca10(PO4)6O Ca10(PO4)6F2 Ca2(PO4)2O CDHA 1.5 1.67 3.16 85.1 1.67 1.67 1.67 2.0 3.16 3.20 3.18 3.05 117.2 Not available Not available 37 42 Compound name Chemical formula Symbol Monocalcium phosphate monohydrate Monocalcium phosphate anhydrous Dicalcium phosphate dehydrate Dicalcium phosphate anhydrous Octacalcium phosphate α-Tricalcium phosphate β-Tricalcium phosphate Amorphous calcium phosphate Precipitated hydroxyapatite Calcium-deficient hydroxyapatite Hydroxyapatite Oxyapatite Fluorapatite Tetracalcium phosphate Ca(H2PO4)2 H2O HA OXA FA TTCP Mineral Hydroxyapatite Hilgenstockite
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 359 tissue fluid and cell-mediated processes such as phagocytosis of particles for resorption (Jarcho, 1981; De Groot, 1980). The degradation rate of HA depends on the method of ceramic formation, the calcium to phosphate ratio, crystallographic structure, and porosity (Osborn and Newesely, 1980; Jarcho, 1981). The ability of HA to resorb is also heavily dependent upon the processing temperature. HA grafts synthesized at high temperatures are very dense with very limited biodegradability (Klein et al., 1983). These dense grafts are usually used as inert biocompatible fillers (Rabalais et al., 1981; Meffert et al., 1985). At lower temperatures, the particulate HA is porous and undergoes slow resorption (Ricci et al., 1992). Early implant loading studies in alveolar ridges augmented with nanostructured HA have shown promise (Canullo et al., 2006; Sugar et al., 1995). Also, ridge augmentation with HA granules alone (Sugar et al., 1995) or in combination with autografts has been investigated (Small et al., 1993). Tricalcium phosphate Over the last few years, TCP has been used and extensively investigated as a bone substitute. TCP has two crystallographic forms; α-TCP and β-TCP (Tamimi et al., 2012b). β-TCP exhibits good biocompatibility and osteoconductivity and is used commonly as a partially resorbable filler allowing replacement with newly formed bone (Shetty and Han, 1991). Resorption of TCP grafts is thought to be dependent on dissolution by biological fluids in the absence of osteoclasts around the materials (Lu et al., 1998) and by presence of osteoclast-mediated resorption based on the osteoclast-like giant cells in defect areas in many studies (Renooij et al., 1985). In terms of bone regenerative potential, β-TCP grafts have been shown to be similar to autogenous bone, FDBA, DFDBA, and collagen sponge (Nakajima et al., 2007). TCP biomaterials have been used in human clinical studies to repair periapical and marginal periodontal defects, as well as alveolar bony defects (Stavropoulos et al., 2010; Metsger et al., 1982). In addition, there are studies using β-TCP that report alveolar ridge augmentation in vertical and horizontal dimensions with variable results (Shalash et al., 2013; Nyan et al., 2014). Dicalcium phosphates These are acidic calcium phosphates that have a high solubility at physiological pH. Dicalcium phosphate dihydrate (DCPD or Brushite) has been investigated for both bone defect repair and vertical bone augmentation applications as injectable cements or as preset cement granules (Tamimi et al., 2009, 2010; Sheikh et al., 2016). It has been demonstrated that injectable brushite cements are capable of regenerating bone in atrophic alveolar ridges, buccal dehiscence defects, and maxillary sinus floor elevation procedures (Gehrke and Famà, 2010). Bone growth in the vertical direction obtained with brushite cement granules has been seen to be higher than that obtained with commercially available bovine HA materials (Tamimi et al., 2006). However, brushite grafts after implantation undergo phase conversion to insoluble HA that ultimately limits their resorption rate and extent (Sheikh et al., 2015f; Tamimi et al., 2012b). Brushite can be used as precursor to the anhydrous form of DCP, dicalcium phosphate anhydrous, also known as DCPA
360 Advanced Dental Biomaterials Figure 15.2 Scanning electron microscopic micrographs of (A) hydroxyapatite; (B) β-tricalcium phosphate; (C) octacalcium phosphate; (D) dicalcium phosphate dihydrate (brushite); (E) calcium polyphosphate; and (F) dicalcium phosphate anhydrous (monetite). or monetite. Monetite can be precipitated by dehydration of brushite or by modifying the precipitation conditions of brushite cements in order to favor DCP crystallization into monetite instead of brushite (Tamimi et al., 2012b). Monetite does not convert to HA after implantation (Sheikh et al., 2015f, 2016, 2017c; Tamimi et al., 2006, 2008) and resorbs at faster rates compared to brushite cement grafts (Gbureck et al., 2007; Tamimi et al., 2012a; Idowu et al., 2014; Sheikh et al., 2017c). Monetite granules have been compared with commercially available bovine HA (Bio-Oss), and have shown greater resorption and bone formation in the extraction sockets (Tamimi et al., 2010) (Fig. 15.2). 15.3.4.3 Bioactive glasses These graft materials are composed of silicon dioxide, calcium oxide, sodium oxide, and phosphorus pentoxide (Schepers et al., 1991; Shue et al., 2012). The particle sizes of bioactive glasses (Bio-Glass) range from 90 710 to 300 355 μm (Schepers et al., 1991; Schepers and Ducheyne, 1997). After implantation of bioactive glass, a silicon-rich gel is formed on the bioactive ceramic surface with the outer layer serving as a bonding surface for osteogenic cells and collagen fibers
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 361 (Hall et al., 1999; Xynos et al., 2001). Bioactive glass nanoparticles have been shown to induce cementoblasts to proliferate in an in vivo study (Carvalho et al., 2012). Clinical reports of alveolar ridge grafting performed with bioactive glass reveal bone formation in close contact to the particles (Schepers et al., 1991). However, limited true periodontal regenerative outcomes based on human histological analysis have been demonstrated with the use of bioactive glass (Nevins et al., 2000; Knapp et al., 2003). 15.3.4.4 Calcium polyphosphate Inorganic polyphosphates (polyP) are polymers of orthophosphate, linked by energy-rich phosphoanhydride bonds to form polymeric chains. These phosphate polymers exhibit varied biological roles and modulate many biological processes. Calcium polyphosphate (CPP) was first developed as a bone substitute to address some of the shortcomings of traditional materials in use for bone healing, such as HA, TCP, and bioactive glasses. CPP is a good bone substitute as it can be made with mechanical properties similar to trabecular bone, controlled degradability, and shows very good integration to host bone when implanted in vivo (Grynpas et al., 2002). CPP has been used in different forms, such as sintered porous blocks (Pilliar et al., 2001), particulates (Pilliar et al., 2016), or nanoparticles (Chen et al., 2015). This flexibility in fabrication has made CPP a good material for use in bone regeneration. Nelson et al. were the first to investigate CPP for bone regeneration as they explored its ability to repair canine mandibular alveolar defects. The bone defect was either filled with CPP or autologous bone. Assessment at 4 months showed increased bone and greater rates of union in the CPP group than in the bone graft control (Nelson et al., 1993). There have been multiple studies examining tissue engineering of oral or dental tissues with CPP. El Sayegh et al. (2002) demonstrated that the degradation rate of CPP did not substantially affect the interactions of human gingival fibroblasts with CPP materials but that compared with titanium alloy substrates, cell spreading and attachment were inhibited. These studies suggest that CPP has promise as a biomaterial for biological and periodontal regenerative therapy (Wang et al., 2006). 15.3.4.5 Glass ionomers Glass-ionomer cements were first used in dentistry as calcium/aluminum/fluorosilicate glass powder mixed with polycarboxylic acid to produce a cement paste that hardens in approximately 5 minutes after which it is rendered water insoluble. Bone replacement grafts based on glass-ionomer cements are porous, which leads to bone ingrowth via their osteoconductive nature. After 24 hours of setting they have a compressive strength and modulus of elasticity (stiffness) that is comparable with cortical bone. Glass ionomers have been considered as a replacement for PMMAbased bone fillers, which have an exothermic setting reaction (polymerization) (Jonck et al., 1989). A possible advantage of using this material could be their fluoride release mechanism that may be adapted for controlled (slow release profile)
362 Advanced Dental Biomaterials drug delivery (antibiotics) in the local periodontal defect environment. A major disadvantage is the lack of resorbability (Denissen et al., 1997), limiting its clinical use as a scaffold for bone regeneration application. 15.3.4.6 Calcium sulfate These compounds have a compressive strength greater than that of cancellous bone (Moore et al., 2001). Calcium sulfate is usually applied as a barrier material to improve the clinical outcomes of periodontal regeneration therapy (Pecora et al., 1997). When used as a barrier, calcium sulfate materials work as an adjunct with other graft materials. A combination of β-TCP and calcium phosphate has been investigated, which does not require a membrane, lowers cost, reduces surgical time, and has the potential to treat periodontal intrabony defects (Paolantonio et al., 2008; Sukumar et al., 2011). A randomized controlled clinical trial over 12 months has shown that the use of calcium sulfate is useful in minimizing postsurgical recession when compared with the use of collagen membrane (Paolantonio et al., 2008; Sheikh et al., 2017b; Sheikh, 2014). The clinical outcome of class II mandibular molar furcation defects has also been shown to be enhanced with the use of a mixture of calcium sulfate and DFDBA (Maragos et al., 2002). 15.3.4.7 Magnesium-based biodegradable materials Elemental magnesium (Mg) was discovered in 1808 and Mg and its alloys have generated significant interest for use in biomedical applications as implants, osteosynthesis devices, ligatures, wires for aneurysm treatment, and connectors for vessel anastomosis (Seal et al., 2009; Staiger et al., 2006). Mg21 is a cation that is mostly stored in bone tissues and is the fourth most abundant ion in the human body. Mgbased metals corrode in aqueous environments via electrochemical reactions that result in the production of Mg hydroxide and hydrogen gas (Staiger et al., 2006). The corrosion product of Mg (Mg21) is easily excreted in urine resulting in the good biological behavior observed when Mg and its alloys are used for medical applications (Staiger et al., 2006). Mg-based biomaterials have better mechanical properties when compared with other conventional biodegradable materials such as polymers and ceramics (Tan et al., 2013). Based on the distinct advantages of Mg-based metals, they have been extensively investigated both in vitro and in vivo for osteologic repair and regeneration applications. Mostly the focus has been on fabricating screws and plates for fracture fixation and porous scaffold (Staiger et al., 2006). However, since these have inferior mechanical properties than the conventional metallic nondegradable devices, Mgbased devices are not being used for load bearing application (Witte et al., 2008; Brar et al., 2009; Tan et al., 2013). After implantation of Mg-Ca alloy pins in rabbit femoral shafts, no cytotoxicity was observed and elevated activity of osteocytes and osteoblasts was shown around the implants indicating good biocompatibility and bioactivity (Li et al., 2008).
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 15.4 363 Biocompatibility of bone replacement graft materials and their degradation products To perform successfully, biomaterial grafts must not cause abnormal responses in local tissues and should not produce toxic or carcinogenic effects. Biodegradable materials in particular should serve their intended function while releasing products of degradation that are biocompatible and do not interfere with tissue healing (Ratner, 2004). A major concern associated with using biodegradable materials especially polymers is the possibility of local inflammation due to themselves or via their degradation products (Gooptu and Lomas, 2008). Once implanted, the biodegradation and resorption processes begin and are accompanied by a release of acidic by-products, which can result in inflammatory reactions (Van Der Giessen et al., 1996). If the capacity of the surrounding tissue to eliminate the by-products is low, due to the poor vascularization or low metabolic activity, the chemical composition of the by-products may lead to local and systemic disturbances (Anderson, 1988). The inflammatory response to polymer degradation can be controlled somewhat by the incorporation of basic salts such as sodium bicarbonate, calcium bicarbonate, and calcium HA (Hollinger and Chaudhari, 1992). Also, the incorporation of TCP (Hutmacher et al., 2001), HA (Shikinami and Okuno, 1999), and basic salts (Agrawal and Athanasiou, 1997) into the polymeric matrix results in the production of a hybrid/composite material. Calcium phosphate based bioceramics are also widely used for bone regeneration applications. Biodegradable dicalcium phosphates are generally well tolerated by bone and soft tissues and do not cause inflammations in the long term (Apelt et al., 2004; Lu et al., 1999). Following implantation these grafts are enclosed in loose connective tissue (Frayssinet et al., 1998), although they can also be surrounded by fibrous connective tissue if the cement composition is acidic (Tamimi et al., 2009). In vivo studies have shown that early resorption of calcium phosphate cements is regulated by macrophages rather than osteoclasts (Constantz et al., 1998; Kuemmerle et al., 2005; Sheikh et al., 2015a,b). Similar to in vitro studies, implanted cement grafts can resorb via disintegration/fragmentation and rather passive dissolution based upon the solubility constant product of the material (Grover et al., 2003). This is critical, since it is known that particles released from calcium phosphate grafts can affect osteoblast function, viability, and proliferation and production of extracellular matrix adversely (Tamimi et al., 2012b). The maximum number of particles that a single osteoblast can support is B50, and the smaller the disintegration products are, the stronger the negative effect is observed (Pioletti et al., 2000). These released particles can also potentially result in peri-implant osteolysis and failure if the microenvironment around the implanted biomaterial is not cleared by extracellular media refreshment (Meijer et al., 2008). 15.5 Biodegradation of implanted graft materials and bone formation The resorptive capability of implanted bone replacement graft materials allows for space to be produced for newly forming bone tissue to not only grow along the
364 Advanced Dental Biomaterials implant surface (creeping substitution via osteoconduction) but also to infiltrate within the resorbing matrix along with new blood vessels (Sheikh et al., 2015e). This infiltration of biomaterial scaffold matrix with blood vessels allows for the bone formation front to progress and be provided with oxygen that is mandatory for the survival of the regenerating tissues (Klein-Nulend et al., 2005). It is crucial for the biodegradable scaffold to retain its strength during the healing period so as to provide fixation at the fracture site but degrade after the healing as completed. Initial resorption of calcium phosphate cement grafts is affected by the inherent cement properties such as porosity, as well as the site of implantation, which affects the rate of fluid exchange and the properties of the surrounding medium (Grover et al., 2003; Sheikh et al., 2015c,f). The amount of new bone formed is also highly dependent on implantation site and vascular supply, as an adequate blood supply increases the speed of cement resorption and replacement by new woven bone (Constantz et al., 1998). It is known for serum proteins to be adsorbed onto the cement surface, altering the interfacial properties of the calcium phosphate crystals (Giocondi et al., 2010), and favoring in vivo resorption (Grover et al., 2003). Research shows that unlike HA cements that undergo negligible resorption over time, dicalcium phosphate cements resorb to a much greater extent in vivo (Oberle et al., 2005; Apelt et al., 2004). The effect of particle size of the implanted graft materials affects the rate and effectiveness of cellular resorption activity (Hannink and Arts, 2011). The cells that take part in cell-mediated calcium phosphate resorption may be osteoclasts, multinucleated giant cells, monocytes, and macrophages directly available in the bone marrow tissue. Phagocytic mechanisms regulated by the monocytes/macrophages or acidic mechanisms via osteoclasts (by reduction of pH in the microenvironment) result in bioresorption of calcium phosphate cements in vivo (Heymann et al., 1999). Macrophages respond to small fragments and particles (,10 μm in diameter) by internalization via phagocytosis and intracellular digestion (Fig. 15.3). If the particle size is larger than 10 μm and smaller than 100 μm, the macrophages fuse together forming giant cells which in turn engulf the particles and digest them (Fig. 15.3) (Sheikh et al., 2015b). If the particles are larger, the bulk resorption is carried out via osteoclast-mediated resorption (Sheikh et al., 2015b; Xia and Triffitt, 2006). As mentioned earlier, the resorption of cement matrix is an important feature with respect to bone formation at the implanted sites, since it frees up the space needed for new bone formation ideally without compromising mechanical stability. The surfaces of dicalcium phosphates have been shown to stimulate osteoblasts activity (Klammert et al., 2009). The release of growth factors incorporated into cement matrices has also been used to stimulate the bone formation. Vascular endothelial growth factor (VEGF), PDGF, and receptor activator of nuclear factor jB ligand (RANKL) are some of the growth factors that have been assessed to enhance bone regenerative capacity in vivo (Le Nihouannen et al., 2008; Tamimi et al., 2012b). Bone formation has been observed to be considerably greater with PDGFloaded brushite chitosan scaffolds, as well as with the combination PDGF/VEGF (De la Riva et al., 2010). RANKL is a growth factor that promotes osteoclast differentiation and is important toward the biodegradation of calcium phosphate grafts
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 365 Figure 15.3 (A) Macrophages respond to small fragments and particles (,10 μm in diameter) by internalization via phagocytosis and intracellular digestion. (B) If the particle size is larger than 10 μm and smaller than 100 μm, the macrophages fuse together forming giant cells which in turn engulf the particles and digest them. (C) If the particles are larger, the bulk digestion is carried out via extracellular degradation by macrophages and macrophage fused giant cells through release of enzymes and/or pH lowering mechanisms. (Silva and Branco, 2010). Results from studies suggest that the application of growth factors using biodegradable materials could improve the tissue response and promote bone formation in bone regeneration applications (Le Nihouannen et al., 2008). 15.6 Future of bone tissue graft materials There is an increasing demand to replace or restore the function of lost, damaged human bone. Autologous bone grafts remain the gold standard for bone reconstruction, particularly in large bone defects; however, due to their limited supply and other drawbacks, alternative strategies are essentially required to regenerate and/or restore bone. In the past several decades, there has been tremendous research activity in the field of bone engineering and increased understanding of bone biology that has translated into the development of new possible approaches to regenerate bone. Despite all these vast advancements in the field of bone tissue engineering, the development of synthetic bone graft materials for maxillofacial and periodontal applications remains a challenge from a biological, engineering, and surgical perspective. Possibly one of the biggest challenges facing bone tissue engineering is fabricating a mechanically strong porous interconnected scaffold that promotes favorable host integration and proper vascularization. Currently, most of the synthetic grafts
366 Advanced Dental Biomaterials experience bone tissue regeneration that is limited to the outer surface layers of the biomaterial upon implantation, as a result of insufficient vascularization of the scaffold. Therefore there is a need to investigate the complex biological parameters to be considered when designing the synthetic materials for proper bone regeneration and vascularization (e.g., pore size, porosity, resorption rate). Moreover, the interplay between inflammatory response, cells and the biomaterial surface, and other molecules involved in the bone formation/bone resorption pathways is far from being fully understood. One approach to tackle this problem would be the incorporation of immunomodulatory strategies to modulate the host’s foreign-body response (i.e., fibrous tissue encapsulation). In addition, further investigation is needed to understand how endogenously or exogenously supplied growth factors with the grafting material can promote favorable cellular differentiation and reproduction as well as inhibit unfavorable inflammatory reactions. The challenges in the field are further compounded by the high cost of some of the synthetic alternatives and the lack of strong evidence (i.e., lack of randomized controlled human studies), thus being scarcely recommendable for clinical use. Accordingly, the future of bone tissue engineering primarily depends on the ability of clinicians, researchers, and engineers to communicate together to overcome the previous hurdles and fabricate an optimal clinically relevant graft with the best possible clinical outcomes, and the lowest possible morbidity and cost. References Agrawal, C.M., Athanasiou, K.A., 1997. Technique to control pH in vicinity of biodegrading PLA-PGA implants. J. Biomed. Mater. Res. 38, 105 114. Alghamdi, A.S., Shibly, O., Ciancio, S.G., 2010. Osseous grafting part II: xenografts and alloplasts for periodontal regeneration—a literature review. J. Int. Acad. Periodontol. 12, 39 44. Al Ruhaimi, K.A., 2001. Bone graft substitutes: a comparative qualitative histologic review of current osteoconductive grafting materials. Int. J. Oral Maxillofac. Implants 16, 105 114. Aludden, H., Mordenfeld, A., Hallman, M., Dahlin, C., Jensen, T., 2017. Lateral ridge augmentation with Bio-Oss alone or Bio-Oss mixed with particulate autogenous bone graft: a systematic review. Int. J. Oral Maxillofac. Surg. 46, 1030 1038. Anderson, J.M., 1988. Inflammatory response to implants. ASAIO J. 34, 101 107. Apelt, D., Theiss, F., El-Warrak, A.O., Zlinszky, K., Bettschart-Wolfisberger, R., Bohner, M., et al., 2004. In vivo behavior of three different injectable hydraulic calcium phosphate cements. Biomaterials 25, 1439 1451. Araujo, P.P., Oliveira, K.P., Montenegro, S.C., Carreiro, A.F., Silva, J.S., Germano, A.R., 2013. Block allograft for reconstruction of alveolar bone ridge in implantology: a systematic review. Implant. Dent. 22, 304 308. Asghari, F., Samiei, M., Adibkia, K., Akbarzadeh, A., Davaran, S., 2017. Biodegradable and biocompatible polymers for tissue engineering application: a review. Artif. Cells Nanomed. Biotechnol. 45, 185 192. Bagambisa, F.B., Joos, U., Schilli, W., 1993. Mechanisms and structure of the bond between bone and hydroxyapatite ceramics. J. Biomed. Mater. Res. A 27, 1047 1055.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 367 Berglundh, T., Lindhe, J., 1997. Healing around implants placed in bone defects treated with Bio-Osss. An experimental study in the dog. Clin. Oral. Implants. Res. 8, 117 124. Blaggana, V., Gill, A.S., Blaggana, A., 2014. A clinical and radiological evaluation of the relative efficacy of demineralized freeze-dried bone allograft versus anorganic bovine bone xenograft in the treatment of human infrabony periodontal defects: a 6 months follow-up study. J. Indian Soc. Periodontol. 18, 601 607. Block, M.S., Degen, M., 2004. Horizontal ridge augmentation using human mineralized particulate bone: preliminary results. J. Oral Maxillofac. Surg. 62, 67 72. Boyne, P.J., 1988. Bone induction and the use of HTR polymer as a vehicle for osseous inductor materials. Compend. Suppl. S337 S341. Brar, H.S., Platt, M.O., Sarntinoranont, M., Martin, P.I., Manuel, M.V., 2009. Magnesium as a biodegradable and bioabsorbable material for medical implants. JOM 61, 31 34. Buck, B., Malinin, T.I., Brown, M.D., 1989. Bone transplantation and human immunodeficiency virus: an estimate of risk of acquired immunodeficiency syndrome (AIDS). Clin. Orthop. Relat. Res. 240, 129 136. Canullo, L., Trisi, P., Simion, M., 2006. Vertical ridge augmentation around implants using e-PTFE titanium-reinforced membrane and deproteinized bovine bone mineral (bio-oss): a case report. Int. J. Periodontics. Restorative. Dent. 26, 355 361. Carvalho, S.M., Oliveira, A.A., Jardim, C.A., Melo, C., Gomes, D.A., Leite, M., et al., 2012. Characterization and induction of cementoblast cell proliferation by bioactive glass nanoparticles. J. Tissue Eng. Regen. Med. 6, 813 821. Chen, X.R., Bai, J., Yuan, S.J., Yu, C.X., Huang, J., Zhang, T.L., et al., 2015. Calcium phosphate nanoparticles are associated with inorganic phosphate-induced osteogenic differentiation of rat bone marrow stromal cells. Chem. Biol. Interact. 238, 111 117. Cho, B.C., Chung, H.Y., Lee, D.G., Yang, J.D., Park, J.W., Roh, K.H., et al., 2005. The effect of chitosan bead encapsulating calcium sulfate as an injectable bone substitute on consolidation in the mandibular distraction osteogenesis of a dog model. J. Oral Maxillofac. Surg. 63, 1753 1764. Committee On Research, S, 2001. Tissue banking of bone allografts used in periodontal regeneration. J. Periodontol. 72, 834. Constantz, B.R., Barr, B.M., Ison, I.C., Fulmer, M.T., Baker, J., Mckinney, L., et al., 1998. Histological, chemical, and crystallographic analysis of four calcium phosphate cements in different rabbit osseous sites. J. Biomed. Mater. Res. 43, 451 461. Contar, C.M., Sarot, J.R., Bordini JR., J., Galvao, G.H., Nicolau, G.V., Machado, M.A., 2009. Maxillary ridge augmentation with fresh-frozen bone allografts. J. Oral Maxillofac. Surg. 67, 1280 1285. Contar, C.M., Sarot, J.R., Da Costa, M.B., Bordini, J., De Lima, A.A., et al., 2011. Freshfrozen bone allografts in maxillary ridge augmentation: histologic analysis. J. Oral Implantol. 37, 223 231. Cornell, C.N., Lane, J.M., 1998. Current understanding of osteoconduction in bone regeneration. Clin. Orthop. Relat. Res. 355, S267 S273. Cypher, T.J., Grossman, J.P., 1996. Biological principles of bone graft healing. J. Foot Ankle Surg. 35, 413 417. De Groot, K., 1980. Bioceramics consisting of calcium phosphate salts. Biomaterials 1, 47 50. De La Riva, B., Sanchez, E., Hernandez, A., Reyes, R., Tamimi, F., Lopez-Cabarcos, E., et al., 2010. Local controlled release of VEGF and PDGF from a combined brushitechitosan system enhances bone regeneration. J. Control Release 143, 45 52.
368 Advanced Dental Biomaterials Denissen, H., Van Beek, E., Martinetti, R., Klein, C., Van Der Zee, E., Ravaglioli, A., 1997. Net-shaped hydroxyapatite implants for release of agents modulating periodontal-like tissues. J. Periodontal Res. 32, 40 46. Dias, R.R., Sehn, F.P., De Santana Santos, T., Silva, E.R., Chaushu, G., Xavier, S.P., 2014. Corticocancellous fresh-frozen allograft bone blocks for augmenting atrophied posterior mandibles in humans. Clin. Oral Implants Res. 27 (1), 39 46. Dodson, S.A., Bernard, G.W., Kenney, E.B., Carranza, F.A., 1996. In vitro comparison of aged and young osteogenic and hemopoietic bone marrow stem cells and their derivative colonies. J. Periodontol. 67, 184 196. Dorozhkin, S.V., 2009. Calcium orthophosphates in nature, biology and medicine. Materials 2, 399 498. Dorozhkin, S.V., Epple, M., 2002. Biological and medical significance of calcium phosphates. Angew. Chem. Int. Ed. 41, 3130 3146. Draenert, F.G., Huetzen, D., Neff, A., Mueller, W.E., 2014. Vertical bone augmentation procedures: basics and techniques in dental implantology. J. Biomed. Mater. Res. A 102, 1605 1613. El Sayegh, T.Y., Pilliar, R.M., Mcculloch, C.A., 2002. Attachment, spreading, and matrix formation by human gingival fibroblasts on porous-structured titanium alloy and calcium polyphosphate substrates. J. Biomed. Mater. Res. 61, 482 492. Ewers, R., 2005. Maxilla sinus grafting with marine algae derived bone forming material: a clinical report of long-term results. J. Oral Maxillofac. Surg. 63, 1712 1723. Felice, P., Marchetti, C., Iezzi, G., Piattelli, A., Worthington, H., Pellegrino, G., et al., 2009. Vertical ridge augmentation of the atrophic posterior mandible with interpositional bloc grafts: bone from the iliac crest vs. bovine anorganic bone. Clinical and histological results up to one year after loading from a randomized-controlled clinical trial. Clin. Oral Implants Res. 20, 1386 1393. Fernández, E., Gil, F., Ginebra, M., Driessens, F., Planell, J., Best, S., 1999. Production and characterization of new calcium phosphate bone cements in the CaHPOv 4 α-Ca 3 (PO 4) 2 system: pH, workability and setting times. J. Mater. Sci.: Mater. Med. 10, 223 230. Festa, V.M., Addabbo, F., Laino, L., Femiano, F., Rullo, R., 2013. Porcine-derived xenograft combined with a soft cortical membrane versus extraction alone for implant site development: a clinical study in humans. Clin. Implant Dent Relat. Res. 15, 707 713. Finkemeier, C.G., 2002. Bone-grafting and bone-graft substitutes. J. Bone Joint Surg. 84, 454 464. Frayssinet, P., Gineste, L., Rouquet, N., 1998. Osteointegration of different kinds of calcium phosphate biomaterials: ceramics and hydraulic cements. Morphologie 82, 3 7. Friedlaender, G.E., Strong, D.M., Sell, K.W., 1976. Studies on the antigenicity of bone. I. Freeze-dried and deep-frozen bone allografts in rabbits. J. Bone Joint Surg. Am. 58, 854 858. Froum, S.J., 1996. Human histologic evaluation of HTR polymer and freeze-dried bone allograft. A case report. J. Clin. Periodontol. 23, 615 620. Gbureck, U., Hölzel, T., Klammert, U., Würzler, K., Müller, F.A., Barralet, J.E., 2007. Resorbable dicalcium phosphate bone substitutes prepared by 3D powder printing. Adv. Funct. Mater. 17, 3940 3945. Gehrke, S., Famà, G., 2010. Buccal dehiscence and sinus lift cases predictable bone augmentation with synthetic bone material. Implants 11, 4. Giocondi, J.L., El-Dasher, B.S., Nancollas, G.H., Orme, C.A., 2010. Molecular mechanisms of crystallization impacting calcium phosphate cements. Philos. Trans. R. Soc. Lond. A: Math. Phys. Eng. Sci. 368, 1937 1961.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 369 Gooptu, B., Lomas, D.A., 2008. Polymers and inflammation: disease mechanisms of the serpinopathies. J. Exp. Med. 205, 1529 1534. Grover, L.M., Knowles, J.C., Fleming, G.J.P., Barralet, J.E., 2003. In vitro ageing of brushite calcium phosphate cement. Biomaterials. 24, 4133 4141. Grynpas, M.D., Pilliar, R.M., Kandel, R.A., Renlund, R., Filiaggi, M., Dumitriu, M., 2002. Porous calcium polyphosphate scaffolds for bone substitute applications in vivo studies. Biomaterials 23, 2063 2070. Hall, J.E., 2015. Guyton and Hall Textbook of Medical Physiology e-Book. Elsevier Health Sciences. Hall, E.E., Meffert, R.M., Hermann, J.S., Mellonig, J.T., Cochran, D.L., 1999. Comparison of bioactive glass to demineralized freeze-dried bone allograft in the treatment of intrabony defects around implants in the canine mandible. J. Periodontol. 70, 526 535. Hannink, G., Arts, J.C., 2011. Bioresorbability, porosity and mechanical strength of bone substitutes: what is optimal for bone regeneration? Injury 42, S22 S25. Hench, L.L., 1998. Bioactive materials: the potential for tissue regeneration. J. Biomed. Mater. Res. 41, 511 518. Heymann, D., Pradal, G., Benahmed, M., 1999. Cellular mechanisms of calcium phosphate ceramic degradation. Histol. Histopathol. 14, 871 877. Hollinger, J., Chaudhari, A., 1992. Bone regeneration materials for the mandibular and craniofacial complex. Cells Mater. 2, 143 151. Hopp, S.G., Dahners, L.E., Gilbert, J.A., 1989. A study of the mechanical strength of long bone defects treated with various bone autograft substitutes: an experimental investigation in the rabbit. J. Orthop. Res. 7, 579 584. Hutmacher, D.W., Kirsch, A., Ackermann, K.L., Hürzeler, M.B., 2001. A tissue engineered cell-occlusive device for hard tissue regeneration—a preliminary report. Int. J. Periodontics Restorative Dent. 21, 49 59. Idowu, B., Cama, G., Deb, S., Di Silvio, L., 2014. In vitro osteoinductive potential of porous monetite for bone tissue engineering. J. Tissue Eng. 5, 1 8. 2041731414536572. Jackson, I.T., Helden, G., Marx, R., 1986. Skull bone grafts in maxillofacial and craniofacial surgery. J. Oral Maxillofac. Surg. 44, 949 955. Jacotti, M., Wang, H.L., Fu, J.H., Zamboni, G., Bernardello, F., 2012. Ridge augmentation with mineralized block allografts: clinical and histological evaluation of 8 cases treated with the 3-dimensional block technique. Implant. Dent. 21, 444 448. Jarcho, M., 1981. Calcium phosphate ceramics as hard tissue prosthetics. Clin. Orthop. Relat. Res. 157, 259 278. Jergesen, H.E., Chua, J., Kao, R.T., Kaban, L.B., 1991. Age effects on bone induction by demineralized bone powder. Clin. Orthop. Relat. Res. 253 259. Jonck, L., Grobbelaar, C., Strating, H., 1989. Biological evaluation of glass-ionomer cement (Ketac-0) as an interface material in total joint replacement. A screening test. Clin. Mater. 4, 201 224. Klammert, U., Reuther, T., Jahn, C., Kraski, B., Kubler, A.C., Gbureck, U., 2009. Cytocompatibility of brushite and monetite cell culture scaffolds made by threedimensional powder printing. Acta Biomater. 5, 727 734. Klein, C.P., Driessen, A.A., De Groot, K., Van Den Hooff, A., 1983. Biodegradation behavior of various calcium phosphate materials in bone tissue. J. Biomed. Mater. Res. 17, 769 784. Klein-Nulend, J., Bacabac, R., Mullender, M., 2005. Mechanobiology of bone tissue. Pathol. Biol. 53, 576 580. Knapp, C.I., Feuille, F., Cochran, D.L., Mellonig, J.T., 2003. Clinical and histologic evaluation of bone-replacement grafts in the treatment of localized alveolar ridge defects. Part 2: Bioactive glass particulate. Int. J. Periodontics Restorative Dent. 23, 1 12.
370 Advanced Dental Biomaterials Kuemmerle, J.M., Oberle, A., Oechslin, C., Bohner, M., Frei, C., Boecken, I., et al., 2005. Assessment of the suitability of a new brushite calcium phosphate cement for cranioplasty an experimental study in sheep. J. Cranio-Maxillofac. Surg. 33, 37 44. Kukreja, B.J., Dodwad, V., Kukreja, P., Ahuja, S., Mehra, P., 2014. A comparative evaluation of platelet-rich plasma in combination with demineralized freeze-dried bone allograft and DFDBA alone in the treatment of periodontal intrabony defects: a clinicoradiographic study. J. Indian Soc. Periodontol. 18, 618 623. Le Nihouannen, D., Hacking, S.A., Gbureck, U., Komarova, S.V., Barralet, J.E., 2008. The use of RANKL-coated brushite cement to stimulate bone remodelling. Biomaterials 29, 3253 3259. Lewandrowski, K.U., Gresser, J.D., Wise, D.L., Trantolo, D.J., 2000. Bioresorbable bone graft substitutes of different osteoconductivities: a histologic evaluation of osteointegration of poly(propylene glycol-co-fumaric acid)-based cement implants in rats. Biomaterials 21, 757 764. Li, Z., Gu, X., Lou, S., Zheng, Y., 2008. The development of binary Mg-Ca alloys for use as biodegradable materials within bone. Biomaterials 29, 1329 1344. Liu, X., Li, Q., Wang, F., Wang, Z., 2016. Maxillary sinus floor augmentation and dental implant placement using dentin matrix protein-1 gene-modified bone marrow stromal cells mixed with deproteinized boving bone: a comparative study in beagles. Arch. Oral. Biol. 64, 102 108. Lu, J., Gallur, A., Flautre, B., Anselme, K., Descamps, M., Thierry, B., et al., 1998. Comparative study of tissue reactions to calcium phosphate ceramics among cancellous, cortical, and medullar bone sites in rabbits. J. Biomed. Mater. Res. 42, 357 367. Lu, J.X., About, I., Stephan, G., Van Landuyt, P., Dejou, J., Fiocchi, M., et al., 1999. Histological and biomechanical studies of two bone colonizable cements in rabbits. Bone 25, 41S 45S. Lyford, R.H., Mills, M.P., Knapp, C.I., Scheyer, E.T., Mellonig, J.T., 2003. Clinical evaluation of freeze-dried block allografts for alveolar ridge augmentation: a case series. Int. J. Periodontics Restorative Dent. 23, 417 425. Macedo, L.G., Mazzucchelli-Cosmo, L.A., Macedo, N.L., Monteiro, A.S., Sendyk, W.R., 2012. Fresh-frozen human bone allograft in vertical ridge augmentation: clinical and tomographic evaluation of bone formation and resorption. Cell. Tissue Bank 13, 577 586. Maragos, P., Bissada, N.F., Wang, R., Cole, B.P., 2002. Comparison of three methods using calcium sulfate as a graft/barrier material for the treatment of Class II mandibular molar furcation defects. Int. J. Periodontics Restorative Dent. 22, 493 501. Markou, N., Pepelassi, E., Vavouraki, H., Stamatakis, H.C., Nikolopoulos, G., Vrotsos, I., et al., 2009. Treatment of periodontal endosseous defects with platelet-rich plasma alone or in combination with demineralized freeze-dried bone allograft: a comparative clinical trial. J. Periodontol. 80, 1911 1919. Mastrogiacomo, M., Scaglione, S., Martinetti, R., Dolcini, L., Beltrame, F., Cancedda, R., et al., 2006. Role of scaffold internal structure on in vivo bone formation in macroporous calcium phosphate bioceramics. Biomaterials 27, 3230 3237. Mcallister, B.S., Haghighat, K., 2007. Bone augmentation techniques. J. Periodontol. 78, 377 396. Mcallister, B.S., Margolin, M.D., Cogan, A.G., Buck, D., Hollinger, J.O., Lynch, S.E., 1999. Eighteen-month radiographic and histologic evaluation of sinus grafting with anorganic bovine bone in the chimpanzee. Int. J. Oral Maxillofac. Implants 14, 361 368.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 371 Meffert, R.M., Thomas, J.R., Hamilton, K.M., Brownstein, C.N., 1985. Hydroxylapatite as an alloplastic graft in the treatment of human periodontal osseous defects. J. Periodontol. 56, 63 73. Meijer, G.J., De Bruijn, J.D., Koole, R., Van Blitterswijk, C.A., 2008. Cell based bone tissue engineering in jaw defects. Biomaterials 29, 3053 3061. Mellonig, J.T., 1991. Freeze-dried bone allografts in periodontal reconstructive surgery. Dent. Clin. North Am. 35, 505 520. Mellonig, J.T., 2000. Human histologic evaluation of a bovine-derived bone xenograft in the treatment of periodontal osseous defects. Int. J. Periodontics Restorative Dent. 20, 19 29. Mellonig, J.T., Bowers, G.M., Bailey, R.C., 1981. Comparison of bone graft materials. Part I. New bone formation with autografts and allografts determined by Strontium-85. J. Periodontol. 52, 291 296. Merli, M., Lombardini, F., Esposito, M., 2010. Vertical ridge augmentation with autogenous bone grafts 3 years after loading: resorbable barriers versus titanium-reinforced barriers. A randomized controlled clinical trial. Int. J. Oral Maxillofac. Implants 25, 801 807. Metsger, D.S., Driskell, T., Paulsrud, J., 1982. Tricalcium phosphate ceramic—a resorbable bone implant: review and current status. J. Am. Dent. Assoc. 105, 1035 1038. Misch, C.M., 1997. Comparison of intraoral donor sites for onlay grafting prior to implant placement. Int. J. Oral Maxillofac. Implants 12, 767 776. Moore, W.R., Graves, S.E., Bain, G.I., 2001. Synthetic bone graft substitutes. ANZ. J. Surg. 71, 354 361. Murphy, K.G., Gunsolley, J.C., 2003. Guided tissue regeneration for the treatment of periodontal intrabony and furcation defects. A systematic review. Ann. Periodontol. 8, 266 302. Murray, V.K., 1988. Clinical applications of HTR polymer in periodontal surgery. Compend. Suppl. 10, S342 S347. Nakajima, Y., Fiorellini, J.P., Kim, D.M., Weber, H.P., Dent, M., 2007. Regeneration of standardized mandibular bone defects using expanded polytetrafluoroethylene membrane and various bone fillers. Int. J. Periodontics Restorative Dent. 27, 151. Nannmark, U., Sennerby, L., 2008. The bone tissue responses to prehydrated and collagenated cortico-cancellous porcine bone grafts: a study in rabbit maxillary defects. Clin. Implant. Dent. Relat. Res. 10, 264 270. Nelson, S.R., Wolford, L.M., Lagow, R.J., Capano, P.J., Davis, W.L., 1993. Evaluation of new high-performance calcium polyphosphate bioceramics as bone graft materials. J. Oral Maxillofac. Surg. 51, 1363 1371. Nevins, M.L., Camelo, M., Nevins, M., King, C.J., Oringer, R.J., Schenk, R.K., et al., 2000. Human histologic evaluation of bioactive ceramic in the treatment of periodontal osseous defects. Int. J. Periodontics Restorative Dent. 20, 458 467. Nkenke, E., Weisbach, V., Winckler, E., Kessler, P., Schultze-Mosgau, S., Wiltfang, J., et al., 2004. Morbidity of harvesting of bone grafts from the iliac crest for preprosthetic augmentation procedures: a prospective study. Int. J. Oral Maxillofac. Surg. 33, 157 163. Nyan, M., Miyahara, T., Noritake, K., Hao, J., Rodriguez, R., Kasugai, S., 2014. Feasibility of alpha tricalcium phosphate for vertical bone augmentation. J. Investig. Clin. Dent. 5, 109 116. Oberle, A., Theiss, F., Bohner, M., Muller, J., Kastner, S.B., Frei, C., et al., 2005. Investigation about the clinical use of brushite-and hydroxylapatite-cement in sheep. Schweiz. Arch. Tierheilkd. 147, 482 490.
372 Advanced Dental Biomaterials Osborn, J., Newesely, H., 1980. The material science of calcium phosphate ceramics. Biomaterials 1, 108 111. Paolantonio, M., Perinetti, G., Dolci, M., Perfetti, G., Tetè, S., Sammartino, G., et al., 2008. Surgical treatment of periodontal intrabony defects with calcium sulfate implant and barrier versus collagen barrier or open flap debridement alone: a 12-month randomized controlled clinical trial. J. Periodontol. 79, 1886 1893. Pearce, A., Richards, R., Milz, S., Schneider, E., Pearce, S., 2007. Animal models for implant biomaterial research in bone: a review. Eur. Cell. Mater. 13, 1 10. Pecora, G., Andreana, S., Margarone, J.E., Covani, U., Sottosanti, J.S., 1997. Bone regeneration with a calcium sulfate barrier. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endodontol. 84, 424 429. Pilliar, R.M., Filiaggi, M.J., Wells, J.D., Grynpas, M.D., Kandel, R.A., 2001. Porous calcium polyphosphate scaffolds for bone substitute applications—in vitro characterization. Biomaterials 22, 963 972. Pilliar, R.M., Kandel, R.A., Grynpas, M.D., Theodoropoulos, J., Hu, Y., Allo, B., et al., 2016. Calcium polyphosphate particulates for bone void filler applications. J. Biomed. Mater. Res. B. Appl. Biomater. 105 (4), 874 884. Pioletti, D.P., Takei, H., Lin, T., Van Landuyt, P., Ma, Q.J., Kwon, S.Y., et al., 2000. The effects of calcium phosphate cement particles on osteoblast functions. Biomaterials 21, 1103 1114. Proussaefs, P., Lozada, J., Rohrer, M.D., 2002. A clinical and histologic evaluation of a block onlay graft in conjunction with autogenous particulate and inorganic bovine mineral (Bio-Oss): a case report. Int. J. Periodontics Restorative Dent. 22, 567 574. Proussaefs, P., Lozada, J., Kleinman, A., Rohrer, M.D., Mcmillan, P.J., 2003. The use of titanium mesh in conjunction with autogenous bone graft and inorganic bovine bone mineral (Bio-Oss) for localized alveolar ridge augmentation: a human study. Int. J. Periodontics Restorative Dent. 23, 185 195. Quattlebaum, J.B., Mellonig, J.T., Hensel, N.F., 1988. Antigenicity of freeze-dried cortical bone allograft in human periodontal osseous defects. J. Periodontol. 59, 394 397. Rabalais J.R., M.L., Yukna, R.A., Mayer, E.T., 1981. Evaluation of durapatite ceramic as an alloplastic implant in periodontal osseous defects. I. Initial six-month results. J. Periodontol. 52, 680 689. Ratner, B.D., 2004. Biomaterials Science: An Introduction to Materials in Medicine. Academic Press. Renooij, W., Hoogendoorn, H.A., Visser, W.J., Lentferink, R.H., Schmitz, M.G., Van Ieperen, H., et al., 1985. Bioresorption of ceramic strontium-85-labeled calcium phosphate implants in dog femora: a pilot study to quantitate bioresorption of ceramic implants of hydroxyapatite and tricalcium orthophosphate in vivo. Clin. Orthop. Relat. Res. 197, 272 285. Ricci, J.L., Blumenthal, N.C., Spivak, J.M., Alexander, H., 1992. Evaluation of a lowtemperature calcium phosphate particulate implant material: physical-chemical properties and in vivo bone response. J. Oral Maxillofac. Surg. 50, 969 978. Rocchietta, I., Simion, M., Hoffmann, M., Trisciuoglio, D., Benigni, M., Dahlin, C., 2015. Vertical bone augmentation with an autogenous block or particles in combination with guided bone regeneration: a clinical and histological preliminary study in humans. Clin. Implant. Dent. Relat. Res. 18 (1), 19 29. Roy, D.M., Linnehan, S.K., 1974. Hydroxyapatite Formed From Coral Skeletal Carbonate by Hydrothermal Exchange. 247 (5438), p. 220.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 373 Russell, J., Scarborough, N., Chesmel, K., 1997. Ability of commercial demineralized freezedried bone allograft to induce new bone formation. J. Periodontol. 68, 804 806. Scarano, A., Degidi, M., Iezzi, G., Pecora, G., Piattelli, M., Orsini, G., et al., 2006. Maxillary sinus augmentation with different biomaterials: a comparative histologic and histomorphometric study in man. Implant. Dent. 15, 197 207. Schallhorn, R.G., Hiatt, W.H., Boyce, W., 1970. Iliac transplants in periodontal therapy. J. Periodontol. 41, 566 580. Schepers, E.J., Ducheyne, P., 1997. Bioactive glass particles of narrow size range for the treatment of oral bone defects: a 1-24 month experiment with several materials and particle sizes and size ranges. J. Oral Rehabil. 24, 171 181. Schepers, E., De Clercq, M., Ducheyne, P., Kempeneers, R., 1991. Bioactive glass particulate material as a filler for bone lesions. J. Oral Rehabil. 18, 439 452. Seal, C., Vince, K., Hodgson, M., 2009. Biodegradable surgical implants based on magnesium alloys—a review of current research. IOP Conference Series: Materials Science and Engineering. IOP Publishing, 012011. Shalash, M.A., Rahman, H.A., Azim, A.A., Neemat, A.H., Hawary, H.E., Nasry, S.A., 2013. Evaluation of horizontal ridge augmentation using beta tricalcium phosphate and demineralized bone matrix: a comparative study. J. Clin. Exp. Dent. 5, e253 e259. Sheikh, Z., Javaid, M.A., Abdallah, M.N., 2013. Bone replacement graft materials in dentistry. In: Khurshid, Z.S.Z. (Ed.), Dental Biomaterials (Principle and Its Application), second ed. Paramount Publishing Enterprise. Sheikh, Z., Abdallah, M.N., Hamdan, N., Javaid, M.A., Khurshid, Z., 2014. Barrier membranes for tissue regeneration and bone augmentation techniques in dentistry. In: Matilinna, K.P. (Ed.), Handbook of Oral Biomaterials. Pan Stanford Publishing, Singapore. Sheikh, Z., Abdallah, M.-N., Hanafi, A.A., Misbahuddin, S., Rashid, H., Glogauer, M., 2015a. Mechanisms of in vivo degradation and resorption of calcium phosphate based biomaterials. Materials 8, 7913 7925. Sheikh, Z., Brooks, P.J., Barzilay, O., Fine, N., Glogauer, M., 2015b. Macrophages, foreign body giant cells and their response to implantable biomaterials. Materials 8, 5671 5701. Sheikh, Z., Geffers, M., Christel, T., Barralet, J.E., Gbureck, U., 2015c. Chelate setting of alkali ion substituted calcium phosphates. Ceram. Int. 41 (8), 10010 10017. Sheikh, Z., Najeeb, S., Khurshid, Z., Verma, V., Rashid, H., Glogauer, M., 2015d. Biodegradable materials for bone repair and tissue engineering applications. Materials 8, 5744 5794. Sheikh, Z., Sima, C., Glogauer, M., 2015e. Bone replacement materials and techniques used for achieving vertical alveolar bone augmentation. Materials 8, 2953 2993. Sheikh, Z., Zhang, Y.L., Grover, L., Merle, G.E., Tamimi, F., Barralet, J., 2015f. In vitro degradation and in vivo resorption of dicalcium phosphate cement based grafts. Acta Biomater. 26, 338 346. Sheikh, Z., Drager, J., Zhang, Y.L., Abdallah, M.N., Tamimi, F., Barralet, J., 2016. Controlling bone graft substitute microstructure to improve bone augmentation. Adv. Healthc. Mater. 5 (13), 1646 1655. Sheikh, Z., Hamdan, N., Ikeda, Y., Grynpas, M., Ganss, B., Glogauer, M., 2017a. Natural graft tissues and synthetic biomaterials for periodontal and alveolar bone reconstructive applications: a review. Biomater. Res. 21, 9. Sheikh, Z., Qureshi, J., Alshahrani, A.M., Nassar, H., Ikeda, Y., Glogauer, M., et al., 2017b. Collagen based barrier membranes for periodontal guided bone regeneration applications. Odontology 105, 1 12.
374 Advanced Dental Biomaterials Sheikh, Z., Zhang, Y.L., Tamimi, F., Barralet, J., 2017c. Effect of processing conditions of dicalcium phosphate cements on graft resorption and bone formation. Acta Biomater. 53, 526 535. Shetty, V., Han, T.J., 1991. Alloplastic materials in reconstructive periodontal surgery. Dent. Clin. North Am. 35, 521 530. Shikinami, Y., Okuno, M., 1999. Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic characteristics. Biomaterials 20, 859 877. Shue, L., Yufeng, Z., Mony, U., 2012. Biomaterials for periodontal regeneration: a review of ceramics and polymers. Biomatter 2, 271 277. Silva, I., Branco, J., 2010. Rank/Rankl/opg: literature review. Acta Reumatol. Port. 36, 209 218. Simion, M., Fontana, F., 2004. Autogenous and xenogeneic bone grafts for the bone regeneration. A literature review. Minerva Stomatol. 53, 191 206. Simion, M., Jovanovic, S.A., Tinti, C., Benfenati, S.P., 2001. Long-term evaluation of osseointegrated implants inserted at the time or after vertical ridge augmentation. A retrospective study on 123 implants with 1-5 year follow-up. Clin. Oral Implants Res. 12, 35 45. Small, S.A., Zinner, I.D., Panno, F.V., Shapiro, H.J., Stein, J.I., 1993. Augmenting the maxillary sinus for implants: report of 27 patients. Int. J. Oral Maxillofac. Implants 8, 523 528. Sokolsky-Papkov, M., Agashi, K., Olaye, A., Shakesheff, K., Domb, A.J., 2007. Polymer carriers for drug delivery in tissue engineering. Adv. Drug Deliv. Rev. 59, 187 206. Stahl, S.S., Froum, S.J., Tarnow, D., 1990. Human clinical and histologic responses to the placement of HTR polymer particles in 11 intrabony lesions. J. Periodontol. 61, 269 274. Staiger, M.P., Pietak, A.M., Huadmai, J., Dias, G., 2006. Magnesium and its alloys as orthopedic biomaterials: a review. Biomaterials 27, 1728 1734. Stavropoulos, A., Windisch, P., Szendröi-Kiss, D., Peter, R., Gera, I., Sculean, A., 2010. Clinical and histologic evaluation of granular beta-tricalcium phosphate for the treatment of human intrabony periodontal defects: a report on five cases. J. Periodontol. 81, 325 334. Sterio, T.W., Katancik, J.A., Blanchard, S.B., Xenoudi, P., Mealey, B.L., 2013. A prospective, multicenter study of bovine pericardium membrane with cancellous particulate allograft for localized alveolar ridge augmentation. Int. J. Periodontics Restorative Dent. 33, 499 507. Sugar, A.W., Thielens, P., Stafford, G.D., Willins, M.J., 1995. Augmentation of the atrophic maxillary alveolar ridge with hydroxyapatite granules in a Vicryl (polyglactin 910) knitted tube and simultaneous open vestibuloplasty. Br. J. Oral Maxillofac. Surg. 33, 93 97. Sukumar, S., Drizhal, I., Paulusová, V., Bukac, J., 2011. Surgical treatment of periodontal intrabony defects with calcium sulphate in combination with beta-tricalcium phosphate: clinical observations two years post-surgery. Acta Medica (Hradec Kralove) 54, 13 20. Sunitha Raja, V., Naidu, M., 2008. Platelet-rich fibrin: evolution of a second-generation platelet concentrate. Indian J. Dent. Res. 19, 42. Tamimi, F.M., Torres, J., Tresguerres, I., Clemente, C., Lopez-Cabarcos, E., Blanco, L.J., 2006. Bone augmentation in rabbit calvariae: comparative study between Bio-Oss (R) and a novel beta-TCP/DCPD granulate. J. Clin. Periodontol. 33, 922 928.
Natural and synthetic bone replacement graft materials for dental and maxillofacial applications 375 Tamimi, F., Torres, J., Kathan, C., Baca, R., Clemente, C., Blanco, L., et al., 2008. Bone regeneration in rabbit calvaria with novel monetite granules. J. Biomed. Mater. Res. A 87A, 980 985. Tamimi, F., Torres, J., Lopez-Cabarcos, E., Bassett, D.C., Habibovic, P., Luceron, E., et al., 2009. Minimally invasive maxillofacial vertical bone augmentation using brushite based cements. Biomaterials 30, 208 216. Tamimi, F., Torres, J., Bassett, D., Barralet, J., Cabarcos, E.L., 2010. Resorption of monetite granules in alveolar bone defects in human patients. Biomaterials 31, 2762 2769. Tamimi, F., Le Nihouannen, D., Eimar, H., Sheikh, Z., Komarova, S., Barralet, J., 2012a. The effect of autoclaving on the physical and biological properties of dicalcium phosphate dihydrate bioceramics: brushite vs. monetite. Acta Biomater. 8, 3161 3169. Tamimi, F., Sheikh, Z., Barralet, J., 2012b. Dicalcium phosphate cements: brushite and monetite. Acta Biomater. 8, 474 487. Tan, L., Yu, X., Wan, P., Yang, K., 2013. Biodegradable materials for bone repairs: a review. J. Mater. Sci. Technol. 29, 503 513. Tevlin, R., Mcardle, A., Atashroo, D., Walmsley, G.G., Senarath-Yapa, K., Zielins, E.R., et al., 2014. Biomaterials for craniofacial bone engineering. J. Dent. Res. 93, 1187 1195. Thaller, S.R., Hoyt, J., Borjeson, K., Dart, A., Tesluk, H., 1993. Reconstruction of calvarial defects with anorganic bovine bone mineral (Bio-Oss) in a rabbit model. J. Craniofac. Surg. 4, 79 84. Urist, M.R., 1980. Bone transplants and implants. In: Mr, U. (Ed.), Fundamental and Clinical Bone Physiology. JB Lippincott, Philadelphia, PA. Valentini, P., Abensur, D., 1997. Maxillary sinus floor elevation for implant placement with demineralized freeze-dried bone and bovine bone (Bio-Oss): a clinical study of 20 patients. Int. J. Periodontics Restorative Dent. 17, 32 45. Van Der Giessen, W.J., Lincoff, A.M., Schwartz, R.S., Van Beusekom, H.M., Serruys, P.W., Holmes, D.R., et al., 1996. Marked inflammatory sequelae to implantation of biodegradable and nonbiodegradable polymers in porcine coronary arteries. Circulation 94, 1690 1697. Wallace, S.S., Froum, S.J., 2003. Effect of maxillary sinus augmentation on the survival of endosseous dental implants. A systematic review. Ann. Periodontol. 8, 328 343. Wallace, S., Gellin, R., 2010. Clinical evaluation of freeze-dried cancellous block allografts for ridge augmentation and implant placement in the maxilla. Implant. Dent. 19, 272 279. Wallace, S.S., Froum, S.J., Cho, S.-C., Elian, N., Monteiro, D., Kim, B.S., et al., 2005. Sinus augmentation utilizing anorganic bovine bone (Bio-Oss) with absorbable and nonabsorbable membranes placed over the lateral window: histomorphometric and clinical analyses. Int. J. Periodontics Restorative Dent. 25, 450 465. Wang, F.M., Qiu, K., Hu, T., Wan, C.X., Zhou, X.D., Gutmann, J.L., 2006. Biodegradable porous calcium polyphosphate scaffolds for the three-dimensional culture of dental pulp cells. Int. Endod. J. 39, 477 483. Wang, H., Li, Y., Zuo, Y., Li, J., Ma, S., Cheng, L., 2007. Biocompatibility and osteogenesis of biomimetic nano-hydroxyapatite/polyamide composite scaffolds for bone tissue engineering. Biomaterials 28, 3338 3348. Weber, J., White, E.W., 1973. Carbonate materials as precursors of new ceramic, metal, and polymer materials for biomedical applications. Miner. Sci. Eng. 5, 151 165. Weber, J.N., White, E.W., Lebiedzik, J., 1971. New porous biomaterials by replication of echinoderm skeletal microstructures. Nature 233, 337 339.
376 Advanced Dental Biomaterials White, E., Shors, E., 1986. Biomaterial aspects of Interpore-200 porous hydroxyapatite. Dent. Clin. North Am. 30, 49 67. Wilk, R., 2004. Bony reconstruction of the jaws. In: Michael, M. (Ed.), Peterson’s Principles of Oral and Maxillofacial Surgery, second ed. B C Decker, Hamilton (London). Witte, F., Hort, N., Vogt, C., Cohen, S., Kainer, K.U., Willumeit, R., et al., 2008. Degradable biomaterials based on magnesium corrosion. Curr. Opin. Solid State Mater. Sci. 12, 63 72. Wong, R.W., Rabie, A.B., 2006. Effect of Gusuibu graft on bone formation. J. Oral Maxillofac. Surg. 64, 770 777. Xia, Z., Triffitt, J.T., 2006. A review on macrophage responses to biomaterials. Biomed. Mater. 1, R1. Xynos, I.D., Edgar, A.J., Buttery, L.D., Hench, L.L., Polak, J.M., 2001. Gene-expression profiling of human osteoblasts following treatment with the ionic products of Bioglasss 45S5 dissolution. J. Biomed. Mater. Res. A 55, 151 157. Yildirim, M., Spiekermann, H., Biesterfeld, S., Edelhoff, D., 2000. Maxillary sinus augmentation using xenogenic bone substitute material Bio-Oss in combination with venous blood. A histologic and histomorphometric study in humans. Clin. Oral Implants Res. 11, 217 229. Younger, E.M., Chapman, M.W., 1989. Morbidity at bone graft donor sites. J. Orthop. Trauma 3, 192 195. Yukna, R.A., 1990. HTR polymer grafts in human periodontal osseous defects. I. 6-month clinical results. J. Periodontol. 61, 633 642. Yukna, R.A., 1994. Clinical evaluation of HTR polymer bone replacement grafts in human mandibular Class II molar furcations. J. Periodontol. 65, 342 349. Yukna, R.A., Greer JR., R.O., 1992. Human gingival tissue response to HTR polymer. J. Biomed. Mater. Res. 26, 517 527. Zitzmann, N.U., Naef, R., Scharer, P., 1997. Resorbable versus nonresorbable membranes in combination with Bio-Oss for guided bone regeneration. Int. J. Oral Maxillofac. Implants 12, 844 852.
16 Calcium orthophosphates as a dental regenerative material Sergey V. Dorozhkin Moscow, Russia Chapter Outline 16.1 16.2 16.3 16.4 Introduction 377 General definitions and knowledge 383 Brief information on current biomedical applications of CaPO4 CaPO4 for dental caries prevention and in dentifrices 385 16.4.1 16.4.2 16.4.3 16.4.4 385 Toothpastes 386 Chewing gums 389 Teeth remineralization 390 Dentin hypersensitivity treatments 392 16.5 Clinical applications of CaPO4 in dentistry 394 16.5.1 Classification according to the existing CaPO4 395 16.5.2 Classification according to the dental specialties 404 16.6 Tissue engineering approaches 16.7 Conclusion 416 References 416 Further reading 452 16.1 414 Introduction Dental caries, also known as tooth decay or a cavity, is an infectious disease (usually of bacterial origin), which causes demineralization and destruction of teeth. If left untreated, the disease can lead to pain, tooth loss, and infection. Historically, this disease is very old and it is not exclusive to the human species. Namely, evidence of dental lesions compatible with caries has been observed in creatures as old as Paleozoic fishes (570250 million years ago), Mesozoic herbivores dinosaurs (24565 million years ago), prehominines of the Eocene (6025 million years ago), as well as in the Miocene (255 million years ago), Pliocene (51.6 million years ago), and Pleistocene animals (1.6 million10,000 years ago). Nowadays caries is also detected in bears and other wild animals, as well as being common in domestic animals (Lanfranco and Eggers, 2012). Back to humans, dental caries has been detected in various epochs and societies throughout the world (Wang et al., 2008; Bellagarda, 1965; Fujita, 2009; Corbett and Moore, 1971, 1973, 1975, 1976; Kerr, 1990). Even though in most of the developed countries the prevalence of the Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00016-5 Copyright © 2019 Elsevier Ltd. All rights reserved.
378 Advanced Dental Biomaterials disease has decreased since the introduction of fluoride for dental care, dental caries remains a major public health problem. Very schematically, dental caries occurs as follows. As the most highly mineralized structure in vertebrate bodies, dental enamel is composed of numerous nanodimensional needle-like crystals of ion-substituted calcium orthophosphates (CaPO4) with the apatitic structure (the so-called biological apatite), which are bundled in parallel ordered prisms or rods to ensure their unique mechanical strength, remarkable hardness, and biological protection. Nevertheless, all types of teeth possess some porosity allowing fluids beneath their surface. Organic (mainly, lactic and acetic) acids, produced by dental plaque cariogenic bacteria (such as Streptococcus mutans and Lactobacillus) from fermentable carbohydrates of sugar or from the remaining food debris, initiate the disease. When the sufficient quantity of acids is produced, so that the solution pH drops below B5.5 (a critical pH), saliva and plaque fluids cease to be saturated with calcium and orthophosphate ions. Thus dental enamel begins to be demineralized (dissolved) and the aforementioned pores become larger (Fig. 16.1A). The demineralization process can be described with a simplified chemical reaction: Ca10 ðPO4 Þ6 ðOHÞ2 1 H1 ! Ca21 1 HPO22 4 1 H2 O As seen from this reaction, enamel dissolution increases concentrations of both the major ions (calcium and acid orthophosphate) and the minor ones [magnesium, bicarbonate (not shown)] in the local microenvironment of the caries lesions, leading to the formation of various types of acidic CaPO4 (LeGeros, 1999; González-Cabezas, 2010; West and Joiner, 2014). Simultaneously, H1 ions are consumed which results in the pH increasing. Due to both the accumulation of the aforementioned ions in saliva and the increasing pH, the demineralization processes of teeth slow down. Several models were developed to simulate dental caries (Ilie et al., 2012; Fabregas and Rubinstein, 2014a,b; Yu et al., 2017). Figure 16.1 Schematic drawings of (A) demineralization and (B) remineralization processes of dental enamel.
Calcium orthophosphates as a dental regenerative material 379 Luckily, saliva has some restorative functions, acting not only as a buffer, to reduce the acidity caused by plaque bacteria, but also as the constant source of soluble ions of calcium and orthophosphate (González-Cabezas, 2010; Dowd, 1999). Therefore upon neutralization of the plaque acids, CaPO4 complexes from saliva diffuse back into the channels between the depleted enamel rods, replenishing the supply of the dissolved ions (Fig. 16.1B). Consequently, the surface of dental tissues is remineralized. Additional application of toothpastes, mouthwashes, mouth rinses, tooth mousses, etc. assists the remineralization. Thus under normal circumstances, enamel demineralization is compensated for its remineralization. This dynamic process takes place more or less continually and equally in a favorable oral environment. However, when the demineralization exceeds the combined abilities of saliva, toothpastes, mouthwashes, mouth rinses, tooth mousses, etc. to remineralize, the dental tissues are progressively dissolved and finally break down, producing dental caries, which look like cavities and/or holes in the teeth (Fejerskov and Kidd, 2008). An example of cariogenic teeth is shown in Fig. 16.2. In addition to dental caries, there is dental erosion, which is a chemical wear of the dental hard tissues without the involvement of bacteria resulting in bulk mineral loss with a partly demineralized surface of reduced microhardness. Clinical features are a loss of surface structures with shallow lesions on smooth surfaces and cupping and flattening of cusps; already in early stages, coronal dentine often is exposed. Frequently, it is caused by acid-containing drinks and/or food. The acids that cause erosion are rather strong with an average pH of B2 for the colas, B22.5 for citrus fruits, and B1 for gastric contents. A repeated exposure leads to surface demineralization and, therefore, softening, while the softened surface is susceptible to loss by abrasion from food or a toothbrush. Repeated cycles of acid exposure lead to smooth, cupped out cavities. Palatal surfaces of maxillary anterior teeth appear to be the most susceptible erosion places, although other teeth are also affected. Currently, dental erosion is considered as one of the main tooth pathologies, Figure 16.2 Dental caries (teeth decay).
380 Advanced Dental Biomaterials which is able to cause patient discomfort after periodontal diseases and caries (West and Joiner, 2014; Eccles, 1979; Lussi, 2006). Filling with artificial materials is a conventional treatment to repair damaged places. However, secondary caries frequently arise at the interfaces between the tooth and foreign materials, which always require restoration replacement (Mjör and Toffenetti, 2000). Furthermore, there are other reasons why people need restorative dental biomaterials, such as inlays, onlays, crowns, veneers, or bridges. The causes may be of a medical (decay or periodontal disease), age (population aging), traumatic (accident), or genetic (disorders) nature. All these causes adversely affect masticator efficiency, language function, facial esthetics, and even the psychological health. Still other patients simply wish to change their smile to improve an appearance of their own. Since no one wants to cover up one’s mouth when smiling, the demand for esthetic, tooth-colored (invisible) restorations permanently increases (Höland et al., 2008). Finally, there are dental abrasion and dental attrition processes. The former is defined as a mechanical removal of hard tissues by the repeated introduction of foreign bodies into the oral cavity that are in contact with the teeth, while the latter is a physiological wearing out of dental hard tissues though tooth-to-tooth contact, without an intervention of foreign substances (West and Joiner, 2014). Therefore due to their visibility, the restorative dental biomaterials are fundamentally different from artificial implants for bone replacements (reviewed in Dorozhkin, 2015, 2016a). The greatest driving force to develop biomaterials for dental restoration is to fulfill the customers’ (patients, dentists, and dental technicians) needs. In addition to the esthetic requirements, pressures from the environmental regulations and public apprehension are on the verge of eliminating dental amalgam as a practical and inexpensive restorative filling material (Widström et al., 1992). Thus by the late 1990s, amalgam use in several European countries was phased out. Consequently, development of metal-free restorations with properties close to natural teeth (with respect to translucency, color, and abrasive behavior) or even better mechanical properties and better durability than natural teeth has been a great challenge for many years (Höland et al., 2008). Briefly, all restorative dental biomaterials must meet the following basic requirements (Höland et al., 2008): G G G G They must be durable and biocompatible. Their optical characteristics (gloss, translucency, and color, in particular) must be comparable to those of natural teeth. Their mechanical properties (strength and toughness) must meet the requirements of the indication range (namely, the required strength of an inlay is lower than that of a dental bridge). Their wear behavior must be similar to that of natural teeth. In addition, they should be easily implantable or injectable, which is a critical requirement for any medical application. Hence, a selecting problem of the appropriate biomaterials arises. When all material characteristics and clinical factors are considered, bioceramics offer, perhaps, the best choice for a metal-free dentistry. Namely, bioceramics possess excellent chemical
Calcium orthophosphates as a dental regenerative material 381 durability, wear resistance, biocompatibility, environmental friendliness, and esthetics. The bioceramic restorations can be used in situations such as treatments of primary caries where inlays can be applied without a more excessive removal of tooth structure that is associated with amalgam. Besides, bioceramic onlays or crowns can also be used in place of large amalgam restorations. However, the widespread use of all-ceramic restorations has been hindered by concerns related to marginal fracture resistance and clinical longevity. Therefore the goal of dental bioceramics research is to produce all-ceramic dental restorative systems that utilize the known advantages of ceramic materials and minimize the existing disadvantages (Thompson et al., 2007). Due to the chemical composition, CaPO4 belong to bioceramics but they have some specific advantages due to a chemical similarity to the inorganic part of both human and mammalian bones and teeth. This makes CaPO4 the only bioceramics potentially applicable for remineralization of dental surfaces (Reynolds, 2008). Owing to these known similarities, dentists have been using CaPO4 in clinical practice for over a century. Namely, Dr. Junius E. Cravens (18441920) from the United States proffered creative concepts in pulp capping in the 1870s. He had the opinion that dentin-like material would be the best to keep the pulp vital. Therefore Cravens used a CaPO4 powder, which was mixed with lactic acid to low viscosity. The result was a soluble calcium lactic orthophosphate, which was applied to the exposed pulp tissue (Cravens, 1876). This pulp capping agent was brought to the market by the S.S. White Company with the trade name “Lacto-Phosphate of Lime” (Fig. 16.3) (Pendleton, 1873). The available CaPO4, their standard abbreviations, and solubility values are listed in Table 16.1 (Dorozhkin, 2012a,b). Additional details on CaPO4, their properties, and applications are available in the special monographs on the subject (LeGeros, 1991; Elliott, 1994; Dorozhkin, 2016c). The objective of this overview is to provide the current state-of-the-art on CaPO4 applications in dentistry and dentistry-relevant fields. Figure 16.3 An advertisement of the S.S. White company for “Lacto-Phosphate of Lime” 1873. Source: Reprinted from Dent. Cosmos 1873, 15, 683.
Table 16.1 Existing calcium orthophosphates (CaPO4) and their major properties (Dorozhkin, 2012, 2016). Ca/P molar ratio Compounds and their typical abbreviations Chemical formula Solubility at 25 C, 2 log (Ks) Solubility at 25 C, g/L 0.5 0.5 1.0 1.0 Ca(H2PO4)2 H2O Ca(H2PO4)2 CaHPO4 2H2O CaHPO4 1.14 1.14 6.59 6.90 B18 B17 B0.088 B0.048 1.33 1.5 1.5 1.22.2 MCPM MCPA or MCP DCPD, mineral brushite DCPA or DCP, mineral monetite OCP α-TCP β-TCP ACP 96.6 25.5 28.9 5.57.0 c B0.0081 B0.0025 B0.0005 c B512d 1.51.67 CDHA or Ca-def HAe B85 B0.0094 6.59.5 1.67 1.67 1.67 HA, Hap, or OHAp FA or FAp OA, OAp, or OXAf, mineral voelckerite TTCP or TetCP, mineral hilgenstockite Ca8(HPO4)2(PO4)4 5H2O α-Ca3(PO4)2 β-Ca3(PO4)2 CaxHy(PO4)z nH2O, n 5 34.5; 1520% H2O Ca102x(HPO4)x(PO4)62x(OH)22x (0 , x , 1) Ca10(PO4)6(OH)2 Ca10(PO4)6F2 Ca10(PO4)6O 116.8 120.0 B69 B0.0003 B0.0002 B0.087 9.512 712 Ca4(PO4)2O 3844 B0.0007 b 2.0 pH stability range in aqueous solutions at 25 C 0.02.0 a 2.06.0 a b b b ACP, Amorphous calcium phosphates; CDHA or Ca-def HA, calcium-deficient hydroxyapatite; DCPA or DCP, dicalcium phosphate anhydrous; DCPD, dicalcium phosphate dehydrate; FA or FAp, fluorapatite; HA, Hap, or OHAp, hydroxyapatite; MCPA or MCP, monocalcium phosphate anhydrous; MCPM, monocalcium phosphate monohydrate; OA, OAp or OXA, oxyapatite; OCP, octacalcium phosphate; PHA, precipitated HA; TTCP or TetCP, tetracalcium phosphate; α-TCP, α-tricalcium phosphate; β-TCP, β-tricalcium phosphate. a Stable at temperatures above 100 C. b These compounds cannot be precipitated from aqueous solutions. c Cannot be measured precisely. However, the following values were found: 25.7 6 0.1 (pH 5 7.40), 29.9 6 0.1 (pH 5 6.00), 32.7 6 0.1 (pH 5 5.28). The comparative extent of dissolution in acidic buffer is ACPcα-TCPcβ-TCP . CDHAcHA . FA. d Always metastable. e Occasionally, it is called “PHA.” f Existence of OA remains questionable.
Calcium orthophosphates as a dental regenerative material 16.2 383 General definitions and knowledge According to Wikipedia, the free encyclopedia: Dentistry is the branch of medicine that is involved in the study, diagnosis, prevention, and treatment of diseases, disorders and conditions of the oral cavity, maxillofacial area and the adjacent and associated structures and their impact on the human body. The American Dental Association recognizes nine dental specialties: public health dentistry, endodontics, oral and maxillofacial pathology, oral and maxillofacial radiology, oral and maxillofacial surgery, orthodontics, pediatric dentistry, periodontics, prosthodontics, and general dentistry. There are other dental niches such as oral medicine, dental aesthetics, dental implantation, and orofacial pain and temporomandibular disorders, some of them are recognized as dental specialties in other countries. In the European Union, all member states must recognize the specialties of orthodontics and oral and maxillofacial surgery (Dentistry). Now it is necessary to describe briefly all dental specialties and determine in which of them CaPO4 is used. According to Wikipedia, “Dental public health is involved in the assessment of dental health needs and improving the dental health of populations rather than individuals. One of the controversial subjects relating to dental public health is fluoridation of drinking water” (Dental Public Health). A search in Scopus database has been performed for papers containing in the title a combination of terms (keywords) “public health dentistry” 1 “apatite” and “public health dentistry” 1 “calcium phosphate.” Zero publications have been found for either case (Table 16.2). Thus this direction has nothing in common with CaPO4. (odous) “tooth”] deals Endodontics [from the Greek ε νδo (endo) “inside” and o δoυς with the tooth pulp and tissues surrounding roots of teeth. If the pulp (this is the only vascularized connective tissue of teeth, which contains nerves, arterioles, venules, lymphatic tissue, and fibrous tissue) becomes diseased or injured, endodontic treatment is required to save the tooth (Endodontics). The results of a similar search (Table 16.2) revealed that CaPO4 is used rarely in endodontics. Oral and maxillofacial pathology, radiology, and surgery represent “the study, diagnosis, and sometimes the treatment of oral and maxillofacial related diseases,” “the study and radiologic interpretation of oral and maxillofacial diseases,” and “extractions, implants, and facial surgery,” respectively (Specialty dentistry). Only surgery appears to deal with CaPO4 very rarely (Table 16.2). Furthermore, within these three dental specialties, one needs to differentiate between “oral” and “maxillofacial” terms. The former term is relevant to the subject of this review, while the latter is undoubtedly irrelevant, since it deals with the treatment of the surrounding bones. Various CaPO4-based formulations have been proposed for the reconstruction of the contour and discontinuity defects in maxillofacial surgery (Friedman et al., 1991, 2000; Reddi et al., 1999; Smartt et al., 2005; Kuemmerle et al., 2005; Luaces-Rey et al., 2009; Tamimi et al., 2009; Abe et al., 2009; Benson and Djalilian, 2009; Lee et al., 2010a; Singh et al., 2010; Bambakidis et al., 2010;
384 Advanced Dental Biomaterials Table 16.2 The amount of publications containing the selected keywords in their titles, found in Scopus database. Number of publications Dental specialty Apatite Calcium phosphate Cumulative Public health dentistry Endodontics Oral and maxillofacial pathology Oral and maxillofacial radiology Oral and maxillofacial surgery Orthodontics Pediatric dentistry Periodontics Prosthodontics 0 23 0 0 5 14 0 219 9 0 14 0 0 2 36 0 118 0 0 37 0 0 7 50 0 337 9 427 25 80 5 39 846 249 19 95 2 96 631 676 44 175 7 135 1477 Additional keywords Dental Dentistry Oral Stomatology Caries Totala a Duplications are possible. Huang et al., 2010a; Sanada et al., 2011); however, this subject was contained within the field of bone grafts (Dorozhkin, 2015, 2016a). Orthodontics, formerly orthodontia [from Greek o ρθoς (orthos) “straight, or proper, or perfect” and o δoυς (odous) “tooth”], is the first specialty of dentistry that is concerned with the study and treatment of malocclusions (improper bites), which may be a result of tooth irregularity, disproportionate jaw relationships, or both (Orthodontics). CaPO4 is used rarely in orthodontics (Table 16.2). Pediatric dentistry [formerly pedodontics (American English) or paedodontics (Commonwealth English)] is the branch of dentistry dealing with children from birth through adolescence. It places special importance in preventing tooth decay. In addition, pediatric dentists work toward the maintenance of primary teeth (baby teeth) until they are naturally lost. It is irrelevant to the CaPO4 subject (Table 16.2). Periodontics [also periodontology, from (odous) “tooth”] is the specialty of dentistry Greek περι (peri) “around” and o δoυς that studies supporting structures of teeth, as well as diseases of periodontium (these are specialized tissues investing and supporting teeth, including cementum, periodontal ligament, alveolar bone, and gingiva, characterized by the loss of support around teeth) and conditions that affect them. Although CaPO4 is frequently used in periodontics (Table 16.2); in fact, they are applied to treat alveolar bones, which, again, is another story (Dorozhkin, 2015, 2016a). Prosthodontics [from Greek (odous) “tooth”], also known as dental πρoσθεση (prosthesis) “addition” and o δoυς
Calcium orthophosphates as a dental regenerative material 385 prosthetics or prosthetic dentistry, is a dental specialty pertaining to the diagnosis, treatment planning, rehabilitation, and maintenance of the oral function, comfort, appearance, and health of patients with clinical conditions associated with missing or deficient teeth and/or oral and maxillofacial tissues using biocompatible substitutes (Prosthodontics). CaPO4 is used very rarely in prosthodontics (Table 16.2). In addition, similar searches in Scopus database using keywords “dental,” “dentistry,” “oral,” “stomatology,” and “caries” combined with “apatite” or “calcium phosphate” have been performed (Table 16.2, the bottom lines). 16.3 Brief information on current biomedical applications of CaPO4 Due to a chemical similarity to the inorganic part of normal calcified tissues (bones, teeth, and deer antlers) of mammals, artificially prepared CaPO4 possess good biocompatibility, bioactivity, and osteoconductivity (Dorozhkin, 2012a, 2016b,c; LeGeros, 1991; Elliott, 1994). These properties of CaPO4 are extensively used in medicine for repair or replacement of injured or damaged bones and teeth. Since the diverse biomedical applications require different formulations, configurations, and/or shapes, the biomedically relevant CaPO4 is produced in various physical forms, such as powders, particles, granules, dense blocks, porous scaffolds, selfsetting formulations, suspensions, nonhardening pastes, implant coatings, as well as composite components of different origin (natural, biological, or synthetic) often with the specific shapes, such as implants, prostheses, or prosthetic devices (Dorozhkin, 2012a, 2015, 2016a,b,c; LeGeros, 1991; Elliott, 1994; Xie et al., 2012; Fernández et al., 2017). In view of the fact that several dental specialties deal with an invasion into and/or treatment of the surrounding bones (such as bone drilling to insert an implant), in principle, all the aforementioned forms, formulations, configurations, and shapes of CaPO4 might be applicable to the dentistry field as well. 16.4 CaPO4 for dental caries prevention and in dentifrices Traditionally, caries prevention strategies have focused on reducing bacterial growth, neutralizing oral acids, and teeth remineralization. Among them, only the third strategy appears to deal with the CaPO4 subject. Briefly, the teeth remineralization is a process in which dissolved CaPO4 minerals are returned to the molecular structure of the teeth themselves (Fig. 16.1B). To reduce dental caries by performing remineralization, systemic and/or topical fluoridation of water is commonly used (Reynolds, 2008). In addition, various ion-delivering agents are used in the form of dentifrices, toothpastes, mouthwashes, mouth rinses, chewing gums, etc. Many of these remineralizing agents contain CaPO4 (Roveri et al., 2009a). This is because the focus in caries research has shifted to the development of methodologies for detection of the early stages of caries lesions and noninvasive treatment
386 Advanced Dental Biomaterials of these lesions. For example, in the presence of calcium and orthophosphate ions, topical fluoride ions promote the formation of fluorapatite (FA) (which is the least soluble compound among all known types of CaPO4, Table 16.1) in dental enamel. This property of fluorides has been known since, at least, 1956 (Fischer et al., 1956). However, to form one unit cell of FA, 10 calcium and 6 orthophosphate ions are required for every 2 fluoride ions. Hence, on topical application of fluoride ions, the availability of calcium and orthophosphate ions can be the limiting factor for net enamel remineralization to occur, and this is highly exacerbated under the xerostomic (i.e., a dry mouth) conditions (Reynolds, 2008). Now, let me describe the applications of CaPO4 in dentifrices. According to Wikipedia, “Dentifrice are agents used along with toothbrush to clean and polish natural teeth. They are supplied as paste, powder, gel, or liquid form” (Dentifrice). To the best of my findings, the earliest study dealing with an application of CaPO4 as a dentifrice was presented at the 23rd general meeting of the International Association for Dental Research (held in Chicago, Illinois, May 27, 1945) and the abstract of that presentation was published shortly afterward (McClendon and Carpousis, 1945). Since then, numerous studies devoted to various applications of CaPO4 in dentifrices have been published (Shern et al., 1979, 1984; Wefel and Harless, 1987; Schreiber et al., 1988; Kani et al., 1988, 1989; Okashi et al., 1991; Gaffar et al., 1993; Zhang et al., 1995; Sullivan et al., 1997, 2001; Kodaka et al., 1999; Hicks and Flaitz, 2000; Boneta et al., 2001; Silva et al., 2001; Niwa et al., 2001; Sintes et al., 2002; Kim et al., 2006, 2009; Jeong et al., 2006, 2007; Lv et al., 2007; Roveri et al., 2009b; Tschoppe et al., 2011; Najibfard et al., 2011; Wang et al., 2011a; Kovtun et al., 2012; Vanichvatana and Auychai, 2013; Sun et al., 2014; Browning et al., 2012; Comar et al., 2013; Hannig et al., 2013; Kensche et al., 2017; de Carvalho et al., 2014; Mielczarek and Michalik, 2014; Vyavhare et al., 2015; Souza et al., 2015; Hill et al., 2015; Gomes et al., 2017; Nakata et al., 2018; Hiller et al., 2018). A number of such formulations also contain fluorides (Shern et al., 1979, 1984; Gaffar et al., 1993; Zhang et al., 1995; Hicks and Flaitz, 2000; Sullivan et al., 2001; Boneta et al., 2001; Silva et al., 2001; Sintes et al., 2002; Roveri et al., 2009b; Wang et al., 2011a; Vyavhare et al., 2015; Souza et al., 2015; Hill et al., 2015; Gomes et al., 2017; Nakata et al., 2018; Hiller et al., 2018). 16.4.1 Toothpastes CaPO4-containing toothpastes were found to promote a partial remineralization of the demineralized enamel (Jeong et al., 2006; Lv et al., 2007; Roveri et al., 2009b; Kim et al., 2009; Tschoppe et al., 2011; Wang et al., 2011a; Kovtun et al., 2012; Vanichvatana and Auychai, 2013; Sun et al., 2014; de Carvalho et al., 2014; Mielczarek and Michalik, 2014; Vyavhare et al., 2015; Souza et al., 2015); as well as depending on the addition of other constituents, they also can possess some whitening effect (Niwa et al., 2001; Kim et al., 2006; Hill et al., 2015) and reduce tooth sensitivity (Kim et al., 2009; Browning et al., 2012; Hiller et al., 2018). For example, the polishing and whitening properties of hydroxyapatite (HA)-containing
Calcium orthophosphates as a dental regenerative material 387 toothpastes were investigated in a combined study (Niwa et al., 2001). The polishing properties were evaluated by means of artificial teeth by polishing with different toothpastes, while the brightening and whitening properties were examined in volunteers using two colorimeters with two specially made fiberscopes. The results revealed that the addition of HA to the toothpaste did not alter its polishing properties, while it did result in a marked increase in tooth whitening. It was also found that the brightening and whitening properties increased as the amount of HA in the toothpaste increased. Thus HA-containing toothpaste appeared to be effective at whitening teeth and whitening was not due to their polishing effect on the tooth surface (Niwa et al., 2001). The whitening properties of HA-containing toothpastes were also found by other researchers (Kim et al., 2006; Hill et al., 2015). Further, a randomized study with 181 children (92 boys, 89 girls) from different Japanese schools over a period of 3 years is worth mentioning (Kani et al., 1989). After lunch, the children brushed their teeth under supervision either with a toothpaste containing 5% HA or a control group with a paste without HA. Yearly controls of the DMFT (number of decayed, missing, and filled teeth due to caries) index were diagnosed as well as the caries incidence on newly erupted teeth. The DMFT index appeared to be significantly deeper in the HA-containing toothpaste group, while the incidence for caries in newly erupted teeth was significantly lower if compared to the control (Kani et al., 1989). Besides, dentifrices containing a combination of monofluorophosphate (MFP) with a dicalcium phosphate dehydrate (DCPD) abrasive were evaluated in a variety of in vivo tests (Gaffar et al., 1993). MFP with silicon dioxide abrasive at the equivalent fluoride concentration was used for comparison. The data indicated that DCPD was more effective than silica in preventing pH drop of plaque. A toothpaste containing MFP 1 DCPD was significantly more effective than an MFP 1 silica toothpaste. In addition, a toothpaste containing 45Ca radiolabeled DCPD was applied topically in rats’ teeth. The results showed that 45Ca was incorporated into the enamel with a concomitant reduction in enamel solubility. In a rat caries study using MFP 1 DCPD, matching placebo, and MFP 1 silica, the MFP 1 DCPD dentifrice showed a significantly greater reduction in smooth surface caries. These dentifrices were also tested in an in situ human model for fluoride uptake in artificial root caries lesions where MFP 1 DCPD provided a significantly higher fluoride uptake than MFP 1 silica. A second in situ study in humans evaluated the same dentifrices, MFP 1 DCPD increased salivary plaque calcium and fluoride. These results of laboratory, animal, and in situ studies taken together indicated that the MFP 1 DCPD combination was the unique one in providing extra supersaturation in saliva and plaque with concomitant enhanced anticaries efficacy (Gaffar et al., 1993). Thus due to the aforementioned successful cases of CaPO4 addition to toothpastes, such toothpastes are commercially produced worldwide (Table 16.3). As seen from Table 16.3, toothpastes for both human and animals are available. One should note that HA and amorphous calcium phosphates (ACPs) are added to toothpastes to provide remineralization properties, while DCPD and dicalcium phosphate anhydrous (DCPA) are added to toothpastes as abrasives to provide a gentle polishing action.
Table 16.3 Trademarks of CaPO4-containing commercial toothpastes. Type of CaPO4 Human or animals Trade name and producer (when available) HA Human Aclaim (Group Pharmaceuticals Limited, India) Active Remineralization Toothpaste (A.R.T.) (Pearlie White, Corlison, Singapore) ApaCare (Cumdente, Germany) Apadent (Sangi Co., Japan) Apagard Premio (Sangi Co., Japan) Arcticum (SPLAT-COSMETICA, Russia) Biorepair (Coswell, Italy) Biorepair (Dr. Wolff, Bielefeld, Germany) Coolin Bubble (Canavena Co., Korea) DIO (DIO Co., Korea) Desensibilize Nano P (FGM Produtos Odontológicos, Brazil) Desensin repair (Dentaid) Hakusanshiko (Japan) Janina (Janina Ultra White, United Kingdom) Kalident—calcium hydroxyapatite (Kalichem, Italia) MAXDENT (STS Cosmetics, Bulgaria) Megasonex (Goldspire Group, Hong Kong) nanoXIM CarePaste (FLUIDINOVA, Portugal) Parodontol Active (Svoboda Ltd., Russia) PrevDent (PrevDent International, The Netherlands) Renamel AfterBleach (Sangi Co., Japan) Remin (X-PUR, Oral Science, QC, Canada) R.O.C.S. SENSITIVE (DRC Group, Russia) Sensitive Reminx (Pharma Jenistec Co., Korea) Triple Denta (TripleLife Co., Ltd., Korea) Ultracomplex (SPLAT-COSMETICA, Russia) UltraDEX Recalcifying and Whitening (Periproducts Ltd., United Kingdom) VITIS anticaries toothpaste (Dentaid) YP Dental (You Co., Ltd. Japan) Clinpro Tooth Crème (3M ESPE, United States) Age Defying (Arm & Hammer, Church & Dwight Co. New Jersey, United States) Complete Care (Arm & Hammer, Church & Dwight Co. New Jersey, United States) EnaFix (Group Pharmaceuticals Limited, India) Enamel Care (Arm & Hammer, Church & Dwight Co. New Jersey, United States) Enamel Pro (Premier Dental Products Company, United States) Enamelon (Premier Dental Products Company, United States) GC Tooth Mousse (GC Corporation, Japan) INNOVA (SPLAT-COSMETICA, Russia) MI paste (GC America, Illinois, United States) MI paste plus (GC America, Illinois, United States) G TCP ACP Human Human (Continued)
Calcium orthophosphates as a dental regenerative material 389 Table 16.3 (Continued) Type of CaPO4 Human or animals Trade name and producer (when available) DCPD or DCPA Human All White (Dr. Collins, United States) Colgate Cavity Protection (Colgate-Palmolive, New York, United States) Dentu-Creme Denture (Polident, GlaxoSmithKline, United Kingdom) Plus White (CCA Industries Inc., New Jersey, United States) Pureen (Singapore) R.O.C.S. PRO Baby (DRC Group, Russia) Snappy Jaws (Australia) Supersmile (United States) Triple Action Whitening (Pearl Drops, Church & Dwight, New Jersey, United States) Triple Power Whitening (Pearl Drops, Church & Dwight, New Jersey, United States) VITA-MYR (Nevada, United States) Advanced Oral Care (Nylabone, New Jersey, United States) C.E.T. Enzymatic (Virbac, Texas, United States) Dental Care Kit (Sentry Petrodex, Sergeant’s Pet Care Products, Nebraska, United States) Dentifresh (Hatchwell, United Kingdom) Enzymatic toothpaste (Sentry Petrodex, Sergeant’s Pet Care Products, Nebraska, United States) Four Paws Pet Dental (Four Paws Products, New York, United States) VetOne (VetOne, Idaho, United States) TriMedica Pure MSM (TriMedica, Northern Ireland, United Kingdom) Animals Undisclosed CaPO4 Human ACP, Amorphous calcium phosphates; DCPA, dicalcium phosphate anhydrous; DCPD, dicalcium phosphate dehydrate; HA, hydroxyapatite; TCP, tricalcium phosphate. To finalize this section, one should mention one of the studies in which the addition of CaPO4 to toothpastes did not show any positive influence on enamel and/or dentin demineralization/remineralization properties (Esteves-Oliveira et al., 2017). 16.4.2 Chewing gums Apart from toothpastes, CaPO4 is added to chewing gums to reduce dental caries (Pickel and Bilotti, 1965; Finn and Jamison, 1967; Finn, 1967; Richardson et al., 1972; Wilson, 1975; Chow et al., 1994; Vogel et al., 1998, 2000; Shen et al., 2001; Iijima et al., 2004; Itthagarun et al., 2005; Cai et al., 2007; Morgan et al., 2008; Thaweboon et al., 2009; Dodds et al., 2012; Shammukha et al., 2012; Porciani et al., 2014; Emamieh et al., 2015; Sultan et al., 2016; Hasnamudhia et al., 2017).
390 Advanced Dental Biomaterials In the vast majority of cases, a positive effect was noticed. Namely, to evaluate chewing gums as a vehicle to increase salivary mineral saturation levels and enhance salivation, both monocalcium phosphate monohydrate (MCPM) and the equimolar mixture of tetracalcium phosphate (TTCP) with DCPA were chosen as experimental chewing gum additives (Chow et al., 1994). Each subject chewed a commercial sugar-free bubble gum (control) for 16 minutes or the same gum to which 5 wt.% of MCPM or TTCP 1 DCPA mixture had been added. Both experimental gums were found to increase significantly the concentrations of calcium and orthophosphate ions in saliva during the 16-minute period even more than with a previously evaluated gum that contained DCPD. The degree of saturation of tooth mineral was significantly increased by both experimental gums, with a greater increase being produced by the TTCP 1 DCPA gum. The MCPM gum produced a significantly greater saliva flow and a lower salivary pH than did the control and TTCP 1 DCPA gums. The results suggested that the experimental gums could be useful for promoting remineralization in general and for inducing salivation in xerostomic patients (Chow et al., 1994). In other studies, both sugar-free gums (control) and casein phosphopeptideACP (CPPACP) containing gums were chewed for either 20-minute periods, four times a day or 5-minute periods, seven times a day. Microradiography and computer-assisted densitometric image analysis demonstrated that regardless the gum type and chewing duration (e.g., 20 or 5 minutes), the nanodimensional CPPACP complexes produced a dose-related remineralization of enamel subsurface lesions in situ. The gums containing 18.8 and 56.4 mg of the nanodimensional CPPACP complexes, chewed for 20 minutes, four times per day for 14 days, increased enamel subsurface remineralization by 101% and 151%, respectively, relative to the control sugar-free gums. Microradiographs of the enamel lesions before and after remineralization showed that the nanodimensional CPPACP complexes promoted remineralization throughout the body of the lesion. Electron microprobe wavelength dispersive spectrometric analyses of sections of the remineralized enamel indicated that the mineral deposited was apatite with a higher Ca/P ratio than that in the stoichiometric HA. Acid challenge of the enamel remineralized by the nanodimensional CPPACP complexes in situ showed that the remineralized apatite was more resistant to acid challenge than the normal calcium-deficient carbonated tooth enamel. Thus the clinical trials of CPPACP-containing sugar-free chewing gums demonstrated that these gums significantly slowed progression of caries and enhanced regression of caries compared with the control sugar-free gums (Shen et al., 2001; Iijima et al., 2004; Itthagarun et al., 2005; Cai et al., 2007; Morgan et al., 2008). 16.4.3 Teeth remineralization In general, remineralization of teeth can be defined as the process in which calcium and orthophosphate ions are supplied from a source external to teeth to promote their deposition into crystal voids in demineralized enamel, to produce net mineral gain (Cochrane et al., 2010). The earliest found paper on a possibility of a remineralization phenomenon occurring in caries was published in 1912 (Head, 1912), while that using CaPO4 for rehardening was performed in 1961 with solutions
Calcium orthophosphates as a dental regenerative material 391 containing dissolved ions of calcium and orthophosphate (Koulourides et al., 1961), followed by the set of the studies performed by Silverstone (Silverstone and Johnson, 1971; Silverstone, 1972), as well as by 10 Cate and Arends (ten Cate and Arends, 1977, 1978, 1980; ten Cate et al., 1981). The early attempts to use CaPO4 for remineralization of dental surface were unsuccessful due to their low solubility, particularly in the presence of fluoride ions. Namely, the insoluble CaPO4 cannot be applied easily; they do not localize effectively on the tooth surface and require an acidic environment for solubility levels sufficient to produce ions capable of diffusing into enamel subsurface lesions. Furthermore, due to the intrinsic insolubility of CaPO4 at the physiological pH values, soluble calcium and orthophosphate ions can only be used at very low concentrations. Besides, the soluble ions of calcium and orthophosphate are neither substantially incorporated into the dental plaque nor localized at the tooth surface to produce the effective concentration gradients to drive diffusion into the subsurface enamel (Reynolds, 2008). Nevertheless, studies on teeth remineralization keep going. For example, a remineralization potential of HA itself for caries lesion treatment was investigated (Okashi et al., 1991). Previously demineralized enamel blocks were immersed into an aqueous solution of sludgy HA at 37 C for 55 hours, followed by 24 hours washing with synthetic saliva and another group was washed only with synthetic saliva. Artificial caries lesions were remineralized slightly by immersion into artificial saliva but a significant acceleration of remineralization was observed in the sludgy HA group (Okashi et al., 1991). Positive results were also obtained in other studies (Kani et al., 1988; Yamagishi et al., 2005; Kim et al., 2007; Lu et al., 2007; Zhen et al., 2007; Rimondini et al., 2007; Roveri et al., 2008; Huang et al., 2009, 2011; Yin et al., 2009; Lv et al., 2010; Wu et al., 2013; Haghgoo et al., 2014; Besinis et al., 2014; Kensche et al., 2016). In 2005 a study on the potential of HA in repairing enamel lesions was published in which an HA-containing paste demonstrated an enamel-repairing effect that was confirmed by transmission electron microscopy and atomic force microscopy (Yamagishi et al., 2005). The results of an in situ study showed a potential accumulation of HA from a mouthwash in the dental pellicle; however, no significant erosion protection was found from this accumulation (Kensche et al., 2016). Furthermore, remineralization of caries lesions could be performed by supersaturated solutions (Chow and Takagi, 1995; Reynolds, 1997; Wen et al., 2016) and/or gels (Wang et al., 2011b; Ning et al., 2012; Tian et al., 2012; Gu et al., 2013; Chen et al., 2013b; Cao et al., 2014; Wu et al., 2015; Wang et al., 2015a; Zaharia et al., 2017) containing dissolved ions of calcium and orthophosphate. In addition, supersaturated CaPO4 mouth rinses were found to experience a significant increase in reversals of caries in high risk for caries patients due to xerostomia (salivary hypofunction) (Singh and Papas, 2009). A remineralization potential of sports drinks, containing nano-sized HA particles, was also studied (Lee et al., 2007; Min et al., 2011). More complicated formulations, such as CaPO4-loaded liposomes combined with amelogenin-inspired oligopeptides (Luo et al., 2012), as well as biocomposites of nanodimensional ACP with various polymeric compounds (Liang et al., 2017a,b; Weir et al., 2017; Xiao et al., 2017) have been also developed to promote
392 Advanced Dental Biomaterials remineralization of dental enamel. Thus CaPO4 appear to be the chemicals able to reduce dental caries at the early stages. However, studies performed by using atomic force microscopy nanoindentation technique revealed that previously demineralized samples of dental enamel further exposed to remineralizing solutions did show a crystalline layer of CaPO4 formed on their surface. Unfortunately, the reprecipitated deposits of CaPO4 always consisted of loosely packed crystals and did not protect the underlying enamel from a subsequent acid attack. Furthermore, these surface deposits were completely removed by either a toothbrush or a short exposure to an erosive acidic solution (Jandt, 2006; Lippert et al., 2004a,b,c). In this context, it should be emphasized that the term “remineralization,” which is often misused in the literature, should imply the process of mineral growth that goes hand in hand with a strengthening effect of the weakened enamel surface. Since no strengthening of an exposure to remineralizing solutions was observed, it might be considered that no “passive mineralization” was found (in spite of the real evidence of the reprecipitated surface deposits of CaPO4) (Jandt, 2006; Lippert et al., 2004b,c). Further details on the remineralization attempts of teeth are available in the topical reviews (Li et al., 2014a; Niu et al., 2014; Cao et al., 2015). To finalize this section, one should mention a new technology called “electrically assisted enhanced remineralization” developed by Reminova Ltd. (Scotland, United Kingdom), which relies on the use of a tiny electrical current of a few micro-Amps to reverse tooth decay by boosting remineralization. Although the details are not disclosed, the science behind this process is described as follows: “It is a two-step process. First, a lesion in the enamel is identified by a dentist and conditioned with a quick and simple set of washes to remove decay and exogenous materials that may have built up within the decayed lesion. Some natural minerals are applied locally to the lesion and then a tiny electric field is used to help drive the minerals into the deepest parts of the now cleaned-out lesion. The minerals pack closely together and start precipitating (hardening) onto the existing enamel structure. Very quickly, the lesion will fill and the enamel is strong and healthy again. Because it is based on a natural process, it is understood how minerals (calcium and phosphate molecules) go back into tooth enamel, and how they join and form part of the enamel.” Obviously, either CaPO4 compounds or Ca- and PO4-containing constituents are used as “some natural minerals” in this technology. 16.4.4 Dentin hypersensitivity treatments As written in Wikipedia, the free encyclopedia Dentin hypersensitivity (abbreviated to DH or DHS and also termed sensitive dentin, dentin sensitivity, cervical sensitivity and/or cervical hypersensitivity) is dental pain which is sharp in character and of short duration, arising from exposed dentin surfaces in response to stimuli, typically thermal, evaporative, tactile, osmotic, chemical or electrical and which cannot be ascribed to any other dental disease (Dentin hypersensitivity).
Calcium orthophosphates as a dental regenerative material 393 Dentin hypersensitivity is a frequently reported oral pain condition, which is mostly diagnosed at the buccal surfaces of teeth, where enamel is missing due to erosion, abrasion, and/or attrition. Contrary to enamel, which is dense and contains a small amount of pores, dentine has a great number of tiny tubes (tubules) that lead to the nerve and are filled with fluids. However, until about the third or fourth decade of life in healthy individuals, the surface of dentin is not exposed and the tubules are sealed. When a tooth loses its protection from gum recession and/or tooth enamel wear, these tubules are exposed to the outside, allowing external stimuli to reach the nerve endings. Therefore even mild external stimuli such as hot or cold foods and beverages can cause a change in fluid movement, which causes the nerve endings to react in response, triggering a short but sharp pain. There are several strategies to prevent dentin hypersensitivity in oral care, including occlusion of dentin tubules with different particulate materials (arginine/ calcium carbonate, calcium sodium phosphosilicate, strontium fluoride, etc.) or nerve desensitization with agents such as potassium salts (Enax and Epple, 2018). However, all these cases are beyond the scope of this review. Back on the topic, some types of CaPO4 were found to be able to treat this disease as well (Kim et al., 2009; Browning et al., 2012; Porciani et al., 2014; Hüttemann and Dönges, 1987; Sugawara et al., 1989; Yates et al., 1998; Suge et al., 2002; Geiger et al., 2003; Shetty et al., 2010; Low et al., 2015; Vano et al., 2018; Mehta et al., 2015; Zhou et al., 2016). For example, already in 1987, the results of a clinical study of 140 patients with hypersensitive teeth, in which HA-containing pastes with different particle sizes of HA were used for treatment, were published. A paste containing 2-μm-sized HA particles appeared to be more effective than that containing 6-μm-sized HA particles and 90% of all patients experienced relief from tooth hypersensitivity after 35 days (Hüttemann and Dönges, 1987). In another study, an HA-containing toothpaste was compared with positive control toothpastes, in which the HA-containing toothpaste was similarly effective in reducing dentine hypersensitivity with preexisting benchmark toothpastes (Kim et al., 2009). Positive results were also obtained with both HA-containing Renamel AfterBleach toothpaste (Browning et al., 2012) and an undisclosed nano-HA, potassium nitrate, sodium MFP, and antioxidants-containing toothpaste (Low et al., 2015). In another study, HA-treated teeth showed statistically significant reduction in hypersensitive symptoms compared to the control groups and the authors concluded that HA showed “definite potential as an effective and permanent desensitizer when used as an in-office procedure” (Shetty et al., 2010). Furthermore, a CaPO4 precipitation method was once tried as a treatment for dentin hypersensitivity using the vital teeth of beagle dogs. The results revealed that dentin tubules were occluded homogeneously and completely with an apatitic mineral after application of the CaPO4 precipitation in vital teeth (Suge et al., 2002). A commercial self-setting formulation TEETHMATE DESENSITIZER (Kuraray Noritake Dental Inc., Japan), consisting of a mixture of DCPA 1 TTCP 1 some fluoride-free additives, which formed calciumdeficient HA (CDHA) precipitates upon exposure to saliva, appeared to able to occlude open dentinal tubules, and, by this way, acted as an effective desensitizer compound (Mehta et al., 2015; Zhou et al., 2016; Moogi et al., 2017; Shetty et al., 2017).
394 Advanced Dental Biomaterials Nevertheless, there are similar formulations with fluorides (Medvecky et al., 2018). In addition, there is commercially produced dental gel Relief ACP (Discus dental, United States), containing 0.375% ACP and 1100 ppm fluoride (Arafa, 2017), as well as MI varnish (3M ESPE, United States), containing ACP stabilized by a CPP and NaF (Sharma et al., 2017). Both of them are used to treat dentin hypersensitivity. Additional details on this topic are available in a recent review (Enax and Epple, 2018). 16.5 Clinical applications of CaPO4 in dentistry As said in Section 16.1, dentists have been using CaPO4 for over a century. However, to the best of my findings, the first available publication on decalcification of teeth as the reason for various dental pathologies was published in 1925 (Miloslavich, 1925). Furthermore, the clinical applications of CaPO4 in dentistry started only in the 1970s (Monroe et al., 1971). Namely, the first application of a CaPO4 [erroneously described as “tricalcium phosphate (TCP) of HA structure”] bioceramics in surgically created periodontal defects was reported in 1975 (Nery et al., 1975), followed by a publication on alveolar ridge augmentation in 1978 (Nery et al., 1978), while the use of dense HA cylinders for immediate tooth root replacement was reported in 1979 (Denissen and de Groot, 1979). A summary on early (before 1987) studies can be found in Table 16.3 of. LeGeros (1988), while Table 16.4 of this publication represents the various types of dental applications of CaPO4 in the mid-1980s (LeGeros, 1988, 1991). Overall, the reasons for the clinical application of CaPO4 in dentistry are similar to those for their applications in bone grafting. A chemical similarity to the inorganic phases of teeth and bones appears to be the major reason. Consequently, CaPO4 possess an excellent biocompatibility, biotolerance, an ability to be resorbed by both tooth- and bone-related cells, osteoconductivity, etc. In addition, CaPO4 is less expensive than the majority of inorganic fillers used today. In the following subsections, the clinical applications of CaPO4 in dentistry have been classified into two ways: according to the existing CaPO4, as listed in Table 16.1, and according to the modern dental specialties, as listed in Table 16.2. Table 16.4 Dental applications of CaPO4 in the mid-1980s (LeGeros, 1991; LeGeros, 1988). 1 2 3 4 5 6 7 Restore (augment) alveolar ridge for better denture fit Immediate tooth root replacement to prevent resorption of alveolar ridge Fillers for periodontal defects or bone loss Coatings for metal implants to improve boneimplant adhesion and prevent loosening of the metal implants Repair of cleft palate Repair of maxillofacial defects Pulp capping materials
Calcium orthophosphates as a dental regenerative material 395 16.5.1 Classification according to the existing CaPO4 16.5.1.1 Monocalcium phosphate monohydrate and monocalcium phosphate anhydrous Just a few studies on dental applications of MCPM and monocalcium phosphate anhydrous (MCPA) were found in databases of scientific literature (Table 16.1). According to the available publications, both compounds are used in dentistry as components of self-setting formulations (Kouassi et al., 2003; Mehdawi et al., 2009), including sealers (Koch et al., 2010). For example, a commercial product EndoSequence BC Sealer (Brasseler USA, Savannah, Georgia) is a premixed readyto-use injectable cement paste developed for permanent root canal filling and sealing applications. It contains zirconium oxide, calcium silicates, MCPA, calcium hydroxide, filler, and thickening agents. When this sealer is placed in the root canal, it absorbs water from the dentin tubules causing hydration reactions of calcium silicates. Simultaneously, MCPA reacts with calcium hydroxide to precipitate CDHA. This leads to the formation of a composite network of gel-like calcium silicate hydrates, which intimately mixes with CDHA crystals and forms a hermetic seal inside the root canal (Koch et al., 2010). In addition, MCPM and/or MCPA were tried as components of caries-inhibiting dental biocomposites, releasing ions of calcium and orthophosphate (Mehdawi et al., 2009; Xu et al., 2007a). Once a MCPMcontaining chewing gum was tested, it produced a significantly greater saliva flow and a lower salivary pH than the control gum did (Chow et al., 1994). 16.5.1.2 Dicalcium phosphate dehydrate and dicalcium phosphate anhydrous As seen in Table 16.3, DCPD and/or DCPA are often added to toothpastes as gentle polishing agents. In addition, DCPD and/or DCPA (unfortunately, the authors of the publications on the subject rarely specified which of them was used) are added to chewing gums (Pickel and Bilotti, 1965; Finn and Jamison, 1967; Finn, 1967; Richardson et al., 1972; Chow et al., 1994; Thaweboon et al., 2009) and other types of dentifrices (Shern et al., 1984; Wefel and Harless, 1987; Schreiber et al., 1988; Gaffar et al., 1993; Zhang et al., 1995; Sullivan et al., 1997, 2001; Kodaka et al., 1999; Hicks and Flaitz, 2000; Boneta et al., 2001; Silva et al., 2001; Sintes et al., 2002). Furthermore, they are used as either components or end products of various CaPO4-based self-setting formulations (Kuemmerle et al., 2005; Luaces-Rey et al., 2009; Sugawara et al., 1989, 2008; Mehta et al., 2015; Zhou et al., 2016; Moogi et al., 2017; Shetty et al., 2017; Medvecky et al., 2018; Dickens-Venz et al., 1994; Lee et al., 2003; Briak et al., 2002; Dickens et al., 2003; Michaı̈lesco et al., 2005; Xu et al., 2006a; Wei et al., 2011a,b; Thein-Han et al., 2012; Panahi et al., 2017) and root canal sealers (Yoshikawa et al., 1997). In addition, both compounds are added to biocomposites (Xu et al., 2006b, 2007b,c; Chen, 2010). For example, a decrease of DCPA particle dimensions was found to increase the Ca- and PO4-ions releases from DCPA-based biocomposites. Therefore such biocomposites possess both high strength and good release of Ca and PO4 ions, which may provide the
396 Advanced Dental Biomaterials needed and unique combination of stress-bearing and caries-inhibiting capabilities suitable for dental applications (Xu et al., 2007c). In addition, DCPD was tried in pulpectomy (Jean et al., 1993). 16.5.1.3 Octacalcium phosphate Just a few publications on applications of octacalcium phosphate (OCP) in dentistry and dentistry-related fields were found. Namely, OCP might be used as a coating (Stefanic et al., 2012; Bao et al., 2014), a component of biocomposites (Bao et al., 2014; Miura et al., 2012; Kawai et al., 2018), and in self-setting formulations (Sena et al., 2004; Imamura et al., 2010). In addition, OCP was tried in pulpectomy (Jean et al., 1993), as a direct pulp capping material (Sena et al., 2004), and for alveolar ridge augmentation (Miura et al., 2012; Kamakura et al., 1996). Furthermore, investigations with rats revealed that implanted OCP could serve as a core for initiating bone formation and cause osteoinduction and osteoconduction in experimentally created cranial defects (Sena et al., 2004) and enhanced reparative dentine formation via induction of odontoblast differentiation (Wang et al., 2015b). 16.5.1.4 Amorphous calcium phosphates Unlike OCP, ACPs appear to be very popular compounds for dental applications (Reynolds, 1998, 2008; Shammukha et al., 2012; Emamieh et al., 2015; Sultan et al., 2016; Hasnamudhia et al., 2017; Yates et al., 1998; Geiger et al., 2003; Ambrosio et al., 2001; Skrtic et al., 1996a,b, 2001, 2003, 2004; Skrtic and Antonucci, 2005, 2007, 2011; Oshiro et al., 2007; O’Donnell et al., 2009; Antonucci et al., 2009; Walker et al., 2010; Xu et al., 2011; Uysal et al., 2009a,b,c, 2010a,b,c, 2011; Tung and Eichmiller, 1999; Dunn, 2007; Keçik et al., 2008; Foster et al., 2008; Sun et al., 2008; Bröchner et al., 2011; Antonucci and Skrtic, 2010; Beerens et al., 2010; Zhao et al., 2011, 2012; Gupta and Prakash, 2011; Zhang et al., 2011; Moreau et al., 2011; Fletcher et al., 2011; Tabrizi and Cakirer, 2011; Hamba et al., 2011; Chow et al., 2011; Hegde and Moany, 2012; Bar-Hillel et al., 2012; Çehreli et al., 2012; Hammad et al., 2013; Park et al., 2013; Gurunathan and Somasundaram, 2015; Baysal and Uysal, 2012; Liu et al., 2018a; Prabhakar et al., 2012; Weir et al., 2012; Kato et al., 2013, 2015; Reema et al., 2014). For example, two ACP-based remineralization systems have been developed and are now commercially available: a CPP-stabilized ACP with a trade name Recaldent (Cadbury Enterprises Pte Ltd., Singapore) and an unstabilized ACP with a trade name Enamelon (Enamelon Inc., Cranbury, New Jersey, United States). CPP is produced from milk protein casein and has a remarkable ability to stabilize CaPO4 in solutions and substantially increase the level of CaPO4 in dental plaque. Therefore in Recaldent technology, it is claimed that CPP stabilizes high concentrations of calcium and orthophosphate ions, together with fluoride ions, at the tooth surface by binding to pellicle and plaque. Through the cluster sequence, CPP binds to forming nanodimensional clusters of ACP preventing their growth to the critical size required for nucleation and phase transformation. CPPACP nanodimensional
Calcium orthophosphates as a dental regenerative material 397 complexes with a hydrodynamic radius of 1.526 6 0.044 nm at pH 5 6.0, which increases to 1.923 6 0.082 nm at pH 5 9.0, are formed as a result (Cross et al., 2016). It is believed that these CPPACP nanodimensional complexes enter the porosities of an enamel subsurface lesion and diffuse down concentration gradients into the body of the subsurface lesion. Once present there, the nanodimensional complexes release the weakly bound calcium and orthophosphate ions, which would then deposit into crystal voids (Cochrane et al., 2010). Due to ACPs’ bioactivity, local Ca- and PO4-enriched environments are created with supersaturation conditions favorable for the regeneration of tooth mineral lost to decay or wear. Although all the available ions are stabilized by CPP from promoting dental calculus, they are freely available to diffuse down concentration gradients into enamel subsurface lesions thereby effectively promoting remineralization in vivo. The Enamelon technology applies calcium ions (e.g., calcium sulfate) and orthophosphate ions (e.g., ammonium orthophosphate, sometimes in the presence of fluoride ions) separately (e.g., from a dual chamber device). Therefore as the salts mix with saliva, they dissolve releasing calcium and orthophosphate ions and ACP (or F-containing ACP) forms intraorally. In the intraoral environment, both ACP and F-containing ACP are very unstable and rapidly transform to a more thermodynamically stable, insoluble crystalline phases, such as CDHA and a blend of CDHA 1 FA, respectively. It is believed that this helps rebuild tooth enamel through remineralization (Hicks and Flaitz, 2000; Schemehorn et al., 1999; Mayne et al., 2011); however, this approach may also promote dental calculus (Reynolds, 2008). Thus both previously prepared ACP (Recaldent) and in situ precipitated ACP (Enamelon, Relief ACP) are used in dentistry to remineralize tooth surface. This property of ACPs is used in toothpastes (Table 16.3). In addition, there is a commercially produced MI varnish (3M ESPE, United States), containing CPPACP Recaldent, which was found to be effective in dentinal hypersensitivity reducing (Sharma et al., 2017). As seen from the available references, in dentistry ACPs are generally used as components of various biocomposites. In an acidic oral environment, such biocomposites take advantages of the ability of ACPs to release calcium and orthophosphate ions, which potentially can take part in enamel remineralization (Vyavhare et al., 2015; Shen et al., 2001; Iijima et al., 2004; Cai et al., 2003, 2007; Morgan et al., 2008; Shammukha et al., 2012; Emamieh et al., 2015; Sultan et al., 2016; Hasnamudhia et al., 2017; Liang et al., 2017a,b, 2018; Weir et al., 2012, 2017; Xiao et al., 2017; Ambrosio et al., 2001; Skrtic et al., 1996a,b, 2001, 2003, 2004; Skrtic and Antonucci, 2005, 2007, 2016; Oshiro et al., 2007; O’Donnell et al., 2009; Antonucci et al., 2009; Reynolds, 1998; Tung and Eichmiller, 1999; Gupta and Prakash, 2011; Zhang et al., 2011, 2014, 2016a,b,c; Liu et al., 2018a,b; Prabhakar et al., 2012; Reema et al., 2014; Llena et al., 2009; Langhorst et al., 2009; Kumar et al., 2008; Ranjitkar et al., 2009a,b; Wegehaupt and Attin, 2010; Al-Mullahi and Toumba, 2010; Giniger et al., 2005a,b; Reynolds et al., 2008; Ramalingam et al., 2005; Panich and Poolthong, 2009; Silva et al., 2010; Bayrak et al., 2009; Yengopal and Mickenautsch, 2009; Walker et al., 2006, 2009; Willershausen et al., 2009; Mei et al., 2009; Melo et al., 2013a,b; Aykut-Yetkiner
398 Advanced Dental Biomaterials et al., 2014; Zhou et al., 2014; Li et al., 2014b; Xie et al., 2018; Mendes et al., 2018). The ACP-containing biocomposites and hybrid biomaterials can be prepared in various forms, such as crèmes (Zhang et al., 2011), nanodimensional fibers (Fletcher et al., 2011), or cements (Zhang et al., 2016b). Such formulations are used mainly as anticariogenic and/or remineralizing agents (Shen et al., 2001; Iijima et al., 2004; Cai et al., 2003, 2007; Morgan et al., 2008; Liang et al., 2017a,b, 2018; Weir et al., 2012, 2017; Xiao et al., 2017; Gupta and Prakash, 2011; Zhang et al., 2011, 2014, 2016a,b,c; Liu et al., 2018a,b; Prabhakar et al., 2012; Reema et al., 2014; Llena et al., 2009; Langhorst et al., 2009; Kumar et al., 2008; Ranjitkar et al., 2009a,b; Wegehaupt and Attin, 2010; Al-Mullahi and Toumba, 2010; Giniger et al., 2005a,b; Reynolds et al., 2008; Ramalingam et al., 2005; Panich and Poolthong, 2009; Silva et al., 2010; Bayrak et al., 2009; Yengopal and Mickenautsch, 2009; Walker et al., 2006, 2009; Willershausen et al., 2009; Mei et al., 2009; Melo et al., 2013a,b; Aykut-Yetkiner et al., 2014; Zhou et al., 2014; Li et al., 2014b; Xie et al., 2018; Skrtic and Antonucci, 2016; Mendes et al., 2018), for example, in chewing gums (Shen et al., 2001; Iijima et al., 2004; Cai et al., 2007; Morgan et al., 2008; Shammukha et al., 2012; Emamieh et al., 2015; Sultan et al., 2016; Hasnamudhia et al., 2017), sugar confections (Walker et al., 2010), lozenges (Cai et al., 2003), various tooth mousses (Kumar et al., 2008; Ranjitkar et al., 2009a,b), bleaching gels (Giniger et al., 2005a,b), mouth rinses (Reynolds et al., 2008), various drinks (Ramalingam et al., 2005; Panich and Poolthong, 2009), or even in milk (Walker et al., 2006, 2009). In 2016 rechargeable ACP-containing formulations with sustained ion release and rerelease were introduced (Zhang et al., 2016a,b; Xie et al., 2018; Liang et al., 2018). To improve cell adhesion, coatings composed of ACP and hyaluronic acid were used (Sun et al., 2008). Finally, ultrathin freestanding ACP sheets were manufactured and tested (Kato et al., 2013, 2015). Additional details of dental applications of ACPcontaining biocomposites are available elsewhere (Skrtic and Antonucci, 2016; Shyam et al., 2017). 16.5.1.5 α-Tricalcium phosphate and β-tricalcium phosphate According to the available literature, α-TCP and/or β-TCP (unfortunately, the authors of the publications on the subject do not always specify which of them was used) are widely used in dentistry and dentistry-related fields. For example, they are used for augmentation of the surrounding bones (Nery et al., 1978; Pepelassi et al., 1991; Wiltfang et al., 2003; Zerbo et al., 2004; Zijderveld et al., 2005; Shayesteh et al., 2008; Marukawa et al., 2011; Klijn et al., 2012; Yoshino et al., 2013), in maxillofacial surgery (Fischer-Brandies and Dielert, 1985; Ignatius et al., 2001; Horch et al., 2006; Saito et al., 2012), as a component of root canal sealers (Bilginer et al., 1997) and self-setting formulations to fill dental cavities (Dou et al., 2017), as implant coatings (Niwa et al., 2009; Yamamoto et al., 2015), as remineralization (Karlinsey et al., 2010a,b; Karlinsey and Pfarrer, 2012) and pulpotomy (Shayegan et al., 2008) agents, for dental pulp capping (Heller et al., 1975; Koenigs et al., 1975; Boone and Kafrawy, 1979; Himel et al.,
Calcium orthophosphates as a dental regenerative material 399 1985; Yoshiba et al., 1994; Higashi, 1996), to treat perforations (Sinai et al., 1989; Balla et al., 1991), as endodontic plugs (Harbert, 1991), and to fill various types of bone defects and lesions (Gaberthüel and Strub, 1977; Strub and Gaberthüel, 1978; Baldock et al., 1985; Saffar et al., 1990; Stavropoulos et al., 2010; Asvanund and Chunhabundit, 2012; Okubo et al., 2013; Saito et al., 2014; Matsuura et al., 2015; Matsuse et al., 2018). Microspheres of Mg-doped TCP were tested as delivery vehicles for tooth-bleaching agents (Mellgren et al., 2018). In addition, the investigations to assess the effect of nanodimensional β-TCP on enamel repair revealed that the use of β-TCP had a significant effect on improving enamel mineralization and penetrating into intercrystalline spaces (Rezvani et al., 2015). Furthermore, α-TCP-containing chewing gums were prepared and tested (Vogel et al., 1998, 2000). To extend a variety of applications, β-TCP can be functionalized by various organic compounds, such as sodium lauryl sulfate (Karlinsey et al., 2010a), fumaric acid (Karlinsey et al., 2010b), and some other compounds (Karlinsey and Pfarrer, 2012; Ogawa et al., 2016; Cãlin and Pãtraşcu, 2016). Functionalization of β-TCP served two major purposes: first, it provided a barrier that prevented premature β-TCPfluoride interactions, and second, it provided a targeted delivery of β-TCP when applied to the teeth. Placebo-controlled clinical studies demonstrated that if compared to fluoride alone, the combination of fluoride plus functionalized β-TCP improved remineralization by building stronger, more acid-resistant mineral in both white spot lesions as well as eroded enamel (Karlinsey et al., 2010a,b; Karlinsey and Pfarrer, 2012). A therapy of 36 teeth with deep caries by both HA and undisclosed TCP was carried out. Repeated examinations of patients 1 and 6 months after treatment showed that both HA and TCP normalized the function of the pulp and caused remineralization of dentin in the bottom of carious cavity (Maksimovskii and Zemskova, 1994). To finalize this topic, one should mention a commercially produced Clinpro white varnish (3M ESPE, United States), containing functionalized undisclosed TCP and 22,600 ppm of fluorides, as well as a toothpaste with a similar composition from the same producer (Table 16.3). 16.5.1.6 Apatites (hydroxyapatite, calcium-deficient hydroxyapatite, and fluorapatite) As in Table 16.2, apatites (HA, CDHA, and FA) appear to be the most popular type of CaPO4 used for dental applications. Since nanodimensional and nanocrystalline apatites are often considered as the model compounds of dental enamel due to both the chemical and phase similarities (LeGeros, 1991; Elliott, 1994), their use in restorative dentistry offers several promising advantages, including intrinsic radiopaque response, enhanced polishability, and improved wear performance. In addition, they have hardness similar to that of natural teeth (Wang and Wang, 2010). For example, nanodimensional HA particles were found to have an ability to infiltrate a demineralized collagen matrix of dentin. Afterward, the infiltrated collagen matrix of dentin might provide a suitable scaffold for dentin remineralization,
400 Advanced Dental Biomaterials whereby the infiltrated HA particles could act as seeds within the collagen matrix and, given the appropriate remineralizing environment, dentin remineralization might occur (Besinis et al., 2012). In addition, it was demonstrated that nano-sized HA particles could be self-assembled to form enamel-like structures (Chen et al., 2005). Therefore a localized biomimetic repair of the enamel surface could be achieved by nano-sized (B20 nm) HA, which were analogs to the basic building blocks of the enamel rods. This similarity resulted in a good fixation of artificial biomaterials to natural tissues. Moreover, the enamel structure became reinforced by nano-sized HA since a secondary caries was suppressed and the hardness was retained (Onuma et al., 2005; Meng et al., 2007; Li et al., 2007). Furthermore, nano-sized HA could be adsorbed onto the enamel surface strongly and even be integrated into the natural enamel structure (Li et al., 2008). Generally, these studies also suggest that analogs of nanodimensional building blocks of biominerals should be highlighted in the entire subject of biomineralization. This strategy may have prospective applications in dentistry as it offers an easy but effective method to reconstruct tooth enamel that is suffering from mineral losses. Normally, apatites for dental applications are prepared from the pure chemicals; however, they could also be prepared from biological sources, such as teeth (Lim et al., 2011). Due to the versatile applications in dentistry, apatites could be used in various formulations, configurations, and/or shapes. First, apatites are added to toothpastes (Table 16.3). Second, apatites are used as coatings to enhance the osteoinductivity of various dental implants (Block et al., 1987, 1994, 2000; Jones et al., 1997, 1999; Gineste et al., 1999; Ong and Chan, 1999; Yoshinari et al., 2002; Schliephake et al., 2006; Kim et al., 2011a; Jung et al., 2012; Kano et al., 2012; van Oirschot et al., 2016a; Yu and Choe, 2018). For example, degradation rates of dental implants covered by 50- and 100-μm-thick coatings of HA, FA, and fluorohydroxylapatite (FHA) were studied (Gineste et al., 1999). The implants were inserted in dog jaws and retrieved for histological analysis after 3, 6, and 12 months. The HA and FA coatings (even 100 μm thick) were almost totally degraded within the implantation period. In contrast, the FHA coatings did not show significant degradation during the same period (Gineste et al., 1999). The apatite coatings on titanium implants followed by bisphosphonate-immobilization appeared to be effective in the promotion of osteogenesis on surfaces of dental implants (Yoshinari et al., 2002). Such coatings are commercially produced. For example, there is a commercial “μ-ONE HA implant” (Yamahachi Dental, Japan) in which an HA film of 12 μm thick is sputtered by an undisclosed technique to the surface of fixing device (embedded in maxilla during surgery) of titanium implanted teeth. Regarding their durability, the HA-coated dental implants were found to work well in the short to medium terms [during 46 years (Tinsley et al., 2001), 810 years (Binahmed et al., 2007), and 14 years (Iezzi et al., 2007)]; nevertheless, even longer term clinical results are awaited with a great interest (van Oirschot et al., 2016a). Third, apatites are added as components to intermediate restorative materials (Owadally and Ford, 1994; Owadally et al., 1994), glass-ionomer cements (which are dental restorative materials used for filling teeth) (Nicholson et al., 1993; Yap et al., 2002; Lucas et al., 2003; Moshaverinia et al., 2008a; Moshaverinia et al.,
Calcium orthophosphates as a dental regenerative material 401 2008b; Arita et al., 2011; Lin et al., 2011), enamel resin infiltrants (Neto et al., 2016), various dental biocomposites (Domingo et al., 2001; Brostow et al., 2008; Oduncu et al., 2010; Huang et al., 2010b; Zhang and Wang, 2012; Vargas et al., 2013; Schlafer et al., 2016; Ai et al., 2017), dentifrices (Kani et al., 1988, 1989; Kodaka et al., 1999; Jeong et al., 2007; Najibfard et al., 2011), mouth rinses (Hannig et al., 2013; Kensche et al., 2017), and toothpastes (Comar et al., 2013; de Carvalho et al., 2014; Mielczarek and Michalik, 2014; Vyavhare et al., 2015; Souza et al., 2015; Hill et al., 2015). An artificial dental bite was produced by stereolithography from HA-containing biocomposite (Makvandi et al., 2018). Furthermore, the application of HA powder was found to be effective in apexogenesis of young permanent teeth of dogs (Marković et al., 2007). In addition, an interesting approach to control dental caries by CDHA-osteopontin biocomposites was introduced (Schlafer et al., 2016). Since dental caries is caused by acid production by bacteria in biofilms located on dental surfaces, its prevention involves a control of microorganisms producing the acids. Interestingly that HA-containing mouth rinses were found to be able to reduce the initial bacterial colonization (Kensche et al., 2017). In another study, CDHA-osteopontin biocomposite particles were prepared to bind to bacteria in the biofilms, impede biofilms building-up without killing the microflora, and release orthophosphate ions to buffer bacterial acid production if pH decreased below 6. Analysis of the results revealed that the treatment by either CDHA-osteopontin or pure osteopontin led to less biofilm formation compared to untreated controls. Thus the antibiofilm effect of the CDHA-osteopontin particles was ascribed to osteopontin, while CDHA was responsible for the buffering effect, which kept pH always above 5.5 (Schlafer et al., 2016). Fourth, there are various types of self-setting apatite-forming and/or apatitecontaining formulations (Friedman et al., 1991, 2000; Reddi et al., 1999; Abe et al., 2009; Benson and Djalilian, 2009; Lee et al., 2010a,b; Singh et al., 2010; Bambakidis et al., 2010; Krell and Wefel, 1984; Krell and Madison, 1985; Chohayeb et al., 1987; Sugawara et al., 1990, 2002; Chaung et al., 1996; Dupoirieux and Gard, 2000; Cherng et al., 2001; Fujikawa et al., 2002; Comuzzi et al., 2002; Shirakata et al., 2002; Kim et al., 2004; Noetzel et al., 2006; Witjaksono et al., 2007; Ascherman et al., 2008; Arisan et al., 2010; Barros et al., 2012; Xiao et al., 2015). For example, a cement was injected as a bone filler for gaps around oral implants placed on the medial femoral condyles of six goats and excellent bone formation around the graft material was found. Unfortunately, the degradation rate of the cement appeared to be very slow and no resorption was observed (Comuzzi et al., 2002). In another study, a cement was placed on artificially created periodontal defects but no significant difference was found between the cement and control. Nevertheless, the cement acted as a scaffold for bone formation and provided histocompatible healing of periodontal tissues (Shirakata et al., 2002). Other investigators used cements for direct pulp capping (Chaung et al., 1996; Dupoirieux and Gard, 2000) and compared them to calcium hydroxide. Both materials were found to be equally capable of producing a secondary dentin at B24 weeks (Chaung et al., 1996). Still other investigators extracted all mandibular premolar teeth from beagles (Sugawara et al., 2002). After 1 month of healing,
402 Advanced Dental Biomaterials alveolar bones were reduced to make space for previously fabricated CaPO4 cement blocks. One more month later, 8 mm HA implants were placed in such a manner that the apical half was embedded into alveolar bones and the coronal half in the cement blocks. The investigators observed that the cement blocks were gradually replaced by bone and the histopathologic features of the cement area were similar to that of natural bone. Moreover, the coronal half of the implants, previously surrounded by the cement, was firmly attached by natural bone (Sugawara et al., 2002). In another study, the same researchers used fluorescent labeling analysis and electron microanalysis to measure the extent of new bone formation and elemental (Ca, P, Mg) distribution (Fujikawa et al., 2002). Besides, several apatite-forming and/or apatite-containing self-setting formulations were tested as root canal fillers (Michaı̈lesco et al., 2005; Sugawara et al., 1990; Noetzel et al., 2006) and sealers (Krell and Wefel, 1984; Krell and Madison, 1985; Chohayeb et al., 1987; Cherng et al., 2001; Kim et al., 2004; Witjaksono et al., 2007). Since HA alone does not possess the self-setting abilities, to create a self-setting formulation it could be mixed with an epoxy resin (Witjaksono et al., 2007). To impart an antibacterial effect, an apatite-forming MCPM 1 CaO self-setting formulation with an excess of CaO (which after contact with water was transformed to Ca(OH)2) was elaborated (Koch et al., 2010). Finally, injectable forms of such cements can be used as adjunctive supportive agents for dental implants (Arisan et al., 2010). An interesting approach was performed in an attempt to regenerate tooth enamel in vitro using thin and flexible HA sheets (Hontsu et al., 2012). First, a thin HA film was deposited onto a soluble substrate by pulsed laser deposition technique. Next, the HA film was collected as a freestanding sheet by dissolving the substrate using a solvent. HA sheets of one to several microns thick and up to 50 mm in diameter could be produced by this technique. Then, the HA sheet was adhered to the extracted human teeth using a CaPO4-containing solution with a pH of 5.5. The authors found that the HA sheet was fused with tooth enamel within approximately 1 week and that the HA sheet was effective for the restoration and conservation of the tooth in dental applications (Hontsu et al., 2012). This approach was further developed in later studies by introducing a bit thicker (8 μm thick) HA sheets with an additionally deposited thin layer of undisclosed TCP of 500 nm thick (Yamamoto et al., 2013a,b). One should mention that due to a small thickness, the HA sheets are transparent (therefore invisible) and their coloration is possible. Therefore they could be applied in cosmetic dentistry. In addition, the HA sheets have a number of minute holes that allow liquid and air to escape from underneath to prevent their forming bubbles when it is applied onto a tooth. One problem is that it takes almost a whole day for an HA sheet to adhere firmly to the tooth’s surface. Similar sheets from ACP were developed and tested as well (Kato et al., 2013, 2015). More to the point, dental applications of apatites include direct pulp capping (Higashi, 1996; Jaber et al., 1991, 1992; Li et al., 1998; Hayashi et al., 1999; Zhang et al., 2008), dentin hypersensitivity treatments (Shetty et al., 2010; Low et al., 2015; Vano et al., 2018), using in endodontics (Nakagawa, 1983; Maeda et al., 1984; Roane and Benenati, 1987; Yamaguchi, 1989; Tomizuka, 1990; White and
Calcium orthophosphates as a dental regenerative material 403 Goodis, 1991; MacDonald et al., 1994; Gambarini and Tagger, 1996; Mangin et al., 2003; Yu et al., 2003; Marković et al., 2004; Teodorović and Martinović, 2005; Fathi et al., 2006; Masudi et al., 2010; Vaishnavi et al., 2011; Collares et al., 2012; Wang and Geogi, 2014; Alhashimi et al., 2016), orthodontics (Hara et al., 1989; Müller, 1989; Schneider and Diedrich, 1989; Giordano et al., 1996; Liang et al., 1998; Akhavan et al., 2013; Enan and Hammad, 2013; Seifi et al., 2015; Ajami et al., 2016), oral and maxillofacial surgery (Fischer-Brandies and Dielert, 1985; Block et al., 2000; Frame et al., 1981; Boyne, 1982; Mangano et al., 1984; Cranin et al., 1987a,b; Frame and Brady, 1987; Frame, 1987; Asanami et al., 1988; Shirakawa et al., 1988; Salyer and Hall, 1989; Hemmerle et al., 1995; Kent et al., 1983; Holmes and Hagler, 1987; Wittkampf, 1988; Friedman et al., 1998; Ylinen et al., 2002; Hallman et al., 2002; Wiltfang et al., 2004; Zecha et al., 2011; Scarano et al., 2012; Mercier, 1995; Redondo et al., 1995; Mercier et al., 1996; Lew et al., 1997; Mishra et al., 2010; Zhang et al., 2010; Sverzut et al., 2015), orthognathic surgery (Wolford et al., 1987; Moenning and Wolford, 1989; Rosen and Ackerman, 1991; Cottrell and Wolford, 1998; Wolford and Freitas, 1999), prosthodontics (Larsen et al., 1983, 1984; Balshi, 1987; Nelson and von Gonten, 1988; Tanaka et al., 1988, 1989; Zeltser et al., 1989; Denissen et al., 1989; Ogiso et al., 1994; Ngoc et al., 1997; Sung et al., 2007), and periodontics (Meffert et al., 1985; Stahl and Froum, 1987; Bowen et al., 1989; Mora and Ouhayoun, 1995; Brown et al., 1998; Yukna et al., 1998; Morris et al., 2000; Okuda et al., 2005, 2009; Kawase et al., 2010; Trombelli et al., 2010; Heinz et al., 2010; Jung et al., 2011; Horváth et al., 2013; Yoshinuma et al., 2012; Shirai et al., 2012; Machot et al., 2014; Madhumathi and Kumar, 2014; Pradeep et al., 2017; Vaca-Cornejo et al., 2017; Madhumathi et al., 2018). Several apatite-containing biocomposites are commercially produced. The examples comprise Cavalite (Kerr Italia S.r.l., Italia), which is a light-cured cavity liner containing HA and glass-ionomer powder, as well as NovaPro Fill and NovaPro Flow (Nanova Biomaterials Inc., Missouri, United States), which are nanofiber-reinforced biocomposites containing nanodimensional powder of HA. To conclude this section, one should mention that due to close chemical and phase similarities between apatites and dental enamel, dissolution of apatites in acids is considered a good model of dental caries (Dorozhkin, 2012b). 16.5.1.7 Tetracalcium phosphate According to the available literature, TTCP alone is rarely used in dentistry (Xu and Moreau, 2010; Shen et al., 2010). In the vast majority of the cases, TTCP is combined with either other types of CaPO4 (mainly DCPD or DCPA) or other chemicals to form various self-setting formulations (Sugawara et al., 1989, 2008; Mehta et al., 2015; Zhou et al., 2016; Moogi et al., 2017; Shetty et al., 2017; Medvecky et al., 2018; Dickens-Venz et al., 1994; Lee et al., 2003; Dickens et al., 2003; Xu et al., 2006a; Wei et al., 2011a,b; Thein-Han et al., 2012), biocomposites (Dickens-Venz et al., 1994; Lee et al., 2003; Dickens et al., 2003; Xu et al., 2009), as well as root canal sealers (Yoshikawa et al., 1997) and fillings (Chen et al.,
404 Advanced Dental Biomaterials 2013a). For example, a FA forming self-setting formulation consisting of solid TTCP, solid NaF, and liquid H3PO4 was prepared and used for in vitro filling of big enamel carious cavities. The results revealed that the hardened formulation was tightly combined with the enamel surface due to the chemical interaction between the formulation and enamel apatite (Wei et al., 2011b). A TTCP-containing chewing gum was prepared and tested (Chow et al., 1994). 16.5.1.8 Biphasic and multiphasic CaPO4 formulations According to the definition, biphasic and multiphasic CaPO4 formulations represent various blends of two or more individual types of CaPO4, respectively, and, among them, a biphasic calcium phosphate (abbreviated as BCP) formulation, consisting of HA and β-TCP, appears to be the most popular one (Dorozhkin, 2016a). An injectable bone and dental substitute constituted of BCP and a hydrosoluble cellulose polymer as a carrier was developed (Daculsi et al., 1999). This formulation was used for filling bone defects after tooth extractions in 11 patients. Three years after surgery, small biopsies of the implanted areas were harvested and analyzed by using microcomputed tomography, nondecalcified histology, and histomorphometry. The BCP granules appeared in direct contact with mineralized bone tissue, thereby supporting bone growth. A gradual substitution of the filler by bone tissue was observed thus preserving the height of the alveolar bone crest (Weiss et al., 2007). Similar results were obtained in another study (Struillou et al., 2011). In addition, BCP was found to be effective for healing of dental bones, osseous, and/ or intrabony defects (Huang et al., 2010a; Ellinger et al., 1986; Nery et al., 1990a,b, 1992; Sculean et al., 2008; Shi et al., 2008; Su et al., 2008; Pandit et al., 2010; Kaushick et al., 2011; Kim et al., 2011b; Wagner et al., 2012; Pietruska et al., 2012a; Wang et al., 2012; Seong et al., 2014). For example, micro- and macroporous BCP combined with a fibrin sealant was found to be safe and effective in sinus floor elevation for dental implant placement, supporting bone regeneration (Wagner et al., 2012). Furthermore, BCP was used to fill dental root canals (Bosco et al., 2006) and as a hypersensitivity treatment agent (Ngoc et al., 2018), while a triphasic CaPO4 (α-TCP 1 HA 1 TTCP) was applied as a direct pulp capping material (Kiba et al., 2010). 16.5.2 Classification according to the dental specialties 16.5.2.1 Endodontics Generally, root canal filling materials are divided into core materials and root canal sealers (Table 16.2). Root canal obturation consists of placing an inert filling material in the space previously occupied by pulp tissue. To achieve successful endodontic therapy, it is important to obturate the root canal system completely. Thus the effective endodontic obturation must provide a dimensionally stable, inert fluidtight apical seal that will eliminate any portal of communication between the canal space and the surrounding periapical tissues through the apical foramen. According
Calcium orthophosphates as a dental regenerative material 405 to the databases, the earliest publication on the use of CaPO4 in endodontics was published in Japanese in 1983 (Nakagawa, 1983), followed by a publication in English in 1984 (Maeda et al., 1984). Several examples of endodontic applications of CaPO4 are given below. A case of combined endodonticperiodontic lesions on a mandibular first molar was treated by intentional replantation and application of HA. Four months after the surgery, a porcelain-mental full crown restoration was completed. The 15-month follow-up examination showed that the tooth was clinically and radiographically healthy and functioned well (Yu et al., 2003). Several types of CaPO4 [DCPD, OCP, β-TCP, BCP (HA 1 β-TCP), and HA] in particle sizes of ,5 or ,150 μm were used for pulp capping teeth of pigs, rats, and dogs. All types of CaPO4 showed biocompatibility. Based on these results, it was suggested that these types of CaPO4 might be useful for specific clinical applications in endodontics, such as pulp capping (microparticles of HA, β-TCP, BCP) and pulpectomy (HA, OCP, DCPD) (Jean et al., 1993). Applicability of CaPO4 in pulpotomy and pulpectomy was confirmed in other studies (Shayegan et al., 2008; Nevins et al., 1980; Shayegan et al., 2010). Bone regeneration in endodontically induced periapical lesions using HA, platelet-rich plasma, and a combination of HA with platelet-rich plasma was evaluated for a period of 1 year with 20 systemically healthy patients (Vaishnavi et al., 2011). To qualify, the patient had to have a tooth where nonsurgical root canal therapy had failed, periapical radiolucency was present, and periapical root end surgery was required. The bony defect had to be confined to the apical area, with the bone covering the entire root surface coronally, with an intact lingual cortical plate. The patients were randomly divided into four groups, with five patients in each, as follows: replacement with HA; replacement with platelet-rich plasma; replacement with HA with platelet-rich plasma; and a control group with no substitutes. The radiographic evaluation revealed that the HA patients showed complete bone regeneration with evidence of a trabecular pattern at the end of 1 year, the platelet-rich plasma patients showed complete bone regeneration at the end of 9 months, the HA with platelet-rich plasma patients showed complete bone regeneration at the end of 6 months, while the control patients showed unsatisfactory bone regeneration even after 1 year. Thus HA addition to platelet-rich plasma was proven to facilitate bone regeneration (Vaishnavi et al., 2011). An injectable bone substitute made of a suspension of BCP (HA 1 β-TCP) bioceramics was used to fill dental root canals after removal of canal pulp (Bosco et al., 2006). The aim of that study was to verify the ability of a CaPO4 ceramic suspension to fill the apical zone of teeth both ex vivo and in vivo in a sheep model. The results showed that injection was possible with a good level of BCP granules at the end of the root dental canal with extracted tooth. The results of scanning electron microscopy investigations revealed mineral formation at the apex level with mineral tissue conduction between the BCP granules; however, only one tooth showed a good apical filling with a good sealing. The authors concluded that the sealing of the apex seemed to depend on the amount of BCP granules (Bosco et al., 2006).
406 Advanced Dental Biomaterials Furthermore, there are CaPO4-containing endodontic and/or root canal sealers (Yoshikawa et al., 1997; Bilginer et al., 1997; Krell and Wefel, 1984; Krell and Madison, 1985; Chohayeb et al., 1987; Cherng et al., 2001; Kim et al., 2004; Witjaksono et al., 2007; White and Goodis, 1991; MacDonald et al., 1994; Marković et al., 2004; Teodorović and Martinović, 2005; Masudi et al., 2010; Collares et al., 2012; Wang and Geogi, 2014; Alhashimi et al., 2016; Yang et al., 2007; Khashaba et al., 2009, 2011; Bae et al., 2010, 2011; Shon et al., 2012; Tiwari and Nandlal, 2012; Chang et al., 2014, 2015; Al-Haddad et al., 2015; Portella et al., 2015; Wang et al., 2017; Cahyanto et al., 2017; Shieh et al., 2017). The composition of two examples of such sealers (Sankin apatite root canal sealer and Capseal) are presented in Table 16.5 (Shon et al., 2012). Of them, Capseal was found to result in both higher alkalinity and higher calcium ion releases than Sankin apatite root canal sealers (Bae et al., 2011). The results of their application revealed that the sealers mentioned in Table 16.5 facilitated the periapical dentoalveolar and alveolar healing by controlling cellular mediators from periodontal ligament cells and osteoblast differentiation of precursor cells (Shon et al., 2012). More to the point, endodontic perforations were treated by CaPO4 (Sinai et al., 1989; Balla et al., 1991; Roane and Benenati, 1987; Chau et al., 1997), but once a lack of complete healing was noticed (Balla et al., 1991). Additional examples of the endodontic applications of CaPO4 comprise the following cases. They can be used as components of endodontic cements (Dou et al., 2017; Yamaguchi, 1989; Cahyanto et al., 2017; Shieh et al., 2017; Chau et al., 1997; Cherng et al., 2010) or coatings for endodontic dental implants (Fathi et al., 2006), as well as serve as a root endfilling material (MacDonald et al., 1994; Mangin et al., 2003) and as endodontic endosseous implants (Tomizuka, 1990). Since CaPO4 does not cause inflammation (Sinai et al., 1989), it could be used as a hard plug deep inside teeth (Harbert, 1991). Finally yet importantly, CaPO4 crowns were manufactured (Ishida et al., 1992). Table 16.5 Composition of the available CaPO4-containing sealer materials (Shon et al., 2012). Brand name Manufacturer Components Sankin apatite root canal sealer (I, II, and III) Sankin Kogyo, Tokyo, Japan Capseal (I and II) Experimental Powder: α-TCP and Sankin HA in type I, iodoform is added to powder in type II (30%) and type III (5%). Liquid: polyacrylic acid and water Powder: TTCP and DCPA, Portland cement (gray cement in type I and white cement in type II), zirconium oxide, and others. Liquid: hydroxypropyl methyl cellulose in sodium phosphate solution DCPA, Dicalcium phosphate anhydrous; TTCP, tetracalcium phosphate; α-TCP, α-tricalcium phosphate.
Calcium orthophosphates as a dental regenerative material 407 16.5.2.2 Oral and maxillofacial surgery An insufficient bone volume and a poor bone density are common problems in edentulous patients with resorbed maxilla. One method that makes implant placement possible in such difficult situations is the augmentation of maxillary sinus using various bone grafts (Tiwari and Nandlal, 2012). In addition, there are other cases in which bone grafts appear to be necessary for dentistry-related fields. Due to these cases, CaPO4 has been used in oral and maxillofacial surgery since the 1980s (Fischer-Brandies and Dielert, 1985; Frame et al., 1981; Boyne, 1982; Mangano et al., 1984; Cranin et al., 1987a,b; Frame and Brady, 1987; Frame, 1987; Asanami et al., 1988; Shirakawa et al., 1988; Salyer and Hall, 1989) and up to now many scientific articles have been published on the subject (Friedman et al., 1991, 2000; Reddi et al., 1999; Smartt et al., 2005; Kuemmerle et al., 2005; Luaces-Rey et al., 2009; Tamimi et al., 2009; Abe et al., 2009; Benson and Djalilian, 2009; Lee et al., 2010a; Singh et al., 2010; Bambakidis et al., 2010; Huang et al., 2010a; Sanada et al., 2011; Ignatius et al., 2001; Hemmerle et al., 1995). However, as discussed in Section 16.2, the vast majority of the publications on this subject deal with a treatment of the surrounding bones and, thus, they fall into a category of bone substitutes, which is another story. Nevertheless, the following directions of CaPO4 application in oral and maxillofacial surgery can be outlined: coatings on various types of dental implants (Block et al., 1987, 1994, 2000; Jones et al., 1997, 1999; Gineste et al., 1999; Ong and Chan, 1999; Yoshinari et al., 2002; Schliephake et al., 2006; Kim et al., 2011a; Jung et al., 2012; Kano et al., 2012; van Oirschot et al., 2016a,b; Yu and Choe, 2018; Tinsley et al., 2001; Binahmed et al., 2007; Iezzi et al., 2007; Fathi et al., 2006; Hulshoff and Jansen, 1997; Alexander et al., 2009; Junker et al., 2011; Palarie et al., 2012; Alghamdi et al., 2013), augmentation of the surrounding bones (Tamimi et al., 2009; Miura et al., 2012; Kamakura et al., 1996; Wiltfang et al., 2003, 2004; Zerbo et al., 2004; Zijderveld et al., 2005; Shayesteh et al., 2008; Marukawa et al., 2011; Klijn et al., 2012; Yoshino et al., 2013; Block et al., 2000; Sugawara et al., 2002; Schneider and Diedrich, 1989; Giordano et al., 1996; Liang et al., 1998; Akhavan et al., 2013; Enan and Hammad, 2013; Seifi et al., 2015; Ajami et al., 2016; Frame et al., 1981; Boyne, 1982; Mangano et al., 1984; Cranin et al., 1987a,b; Frame and Brady, 1987; Frame, 1987; Asanami et al., 1988; Shirakawa et al., 1988; Salyer and Hall, 1989; Hemmerle et al., 1995; Kent et al., 1983; Holmes and Hagler, 1987; Wittkampf, 1988; Friedman et al., 1998; Ylinen et al., 2002; Hallman et al., 2002; Zecha et al., 2011; Scarano et al., 2012; Sato et al., 2009), and use as fillers of osseous mandible and/or jaw defects (Horch et al., 2006; Dupoirieux and Gard, 2000; Ascherman et al., 2008; Mercier, 1995; Redondo et al., 1995; Mercier et al., 1996; Lew et al., 1997; Mishra et al., 2010; Zhang et al., 2010; Sverzut et al., 2015). 16.5.2.3 Orthodontics According to the databases, the earliest publications on the use of CaPO4 in orthodontics appeared in 1989 (Hara et al., 1989; Müller, 1989; Schneider and Diedrich,
408 Advanced Dental Biomaterials Figure 16.4 Appearances of CaPO4 ceramic brackets (A) Hyaline and (B) Hyaline II. Source: Reprinted with permission from Meguro, D., Hayakawa, T., Kasai, K., 2006b. Efficacy of using orthodontic adhesive resin in bonding and debonding characteristics of a calcium phosphate ceramic bracket. Orthod. Waves 65, 148154. 1989). Coatings of CaPO4 [both HA (Liang et al., 1998) and α-TCP (Niwa et al., 2009)] were successfully applied to titanium implants and the coated implants were found to be applicable as anchorage for short-term orthodontic treatment (Liang et al., 1998) and both types of coatings appeared to be effective stimulators of new bone formation (Niwa et al., 2009). In another study, HA addition to an orthodontic cement was found to have a protective action on the dental enamel near the orthodontic bands or brackets (Giordano et al., 1996). Furthermore, there are CaPO4 bioceramic brackets Hyaline (Tomy International Inc., Tokyo, Japan) (Fig. 16.4). In addition to excellent biocompatibility, these brackets have a hardness equivalent to that of tooth enamel, which eliminates fears of dental abrasion due to the occluding tooth even when the patient has a deep-bite (Meguro et al., 2006a,b; Joo and Park, 2007). Besides, there is a study, in which researchers improved antibacterial properties of commercially produced light cure adhesive Transbond XT (3M Oral Care, Minnesota, United States) by addition of nanodimensional composite particles of HA with silver. The Transbond XT composite disks containing 5% and 10% of Ag/HA were found to produce bacterial growth inhibition zones and showed antibacterial properties against biofilms (Sodagar et al., 2016). However, among all the available types of CaPO4 (Table 16.1), ACP-containing formulations are most often used in orthodontics (Uysal et al., 2009a,b,c, 2010a,b,c, 2011; Dunn, 2007; Keçik et al., 2008; Foster et al., 2008; Bröchner et al., 2011; Beerens et al., 2010; Tabrizi and Cakirer, 2011; Chow et al., 2011; Hegde and Moany, 2012; Bar-Hillel et al., 2012; Çehreli et al., 2012; Hammad et al., 2013; Park et al., 2013; Gurunathan and Somasundaram, 2015; Baysal and Uysal, 2012; Liu et al., 2018a). For example, an efficacy of an ACP-containing orthodontic biocomposite and a resin-modified glass-ionomer cement on enamel demineralization adjacent to orthodontic brackets was evaluated by a new laser fluorescence device. The authors concluded that both formulations should be recommended for any atrisk orthodontic patient to provide preventive actions and potentially remineralize subclinical enamel demineralization (Uysal et al., 2009a). Similarly, ACPcontaining orthodontic biocomposites were found to reduce both enamel decalcification around orthodontic brackets (Uysal et al., 2010b,c; Chow et al., 2011) and
Calcium orthophosphates as a dental regenerative material 409 bacterial adherence (Chow et al., 2011). Furthermore, ACP-containing orthodontic biocomposites were found to possess a lower but still satisfactory bond strength needed to function as orthodontic adhesives (Uysal et al., 2009b, 2010a; Dunn, 2007; Foster et al., 2008; Sun et al., 2008). Therefore CPPACP biocomposite, either alone or combined with fluoride, may safely be used as a prophylactic agent before bracket bonding (Tabrizi and Cakirer, 2011; Çehreli et al., 2012). In addition, a pretreatment by CPPACP, enamel microabrasion, and the combination of these two methods were found to improve the bonding of orthodontic brackets to demineralized enamel (Çehreli et al., 2012). Rechargeable orthodontic self-setting formulations containing ACP with sustained ion release and rerelease were developed, as well (Zhang et al., 2016b; Xie et al., 2018; Liang et al., 2018). Adiaphorous results were obtained as well. For example, a topical treatment of white spot lesions after debonding of orthodontic appliances with a CPP-stabilized ACP agent resulted in significantly reduced fluorescence and reduced areas of the lesions after 4 weeks; however, the improvement was not superior to the natural regression following daily use of fluoride toothpaste (Bröchner et al., 2011). In addition, no clinical advantages for use of a CPP-fluoridated ACP paste supplementary to normal oral hygiene over the time span of 12 weeks were found in another study (Beerens et al., 2010). 16.5.2.4 Prosthodontics Humans have long used both natural and synthetic materials as replacements for lost teeth. For example, the earliest known dental implant was made of iron and found in a Roman male, who lived around the 1st or 2nd century CE (Crubezy et al., 1998). The first known tooth made from a natural material was found in a Mayan woman, estimated around 600 CE, and was made of nacre from seashells (Bobbio, 1972). Nevertheless, despite a long history of tooth grafts, just a few publications on prosthodontic applications of CaPO4 are available (Table 16.2). According to the databases, the earliest publication on the subject was published in 1983 (Larsen et al., 1983), followed by another publication by the same authors (Larsen et al., 1984). A 4-year study and evaluation of nonresorbable HA to augment different alveolar ridges was performed. The technique used resulted in improved contour, height, and width of the alveolar ridge. The state and health of the tissues were found to be improved with the use of HA or HA combined with bone marrow (Larsen et al., 1983). However, the study dealt with a treatment of bones but not teeth, which is another story. Similar can be said about other publications on the subject (Balshi, 1987; Nelson and von Gonten, 1988; Tanaka et al., 1988, 1989; Zeltser et al., 1989; Denissen et al., 1989; Ogiso et al., 1994; Ngoc et al., 1997). Furthermore, as seen from the publication dates, all these papers were published in the last century and only one recent paper (Sung et al., 2007) has been found. Nevertheless, even this recent paper is devoted to the subject of preparation, with just a possibility of using the material as dental prosthesis. Thus one can mention the past attempts to use CaPO4 in prosthetic dentistry and, since no promising results were obtained, currently CaPO4 is not used in prosthodontics.
410 Advanced Dental Biomaterials To finalize this topic, it is important to mention that one of the challenges in dental implantology is to achieve and maintain a good osseointegration, as well as an epithelial junction of gingival tissues with the implants. An intimate junction among them may prevent bacteria colonization leading to peri-implantitis, while the direct bonding may ensure a biomechanical anchoring of the artificial dental roots (Fig. 16.5) (Lavenus et al., 2010). To achieve this, the presence of sufficient bone volume is an important prerequisite for dental implant placement. However, this is not always the case. Namely, atrophic maxilla and mandible bones are less tolerant to the placement of dental implants due to their reduced height and width; hence, supplementary bone augmentation by CaPO4 might be necessary Figure 16.5 Tissue integration of a dental implant. Note the intimate contact with gingival tissue in the upper part and the desired contact osteogenesis in the tapered lower part rather than distance osteogenesis. Source: Reprinted with permission from Lavenus, S., Louarn, G., Layrolle, P., 2010. Nanotechnology and dental implants. Int. J. Biomater. 2010, 915327.
Calcium orthophosphates as a dental regenerative material 411 (Khoury et al., 2007; Chiapasco et al., 2009). In addition, I would like to point the readers’ attention to a review on dental implants for patients with osteoporosis. According to the authors, osteoporosis is not a contraindication for the implant surgery if the accurate analysis of bone quality has been performed (Gaetti-Jardim et al., 2011). 16.5.2.5 Periodontics (periodontology) In general, the regeneration of tissues affected by periodontal disease is a complex process; it encompasses the formation of bones, cementum, and periodontal ligaments (Reis et al., 2011). According to the databases, the earliest publication on the use of CaPO4 in periodontics was published in 1974 (Levin et al., 1974), followed by research papers of 1975 (Nery et al., 1975) and 1977 (Gaberthüel and Strub, 1977), and a review of 1978 (Strub and Gaberthüel, 1978). A schematic diagram of the management of periodontal defects by a bone graft technique is shown in Fig. 16.6 (Chen et al., 2010). However, as discussed in Section 16.2, the vast majority of the publications on periodontics deal with the treatment of the surrounding bones and, thus, they fall into a category of bone substitutes (Xu et al., 2006a; Pepelassi et al., 1991; Baldock et al., 1985; Saffar et al., 1990; Stavropoulos et al., 2010; Asvanund and Chunhabundit, 2012; Okubo et al., 2013; Saito et al., 2014; Matsuura et al., 2015; Matsuse et al., 2018; Mellgren et al., 2018; Rezvani et al., 2015; Ogawa et al., 2016; Cãlin and Pãtraşcu, 2016; Xiao et al., 2015; Meffert et al., 1985; Stahl and Froum, 1987; Bowen et al., 1989; Mora and Ouhayoun, 1995; Brown et al., 1998; Yukna et al., 1998; Morris et al., 2000; Okuda et al., 2005, 2009; Kawase et al., 2010; Trombelli et al., 2010; Heinz et al., 2010; Jung et al., 2011; Horváth et al., 2013; Yoshinuma et al., 2012; Shirai et al., 2012; Machot et al., 2014; Madhumathi and Kumar, 2014; Pradeep et al., 2017; VacaCornejo et al., 2017; Madhumathi et al., 2018; Nery et al., 1992; Shi et al., 2008; Figure 16.6 A schematic diagram of the management of periodontal defects by a bone graft technique. The CaPO4 grafts stimulate bone growth and a new bone fills the defect, which provides a better support for a tooth (A) open, (B) add, (C) close and fix. Source: Reprinted with permission from Chen, F.M., Zhang, J., Zhang, M., An, Y., Chen, F., Wu, Z.F., 2010. A review on endogenous regenerative technology in periodontal regenerative medicine. Biomaterials 31, 78927927.
412 Advanced Dental Biomaterials Kaushick et al., 2011; Pietruska et al., 2012a; Wang et al., 2012; Seong et al., 2014; Hayashi et al., 2006; Shirakata et al., 2008, 2012; Chitsazi et al., 2011; D’Lima et al., 2014; Bansal et al., 2014; Dan et al., 2014; Nevins and Cymerman, 2015; Elgendy and Shady, 2015). Nevertheless, a few examples are given next. Postextraction bone resorption is an increasing problem in modem dentistry. Namely, after extraction of a tooth, the bony socket heals naturally. First, it is immediately filled with coagulated blood. In a few days afterward, the granular and fibrous tissues are organized to gradually form a new bone tissue. However, due to the tooth absence, maxilla and/or mandibular alveolar atrophies occur simultaneously. These resorptive and remodeling phenomena of the surrounding bone negatively affect the support for the adjacent teeth; the shallow ridge makes it difficult for future prosthesis retention and less bony support remains for any dental implant placement in the future. To promote healing, the socket of an extracted tooth might be filled by CaPO4 bioceramics. For example, an efficacy of commercial HA granules APAFILL-G as a filler to prevent the resorption of alveolar bone after tooth extraction was studied (Garcı́a et al., 2001). After 7 days, the result revealed that only one of all treated patients experienced an adverse response that was observed under the clinical evaluation and that promptly disappeared after analgesic treatment. The remaining 32 had no adverse clinical response. Radiographically, a continuous radiopacity between bone and the implant resorption was detected after 1 year and the surrounding alveolar bone maintained its contour without symptoms of resorption for 100% of the patients (Garcı́a et al., 2001). In another study, two different types of HA grafting materials, biomimetic and nanocrystalline, were placed into fresh extraction sockets aiming to limit bone resorption. The surgical sites were histologically, clinically, and radiographically evaluated for 6 months after tooth extraction. The percentages of bone, osteoid areas, and residual material in the two groups were not statistically different (Checchi et al., 2011). Furthermore, repositioning maxillary and mandibular bone segments in orthognathic surgery frequently creates bone gaps or continuity defects. These often require grafting to provide positional stability and bony continuity and CaPO4 is used for this purpose. For example, as early as 1987 there was a study to evaluate the use of coralline porous HA as a bone graft substitute in orthognathic surgery (Wolford et al., 1987), followed by another study in 1989 (Moenning and Wolford, 1989). Ninety-two individual patients received a total of 355 block implants to the maxilla (294), mandible (41), and midface (20). There were 202 implants positioned directly adjacent to the maxillary sinus. Complications were minimal, the most common being exposure of the implant to the oral or nasal cavity. Histological evaluation of implants that were biopsied in nine patients, 416 months’ postsurgery, revealed connective tissue ingrowth throughout the implants with approximately one-third being bone of variable maturity and two-thirds being soft tissue (Wolford et al., 1987). Similar results were obtained in another study (Moenning and Wolford, 1989). Periodontal ligaments around extracted sockets were found to have an ability to regenerate bone on HA-coated tooth-shaped implants (Kano et al., 2012). Positive results were also observed in another study, in which bone formation around BCP (HA 1 β-TCP) particles in periodontal defects
Calcium orthophosphates as a dental regenerative material 413 of dogs were found to be more discernible if compared to healing in the control (Struillou et al., 2011). In addition, the porosity of the implanted CaPO4 was found to influence the periodontal healing of furcation defects in dogs (Saito et al., 2012). To increase a treatment efficiency of the periodontal defects, CaPO4 might be combined with the biologically active molecules, such as hormones, growth factors, morphogenetic proteins (Sculean et al., 2008; Lee et al., 2010c; Emerton et al., 2011; Ridgway et al., 2008; Jayakumar et al., 2011; Sorensen et al., 2004; Pietruska et al., 2012b; Elangovan et al., 2013). For example, an application of recombinant human growth and differentiation factor-5 (rhGDF-5) lyophilized onto β-TCP granules demonstrated an effective regeneration of the artificially created periodontal defects (Lee et al., 2010c; Emerton et al., 2011). Positive results were also obtained for a combination of a recombinant human platelet-derived growth factor BB (rhPDGF-BB) with β-TCP for the treatment of human intraosseous periodontal defects (Ridgway et al., 2008; Jayakumar et al., 2011). However, a combination of an enamel matrix derivative with BCP (HA 1 β-TCP) resulted in none to minimal new bone formation (Sculean et al., 2008). Furthermore, a combination of human bone morphogenetic protein-2 (rhBMP-2) with a bioresorbable CaPO4 cement Ceredex (ETEX Corporation, Massachusetts, United States) was not suggested for periodontal indications (Sorensen et al., 2004). Besides, there are results indicating that the use of CaPO4 after open flap procedure does not improve the clinical and radiological treatment outcomes of periodontal intrabony defects (Pietruska et al., 2012b). Thus applications of CaPO4 in periodontology were not always positive. Additional details on this topic are available in special reviews (Bayani et al., 2017; Sepantafar et al., 2018). 16.5.2.6 Other types of oral applications Mucositis is the painful inflammation and ulceration of the mucous membranes lining the digestive tract, usually as an adverse effect of chemotherapy and radiotherapy treatment for cancer. Mucositis can occur anywhere along the gastrointestinal tract, but oral mucositis refers to the particular inflammation and ulceration that occurs in the mouth (Mucositis). It can delay discharge, interrupt treatment, and threaten life. To help the patients, rinses, supersaturated by undisclosed types of CaPO4, were prepared and evaluated (Papas, 2008; Markiewicz et al., 2012; Miyamoto et al., 2012; Quinn, 2013; Svanberg et al., 2015; Bhatt et al., 2017). For example, a review paper is available, in which the author looked at data from all known adult and pediatric studies investigating the use of a supersaturated by CaPO4 mouth rinse Caphosol (EUSA Pharma, United Kingdom) in patients receiving high-dose cancer therapy in order to evaluate its efficacy for both the prevention and treatment of oral mucositis. Thirty studies were identified. The majority of these studies (n 5 24) found Caphosol to be efficacious at reducing the grade and/or duration, as well as pain associated with this disease (Quinn, 2013). However, the most recent data claim that supersaturated CaPO4 rinses were found to be ineffective in treating mucositis (Svanberg et al., 2015; Bhatt et al., 2017).
414 16.6 Advanced Dental Biomaterials Tissue engineering approaches As seen from the aforementioned, CaPO4 is widely used in dentistry to restore and/ or repair various types of oral defects and diseases. However, all the previously mentioned approaches have encountered shortcomings if compared to the normal and healthy teeth and surrounding bones. Therefore various tissue engineering approaches to develop new strategies for tooth regeneration have been attempted. The history of tissue engineering in dentistry started in 1982, when the first regeneration technology of periodontium was introduced (Nyman et al., 1982). The modern tissue engineering approaches in dentistry include combinations of cells, engineering materials, and suitable biochemical and physicochemical factors to improve or replace biological functions. Finally, it will cause in vivo formation and growth of new functional tissues instead of reparation and/or replacement of damaged and/or missing ones by artificial materials and/or implants (Chai and Slavkin, 2003; Hu et al., 2006; Duailibi et al., 2008; Ikeda et al., 2009; Horst et al., 2012). From the material point of view, there are two main approaches toward making a bioengineered tooth: scaffold-free and scaffold-based regenerations. The scaffold-free approaches, such as tissue recombination, cell pellet engineering, and chimerical tooth engineering, are being developed and the correct tooth-like structures could be generated after transplantation in the subrenal capsule (Yelick and Vacanti, 2006; Nakao et al., 2007; Yu et al., 2007). However, with an exception of using soluble calcium- and orthophosphate-containing solutions to promote proliferation, osteogenic differentiation, and mineralization of various types of dental cells (An et al., 2012), the scaffold-free approaches do not utilize CaPO4. Therefore in this review, only scaffold-based tooth regeneration approaches are considered. A schematic drawing of this process is shown in Fig. 16.7 (Kitamura et al., 2012). For example, it was hypothesized that dental follicle cells combined with β-TCP might become a novel therapeutic strategy to restore periodontal defects. The Figure 16.7 A schematic diagram of entire tooth regeneration from the proper combination of growth factors and cells (stem cells or progenitor cells) seeded on a CaPO4 scaffold. Source: Adapted with permission from Kitamura, C., Nishihara, T., Terashita, M., Tabata, Y., Washio, A., 2012. Local regeneration of dentin-pulp complex using controlled release of FGF-2 and naturally derived sponge-like scaffolds. Int. J. Dent. 2012, 190561 (8 pp.).
Calcium orthophosphates as a dental regenerative material 415 authors suggested isolation of dental follicle cells from a beagle dog. The isolated cells should be induced by bone morphogenetic protein-2, basic-fibroblast growth factors, and dexamethasone and, then, seeded onto β-TCP bioceramics. Afterward, the complex should be autoimplanted into the periodontal defects in the same dog to observe regeneration of periodontal tissue in vivo (Zuolin et al., 2010). However, this was just a hypothesis. Let me describe the real investigations. The biocompatibility of four different types of 3D scaffolds for the regeneration of tooth tissues was tested (Zheng et al., 2011). The scaffolds consisted of pure poly(lactic-co-glycolic) acid (PLGA) or 50/50 w/w biocomposites of PLGA with HA, β-TCP, or carbonate-containing HA. Afterward, human dental pulp stem cells were seeded onto the scaffolds, followed by implantation into the mesentery or subrenal capsule of mice or rats for 45 weeks. The results showed that while all CaPO4-containing formulations were able to support effective regeneration of the tooth tissues, the PLGA/β-TCP scaffolds appeared to be superior to the other three scaffolds for tooth tissues regeneration, especially for dentin formation (Zheng et al., 2011). Very promising results were also obtained by other researchers for β-TCP/chitosan biocomposites (Liao et al., 2010), recombinant human transforming growth factor-beta 1 (rhTGF-β1) combined with two different bone grafts: calcified freeze-dried bone allograft and porous BCP (Markopoulou et al., 2011) and a complex of recombinant human bone morphogenetic protein-2 (rhBMP-2)mediated dental pulp stem cells and nano-sized HA/collagen/poly(Llactide) for clinical reconstruction of periodontal bone defects (Liu et al., 2011). Positive results were also obtained for polycaprolactone-TCP scaffolds (Goh et al., 2014). In still other studies, polyglycolic acid (PGA) scaffolds were compared with β-TCP, fibrin, and collagen scaffolds for their capacity to grow dental structures when seeded with tooth germs from 6-month-old minipigs. On fibrin and collagen gels, the porcine third molar tooth bud maintained its epithelial structure, resembling tooth buds, whereas on PGA and β-TCP the implanted tooth buds produced more dentin-like material (Ohara et al., 2010). Porous BCP (HA 1 β-TCP), powdered BCP, and PGA fiber mesh were used as scaffolds and transplanted with cultured porcine dental pulp-derived cells into the backs of nude mice for 6 weeks. Although newly formed hard tissues were observed in all implants, a dentin-like hard tissue was observed when porous BCP was used (Tonomura et al., 2010). Besides, incorporation of nano-sized HA into electrospun poly(ε-caprolactone)/gelatin scaffolds was found to enhance dental pulp stem cells differentiation toward an odontoblast-like phenotype both in vitro and in vivo (Yang et al., 2010). The osteoblast marker bone sialoprotein was highly expressed on β-TCP scaffolds seeded by dental follicle cells but almost absent in differentiated dental follicle cells without β-TCP (Viale-Bouroncle et al., 2011). The latter means that dental progenitor cells have to be combined with CaPO4 bioceramics. To conclude this topic, the tissue engineering approaches of dental regeneration, obviously, appear to be the most promising healing technologies and many interesting studies on a combination of CaPO4 scaffolds with cells and/or growth factors are expected to appear in the near future.
416 16.7 Advanced Dental Biomaterials Conclusion The biologically relevant types of CaPO4 are the emerging bioceramics, which are widely used in various biomedical applications, including dentistry. They have excellent biomedical properties and biological behavior because their composition and structure are similar to those of human bones and teeth. Therefore CaPO4 possess exceptional biocompatibility and unique bioactivity, which are widely used in dentistry and dentistry-related fields. For example, incorporation of CaPO4 into various restorative biomaterials was found to improve the mechanical properties of the biomaterials without impeding their inherent biological properties. Other examples have been described earlier. Nevertheless, the versatile employing strategies of CaPO4 in dentistry aim to ultimately reach the same goal, namely, to enhance osseointegration process of dental implants in the context of immediate loading and to augment formation of surrounding bones to guarantee a long-term success. However, still the complete understanding related to use of CaPO4 in clinical dentistry is lacking and further research is needed to improve their efficacy in clinical dentistry. References Abe, T., Anan, M., Kamida, T., Fujiki, M., 2009. Surgical technique for anterior skull base reconstruction using hydroxyapatite cement and titanium mesh. Acta Neurochirur. 151, 13371338. Ai, M., Du, Z., Zhu, S., Geng, H., Zhang, X., Cai, Q., et al., 2017. Composite resin reinforced with silver nanoparticlesladen hydroxyapatite nanowires for dental application. Dent. Mater. 33, 1222. Ajami, S., Pakshir, H.R., Babanouri, N., 2016. Impact of nanohydroxyapatite on enamel surface roughness and color change after orthodontic debonding. Prog. Orthod. 17, 11 (8 pp.). Akhavan, A., Sodagar, A., Mojtahedzadeh, F., Sodagar, K., 2013. Investigating the effect of incorporating nanosilver/nanohydroxyapatite particles on the shear bond strength of orthodontic adhesives. Acta Odontol. Scand. 71, 10381042. Alexander, F., Christian, U., Stefan, T., Christoph, V., Reinhard, G., Georg, W., 2009. Longterm effects of magnetron-sputtered calcium phosphate coating on osseointegration of dental implants in non-human primates. Clin. Oral Implant. Res. 20, 183188. Alghamdi, H.S., van Oirschot, B.A.J.A., Bosco, R., van den Beucken, J.J., Aldosari, A.A.F., Anil, S., et al., 2013. Biological response to titanium implants coated with nanocrystals calcium phosphate or type 1 collagen in a dog model. Clin. Oral Implant. Res. 24, 475483. Al-Haddad, A., Abu Kasim, N.H., Che Ab Aziz, Z.A., 2015. Interfacial adaptation and thickness of bioceramic-based root canal sealers. Dent. Mater. J. 34, 516521. Alhashimi, R.A., Mannocci, F., Sauro, S., 2016. Experimental polyethylene-hydroxyapatite carrier-based endodontic system: an in vitro study on dynamic thermomechanical properties, sealing ability, and measurements of micro-computed tomography voids. Eur. J. Oral Sci. 24, 279286. Al-Mullahi, A.M., Toumba, K.J., 2010. Effect of slow-release fluoride devices and casein phosphopeptide/amorphous calcium phosphate nanocomplexes on enamel remineralization in vitro. Caries Res. 44, 364371.
Calcium orthophosphates as a dental regenerative material 417 Ambrosio, A.M.A., Sahota, J.S., Khan, Y., Laurencin, C.T., 2001. A novel amorphous calcium phosphate polymer ceramic for bone repair: I. Synthesis and characterization. J. Biomed. Mater. Res. 58, 295301. An, S., Ling, J., Gao, Y., Xiao, Y., 2012. Effects of varied ionic calcium and phosphate on the proliferation, osteogenic differentiation and mineralization of human periodontal ligament cells in vitro. J. Periodontol. Res. 47, 374382. Antonucci, J.M., Skrtic, D., 2010. Fine-tuning of polymeric resins and their interfaces with amorphous calcium phosphate. A strategy for designing effective remineralizing dental composites. Polymers 2, 378392. Antonucci, J.M., O’Donnell, J.N.R., Schumacher, G.E., Skrtic, D., 2009. Amorphous calcium phosphate composites and their effect on compositeadhesivedentin bonding. J. Adhes. Sci. Technol. 23, 11331147. Arafa, A., 2017. Synergetic remineralization effectiveness of calcium, phosphate and fluoride based systems in primary teeth. Pediatric Dent. J. 27, 6571. Arisan, V., Anil, A., Wolke, J.G., Özer, K., 2010. The effect of injectable calcium phosphate cement on bone anchorage of titanium implants: an experimental feasibility study in dogs. Int. J. Oral Maxillofac. Surg. 39, 463468. Arita, K., Yamamoto, A., Shinonaga, Y., Harada, K., Abe, Y., Nakagawa, K., et al., 2011. Hydroxyapatite particle characteristics influence the enhancement of the mechanical and chemical properties of conventional restorative glass ionomer cement. Dent. Mater. J. 30, 672683. Asanami, S., Koike, O., Chikata, M., Shiba, H., Ikeuchi, S., Okada, Y., et al., 1988. Studies of the clinical usefulness of porous hydroxylapatite in the field of dental and oral surgery. Keio J. Med. 37, 265281. Ascherman, J.A., Foo, R., Nanda, D., Parisien, M., 2008. Reconstruction of cranial bone defects using a quick-setting hydroxyapatite cement and absorbable plates. J. Craniofac. Surg. 19, 11311135. Asvanund, P., Chunhabundit, P., 2012. Alveolar bone regeneration by implantation of nacre and B-tricalcium phosphate in guinea pig. Implant Dent. 21, 248253. Aykut-Yetkiner, A., Kara, N., Ateş, M., Ersin, N., Ertuğrul, F., 2014. Does casein phosphopeptid amorphous calcium phosphate provide remineralization on white spot lesions and inhibition of Streptococcus mutans? J. Clin. Pediatr. Dent. 38, 302306. Bae, W.J., Chang, S.W., Lee, S.I., Kum, K.Y., Bae, K.S., Kim, E.C., 2010. Human periodontal ligament cell response to a newly developed calcium phosphate-based root canal sealer. J. Endod. 36, 16581663. Bae, K.H., Chang, S.W., Bae, K.S., Park, D.S., 2011. Evaluation of pH and calcium ion release in capseal I and II and in two other root canal sealers. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 112, e23e28. Baldock, W.T., Hutchens Jr., L.H., McFall Jr., W.T., Simpson, D.M., 1985. An evaluation of tricalcium phosphate implants in human periodontal osseous defects of two patients. J. Periodontol. 56, 17. Balla, R., LoMonaco, C.J., Skribner, J., Lin, L.M., 1991. Histological study of furcation perforations treated with tricalcium phosphate, hydroxylapatite, amalgam, and life. J. Endod. 17, 234238. Balshi, T.J., 1987. Preventive durapatite ridge augmentation for esthetic fixed prosthodontics. J. Prost. Dent. 58, 266270. Bambakidis, N.C., Munyon, C., Ko, A., Selman, W.R., Megerian, C.A., 2010. A novel method of translabyrinthine cranioplasty using hydroxyapatite cement and titanium mesh: a technical report. Skull Base 20, 157161.
418 Advanced Dental Biomaterials Bansal, R., Patil, S., Chaubey, K., Thakur, R., Goyel, P., 2014. Clinical evaluation of hydroxyapatite and β-tricalcium phosphate composite graft in the treatment of intrabony periodontal defect: a clinico-radiographic study. J. Indian Soc. Periodontol. 18, 610617. Bao, L., Liu, J., Shi, F., Jiang, Y., Liu, G., 2014. Preparation and characterization of TiO2 and Si-doped octacalcium phosphate composite coatings on zirconia ceramics (Y-TZP) for dental implant applications. Appl. Surf. Sci. 290, 4852. Bar-Hillel, R., Feuerstein, O., Tickotsky, N., Shapira, J., Moskovitz, M., 2012. Effects of amorphous calcium phosphate stabilized by casein phosphopeptides on enamel de- and remineralization in primary teeth: an in vitro study. J. Dent. Child. (Chic.) 79, 914. Barros, C.M.B., de Oliveira, S.V., Marques, J.B., Viana, K.M.S., Costa, A.C.F.M., 2012. Analysis of the hydroxyapatite incorporate MTA dental application. Mater. Sci. Forum 727728, 13811386. Bayani, M., Torabi, S., Shahnaz, A., Pourali, M., 2017. Main properties of nanocrystalline hydroxyapatite as a bone graft material in treatment of periodontal defects. A review of literature. Biotechnol. Biotechnol. Equip. 31, 215220. Bayrak, S., Tunc, E.S., Sonmez, I.S., Egilmez, T., Ozmen, B., 2009. Effects of casein phosphopeptide-amorphous calcium phosphate (CPP-ACP) application on enamel microhardness after bleaching. Am. J. Dent. 22, 393396. Baysal, A., Uysal, T., 2012. Do enamel microabrasion and casein phosphopeptide-amorphous calcium phosphate affect shear bond strength of orthodontic brackets bonded to a demineralized enamel surface? Angle Orthod. 82, 3641. Beerens, M.W., van der Veen, M.H., van Beek, H., Ten Cate, J.M., 2010. Effects of casein phosphopeptide amorphous calcium fluoride phosphate paste on white spot lesions and dental plaque after orthodontic treatment: a 3-month follow-up. Eur. J. Oral Sci. 118, 610617. Bellagarda, G., 1965. La carie dentaria e le sue cure negli scritti degli antichi. [Dental caries and their treatment in the writings of the ancients]. Minerva Med. 56, 892903 (in Italian). Benson, A.G., Djalilian, H.R., 2009. Complications of hydroxyapatite bone cement reconstruction of retrosigmoid craniotomy: two cases. Ear Nose Throat J. 88, E1E4. Besinis, A., van Noort, R., Martin, N., 2012. Infiltration of demineralized dentin with silica and hydroxyapatite nanoparticles. Dent. Mater. 28, 10121023. Besinis, A., van Noort, R., Martin, N., 2014. Remineralization potential of fully demineralized dentin infiltrated with silica and hydroxyapatite nanoparticles. Dent. Mater. 30, 249262. Bhatt, N., Naithani, R., Gupta, S.K., 2017. Supersaturated calcium phosphate rinse in prevention and treatment of mucositis in patients undergoing hematopoietic stem cell transplant. Exp. Clin. Transplant. 15, 567570. Bilginer, S., Esener, T., Söylemezoglu, F., Tiftik, A.M., 1997. The investigation of biocompatibility and apical microleakage of tricalcium phosphate based root canal sealers. J. Endod. 23, 105109. Binahmed, A., Stoykewych, A., Hussain, A., Love, B., Pruthi, V., 2007. Long-term follow-up of hydroxyapatite-coated dental implants a clinical trial. Int. J. Oral Maxillofac. Implants 22, 963968. Block, M.S., Kent, J.N., Kay, J.F., 1987. Evaluation of hydroxylapatite-coated titanium dental implants in dogs. J. Oral Maxillofac. Surg. 45, 601607. Block, M.S., Kent, J.N., Beirne, O.R., 1994. Long-term follow-up on hydroxylapatite-coated cylindrical dental implants: a comparison between developmental and recent periods. J. Oral Maxillofac. Surg. 52, 937944.
Calcium orthophosphates as a dental regenerative material 419 Block, M.S., Gardiner, D., Almerico, B., Neal, C., 2000. Loaded hydroxylapatite-coated implants and uncoated titanium-threaded implants in distracted dog alveolar ridges. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 89, 676685. Bobbio, A., 1972. The first endosseous alloplastic implant in the history of man. Bull. Hist. Dent. 20, 16. Boneta, A.E., Neesmith, A., Mankodi, S., Berkowitz, H.J., Sánchez, L., Mostler, K., et al., 2001. The enhanced anticaries efficacy of a sodium fluoride and dicalcium phosphate dihydrate dentifrice in a dual-chambered tube. A 2-year caries clinical study on children in the United States of America. Am. J. Dent. 14 (Spec. Iss. 5), 13A17A. Boone, M.E., Kafrawy, H., 1979. Pulp reaction to a tricalcium phosphate ceramic capping agent. Oral Surg. Oral Med. Oral Pathol. 47, 369371. Bosco, J., Enkel, B., Armengol, V., Daculsi, G., Jean, A., Weiss, P., 2006. Bioactive calcium phosphate material for dental endodontic treatment. Root apical deposition. Key Eng. Mater. 309311, 11571160. Bowen, J.A., Mellonig, J.T., Gray, J.L., Towle, H.T., 1989. Comparison of decalcified freeze-dried bone allograft and porous particulate hydroxyapatite in human periodontal osseous defects. J. Periodontol. 60, 647654. Boyne, P., 1982. Impact of durapatite as a bone grafting material in oral and maxillofacial surgery. Compend. Contin. Educ. Dent. (Suppl. 2), S83S86. Briak, H.E., Durand, D., Nurit, J., Munier, S., Pauvert, B., Boudeville, P., 2002. Study of a hydraulic dicalcium phosphate dihydrate/calcium oxide-based cement for dental applications. J. Biomed. Mater. Res. 63, 447453. Bröchner, A., Christensen, C., Kristensen, B., Tranæus, S., Karlsson, L., Sonnesen, L., et al., 2011. Treatment of post-orthodontic white spot lesions with casein phosphopeptidestabilised amorphous calcium phosphate. Clin. Oral Investig. 15, 369373. Brostow, W., Estevez, M., Lobland, H.E.H., Hoang, L., Rodriguez, J.R., Vargar, S., 2008. Porous hydroxyapatite-based obturation materials for dentistry. J. Mater. Res. 23, 15871596. Brown, G.D., Mealey, B.L., Nummikoski, P.V., Bifano, S.L., Waldrop, T.C., 1998. Hydroxyapatite cement implant for regeneration of periodontal osseous defects in humans. J. Periodontol. 69, 146157. Browning, W.D., Cho, S.D., Deschepper, E.J., 2012. Effect of a nano-hydroxyapatite paste on bleaching-related tooth sensitivity. J. Esthet. Restor. Dent. 24, 268276. Cahyanto, A., Megasari, E., Zakaria, M.N., Djustiana, N., Sunarso, Usri, K., Aripin, D., et al., 2017. Fabrication of carbonate apatite cement as endodontic sealer. Key Eng. Mater. 758, 6165. Cai, F., Shen, P., Morgan, M.V., Reynolds, E.C., 2003. Remineralization of enamel subsurface lesions in situ by sugar-free lozenges containing casein phosphopeptide-amorphous calcium phosphate. Aust. Dent. J. 48, 240243. Cai, F., Manton, D.J., Shen, P., Walker, G.D., Cross, K.J., Yuan, Y., et al., 2007. Effect of addition of citric acid and casein phosphopeptide-amorphous calcium phosphate to a sugar-free chewing gum on enamel remineralization in situ. Caries Res. 41, 377383. Cãlin, C., Pãtraşcu, I., 2016. Growth factors and beta-tricalcium phosphate in the treatment of periodontal intraosseous defects: a systematic review and meta-analysis of randomised controlled trials. Arch. Oral Biol. 66, 4454. Cao, C.Y., Mei, M.L., Li, Q.L., Lo, E.C., Chu, C.H., 2014. Agarose hydrogel biomimetic mineralization model for the regeneration of enamel prism-like tissue. ACS Appl. Mater. Interfaces 6, 410420.
420 Advanced Dental Biomaterials Cao, C.Y., Mei, M.L., Li, Q.L., Lo, E.C.M., Chu, C.H., 2015. Methods for biomimetic mineralisation of human enamel: a systematic review. Materials 8, 28732886. Çehreli, S.B., Şar, C., Polat-Özsoy, O., Ünver, B., Özsoy, S., 2012. Effects of a fluoridecontaining casein phosphopeptide-amorphous calcium phosphate complex on the shear bond strength of orthodontic brackets. Eur. J. Orthod. 34, 193197. Chai, Y., Slavkin, H.C., 2003. Prospects for tooth regeneration in the 21st century: a perspective. Microsc. Res. Tech. 60, 469479. Chang, S.W., Lee, S.Y., Kang, S.K., Kum, K.Y., Kim, E.C., 2014. In vitro biocompatibility, inflammatory response, and osteogenic potential of 4 root canal sealers: Sealapex, Sankin apatite root sealer, MTA Fillapex, and iRoot SP root canal sealer. J. Endod. 40, 16421648. Chang, S.W., Lee, Y.K., Zhu, Q., Shon, W.J., Lee, W.C., Kum, K.Y., et al., 2015. Comparison of the rheological properties of four root canal sealers. Int. J. Oral Sci. 7, 5661. Chau, J.Y.M., Hutter, J.W., Mork, T.O., Nicoll, B.K., 1997. An in vitro study of furcation perforation repair using calcium phosphate cement. J. Endod. 23, 588592. Chaung, H.M., Hong, C.H., Chiang, C.P., Lin, S.K., Kuo, Y.S., Lan, W.H., et al., 1996. Comparison of calcium phosphate cement mixture and pure calcium hydroxide as direct pulp-capping agents. J. Formos. Med. Assoc. 95, 545550. Checchi, V., Savarino, L., Montevecchi, M., Felice, P., Checchi, L., 2011. Clinicalradiographic and histological evaluation of two hydroxyapatites in human extraction sockets: a pilot study. Int. J. Oral Maxillofac. Surg. 40, 526532. Chen, M.H., 2010. Update on dental nanocomposites. J. Dent. Res. 89, 549560. Chen, H., Clarkson, B.H., Sun, K., Mansfield, J.F., 2005. Self-assembly of synthetic hydroxyapatite nanorods into an enamel prism-like structure. J. Colloid Interface. Sci. 288, 97103. Chen, F.M., Zhang, J., Zhang, M., An, Y., Chen, F., Wu, Z.F., 2010. A review on endogenous regenerative technology in periodontal regenerative medicine. Biomaterials 31, 78927927. Chen, Y.Z., Lü, X.Y., Liu, G.D., 2013a. A novel root-end filling material based on hydroxyapatite, tetracalcium phosphate and polyacrylic acid. Int. Endod. J. 46, 556564. Chen, L., Liang, K., Li, J., Wu, D., Zhou, X., Li, J., 2013b. Regeneration of biomimetic hydroxyapatite on etched human enamel by anionic PAMAM template in vitro. Arch. Oral Biol. 58, 975980. Cherng, A.M., Chow, L.C., Takagi, S., 2001. In vitro evaluation of a calcium phosphate cement root canal filler/sealer. J. Endod. 27, 613615. Cherng, A.M., Takagi, S., Chow, L.C., 2010. Acid neutralization capacity of a tricalcium silicate-containing calcium phosphate cement as an endodontic material. J. Res. Natl. Inst. Stand. Technol. 115, 471476. Chiapasco, M., Casentini, P., Zaniboni, M., 2009. Bone augmentation procedures in implant dentistry. Int. J. Oral Maxillofac. Implants 24 (Suppl. 2), 3759. Chitsazi, M.T., Shirmohammadi, A., Faramarzie, M., Pourabbas, R., Rostamzadeh, A.N., 2011. A clinical comparison of nano-crystalline hydroxyapatite (Ostim) and autogenous bone graft in the treatment of periodontal intrabony defects. Med. Oral Patol. Oral Cir. Bucal 16, 448453. Chohayeb, A.A., Chow, L.C., Tsaknis, P., 1987. Evaluation of calcium phosphate as a root canal sealer-filler material. J. Endod. 13, 384387. Chow, L.C., Takagi, S., 1995. Remineralization of root lesions with concentrated calcium and phosphate solutions. Dent. Mater. J. 14, 3136.
Calcium orthophosphates as a dental regenerative material 421 Chow, L.C., Takagi, S., Shern, R.J., Chow, T.H., Takagi, K.K., Sieck, B.A., 1994. Effects on whole saliva of chewing gums containing calcium phosphates. J. Dent. Res. 73, 2632. Chow, C.K.W., Wu, C.D., Evans, C.A., 2011. In vitro properties of orthodontic adhesives with fluoride or amorphous calcium phosphate. Int. J. Dent. 2011, 583521 (8 pp.). Cochrane, N.J., Cai, F., Huq, N.L., Burrow, M.F., Reynolds, E.C., 2010. New approaches to enhanced remineralization of tooth enamel. J. Dent. Res. 89, 11871197. Collares, F.M., Leitune, V.C.B., Rostirolla, F.V., Trommer, R.M., Bergmann, C.P., Samuel, S.M.W., 2012. Nanostructured hydroxyapatite as filler for methacrylate-based root canal sealers. Int. Endod. J. 45, 6367. Comar, L.P., Souza, B.M., Gracindo, L.F., Buzalaf, M.A.R., Magalhães, A.C., 2013. Impact of experimental nano-HAP pastes on bovine enamel and dentin submitted to a pH cycling model. Braz. Dent. J. 24, 273278. Comuzzi, L., Ooms, E., Jansen, J.A., 2002. Injectable calcium phosphate cement as a filler for bone defects around oral implants, an experimental study in goats. Clin. Oral Implants Res. 13, 304311. Corbett, M.E., Moore, W.J., 1971. Distribution of dental caries in ancient British populations. 1. Anglo-Saxon period. Caries Res. 5, 151168. Corbett, M.E., Moore, W.J., 1973. Distribution of dental caries in ancient British populations. II. Iron age, Romano-British and Mediaeval periods. Caries Res. 7, 139153. Corbett, M.E., Moore, W.J., 1975. Distribution of dental caries in ancient British populations. III. The 17th century. Caries Res. 9, 163175. Corbett, M.E., Moore, W.J., 1976. Distribution of dental caries in ancient British populations. IV. The 19th century. Caries Res. 10, 401414. Cottrell, D.A., Wolford, L.M., 1998. Long-term evaluation of the use of coralline hydroxyapatite in orthognathic surgery. J. Oral Maxillofac. Surg. 56, 935941. discussion 941942. Cranin, A.N., Tobin, G.P., Glebman, J., 1987a. Applications of hydroxylapatite in oral and maxillofacial surgery. Part I: Periodontal and endosteal-implant repairs. Compendium (Newtown, PA) 8, 254256. 258, 261. Cranin, A.N., Tobin, G.P., Gelbman, J., 1987b. Applications of hydroxylapatite in oral and maxillofacial surgery. Part II: Ridge augmentation and repair of major oral defects. Compendium (Newtown, PA) 8, 334335. 337. Cravens, J.E., 1876. Lacto-phosphate of lime; pathology and treatment of exposed dental pulps and sensitive dentine. Dent. Cosmos 18, 463469. discussion 469476. Cross, K.J., Huq, N.L., Reynolds, E.C., 2016. Casein phosphopeptideamorphous calcium phosphate nanocomplexes: a structural model. Biochemistry 55, 43164325. Crubezy, E., Murail, P., Girard, L., Bernadou, J.P., 1998. False teeth of the Roman world. Nature 391, 29. Daculsi, G., Weiss, P., Bouler, J.M., Gauthier, O., Millot, F., Aguado, E., 1999. Biphasic calcium phosphate/hydrosoluble polymer composites: a new concept for bone and dental substitution biomaterials. Bone 25 (Suppl. 1), 59S61S. Dan, H., Vaquette, C., Fisher, A.G., Hamlet, S.M., Xiao, Y., Hutmacher, D.W., et al., 2014. The influence of cellular source on periodontal regeneration using calcium phosphate coated polycaprolactone scaffold supported cell sheets. Biomaterials 35, 113122. de Carvalho, F.G., Vieira, B.R., Santos, R.L., Carlo, H.L., Lopes, P.Q., de Lima, B.A., 2014. In vitro effects of nano-hydroxyapatite paste on initial enamel carious lesions. Pediatr. Dent. 36, 8589. Denissen, H.W., de Groot, K., 1979. Immediate dental root implants from synthetic dense calcium hydroxylapatite. J. Prost. Dent. 42, 551556.
422 Advanced Dental Biomaterials Denissen, H.W., Kalk, W., Veldhuis, A.A.H., van den Hooff, A., 1989. Eleven-year study of hydroxyapatite implants. J. Prost. Dent. 61, 706712. Dickens, S.H., Flaim, G.M., Takagi, S., 2003. Mechanical properties and biochemical activity of remineralizing resin-based CaPO4 cements. Dent. Mater. 19, 558566. Dickens-Venz, S.H., Takagi, S., Chow, L.C., Bowen, R.L., Johnston, A.D., Dickens, B., 1994. Physical and chemical properties of resin-reinforced calcium phosphate cements. Dent. Mater. 10, 100106. D’Lima, J.P., Paul, J., Palathingal, P., Varma, B.R.R., Bhat, M., Mohanty, M., 2014. Histological and histometrical evaluation of two synthetic hydroxyapatite based biomaterials in the experimental periodontal defects in dogs. J. Clin. Diagn. Res. 8, ZC52ZC55. Dodds, M.W., Chidichimo, D., Haas, M.S., 2012. Delivery of active agents from chewing gum for improved remineralization. Adv. Dent. Res. 24, 5862. Domingo, C., Arcı́s, R.W., López-Macipe, A., Osorio, R., Rodrı́guez-Clemente, R., Murtra, J., et al., 2001. Dental composites reinforced with hydroxyapatite: mechanical behavior and absorption/elution characteristics. J. Biomed. Mater. Res. 56, 297305. Dorozhkin, S.V., 2012a. Calcium Orthophosphates: Applications in Nature, Biology, and Medicine. Pan Stanford, Singapore, 854 pp. Dorozhkin, S.V., 2012b. Dissolution mechanism of calcium apatites in acids: a review of literature. World J. Methodol. 2, 117. Dorozhkin, S.V., 2015. Calcium orthophosphates bioceramics. Ceram. Int. 41, 1391313966. Dorozhkin, S.V., 2016a. Multiphasic calcium orthophosphate (CaPO4) bioceramics and their biomedical applications. Ceram. Int. 42, 65296554. Dorozhkin, S.V., 2016b. Calcium orthophosphates (CaPO4): occurrence and properties. Prog. Biomater. 5, 970. Dorozhkin, S.V., 2016c. Calcium Orthophosphate-Based Bioceramics and Biocomposites. Wiley-VCH, Weinheim, Germany, 405 pp. Dou, W., Mastrogiacomo, S., Veltien, A., Alghamdi, H.S., Walboomers, X.F., Heerschap, A., 2017. Visualization of calcium phosphate cement in teeth by zero echo time 1H MRI at high field. NMR Biomed. 31, e3859 (12 pp.). Dowd, F.J., 1999. Saliva and dental caries. Dent. Clin. North Am. 43, 579597. Duailibi, S.E., Duailibi, M.T., Zhang, W., Asrican, R., Vacanti, J.P., Yelick, P.C., 2008. Bioengineered dental tissues grown in the rat jaw. J. Dent. Res. 87, 745750. Dunn, W.J., 2007. Shear bond strength of an amorphous calcium-phosphate-containing orthodontic resin cement. Am. J. Orthod. Dentofac. Orthoped. 131, 243247. Dupoirieux, L., Gard, C., 2000. Hydroxyapatite cement for calvarial reconstruction. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 89, 140142. Eccles, J.D., 1979. Dental erosion of nonindustrial origin. A clinical survey and classification. J. Prosthet. Dent. 42, 649653. Elangovan, S., Jain, S., Tsai, P.C., Margolis, H.C., Amiji, M., 2013. Nano-sized calcium phosphate particles for periodontal gene therapy. J. Periodontol. 84, 117125. Elgendy, E.A., Shady, T.E.A., 2015. Clinical and radiographic evaluation of nanocrystalline hydroxyapatite with or without platelet-rich fibrin membrane in the treatment of periodontal intrabony defects. J. Indian Soc. Periodontol. 19, 6165. Ellinger, R.F., Nery, E.B., Lynch, K.L., 1986. Histological assessment of periodontal osseous defects following implantation of hydroxyapatite and biphasic calcium phosphate ceramics: a case report. Int. J. Periodontics Restorative Dent. 6, 2233. Elliott, J.C., 1994. Structure and chemistry of the apatites and other calcium orthophosphates, Studies in Inorganic Chemistry, Vol. 18. Elsevier, Amsterdam, The Netherlands, 389 pp.
Calcium orthophosphates as a dental regenerative material 423 Emamieh, S., Khaterizadeh, Y., Goudarzi, H., Ghasemi, A., Baghban, A.A., Torabzadeh, H., 2015. The effect of two types chewing gum containing casein phosphopeptideamorphous calcium phosphate and xylitol on salivary Streptococcus mutans. J. Conserv. Dent. 18, 192195. Emerton, K.B., Drapeau, S.J., Prasad, H., Rohrer, M., Roffe, P., Hopper, K., et al., 2011. Regeneration of periodontal tissues in non-human primates with rhGDF-5 and betatricalcium phosphate. J. Dent. Res. 90, 14161421. Enan, E.T., Hammad, S.M., 2013. Microleakage under orthodontic bands cemented with nano-hydroxyapatite-modified glass ionomer: an in vivo study. Angle Orthod. 83, 981986. Enax, J., Epple, M., 2018. Synthetic hydroxyapatite as a biomimetic oral care agent. Oral Health Prev. Dent. 16, 719. Esteves-Oliveira, M., Santos, N.M., Meyer-Lueckel, H., Wierichs, R.J., Rodrigues, J.A., 2017. Caries-preventive effect of anti-erosive and nano-hydroxyapatite-containing toothpastes in vitro. Clin. Oral. Investig. 21, 291300. Fabregas, R., Rubinstein, J., 2014a. On the initial propagation of dental caries. J. R. Soc. Interface 11, 20140809. Fabregas, R., Rubinstein, J., 2014b. A mathematical model for the progression of dental caries. Math. Med. Biol. 31, 319337. Fathi, M.H., Salehi, M., Mortazavi, V., Mousavi, S.B., Parsapour, A., 2006. Novel hydroxyapatite/niobium surface coating for endodontic dental implant. Surf. Eng. 22, 353358. Fejerskov, O., Kidd, E. (Eds.), 2008. Dental Caries: The Disease and Its Clinical Management. second ed. Wiley-Blackwell, 640 pp. Fernández, M.P.R., Gehrke, S.A., Mazón, P., Calvo-Guirado, J.L., de Aza, P.N., 2017. Implant stability of biological hydroxyapatites used in dentistry. Materials 10, 644. Finn, S.B., 1967. The role of a dicalcium phosphate chewing gum in the control of human dental caries. Int. Dent. J. 17, 339352. Finn, S.B., Jamison, H.C., 1967. The effect of a dicalcium phosphate chewing gum on caries incidence in children: 30-month results. J. Am. Dent. Assoc. 74, 987995. Fischer, R.B., Muhler, J.C., Ring, C.E., 1956. X-ray study of fluorapatite formation during the fluoride treatment of powdered dental enamel. J. Dent. Res. 35, 773777. Fischer-Brandies, E., Dielert, E., 1985. Clinical use of tricalciumphosphate and hydroxyapatite in maxillofacial surgery. J. Oral Implant. 12, 4044. Fletcher, J., Walsh, D., Fowler, C.E., Mann, S., 2011. Electrospun mats of PVP/ACP nanofibres for remineralization of enamel tooth surfaces. CrystEngComm 13, 36923697. Foster, J.A., Berzins, D.W., Bradley, T.G., 2008. Bond strength of an amorphous calcium phosphate-containing orthodontic adhesive. Angle Orthod. 78, 339344. Frame, J.W., 1987. Hydroxyapatite as a biomaterial for alveolar ridge augmentation. Int. J. Oral Maxillofac. Surg. 16, 642655. Frame, J.W., Brady, C.L., 1987. The versatility of hydroxyapatite blocks in maxillofacial surgery. Br. J. Oral Maxillofac. Surg. 25, 452464. Frame, J.W., Brady, C.L., Browne, R.M., 1981. Augmentation of the edentulous mandible using bone and hydroxyapatite: a comparative study in dogs. Int. J. Oral Surg. 10, 8892. Friedman, C.D., Costantino, P.D., Jones, K., Chow, L.C., Pelzer, H., Sisson, G., 1991. Hydroxyapatite cement. II. Obliteration and reconstruction of the cat frontal sinus. Arch. Otolaryngol. Head Neck Surg. 117, 385389. Friedman, C.D., Costantino, P.D., Takagi, S., Chow, L.C., 1998. BoneSourcet hydroxyapatite cement, a novel biomaterial for craniofacial skeletal tissue engineering and reconstruction. J. Biomed. Mater. Res. 43, 428432.
424 Advanced Dental Biomaterials Friedman, C.D., Costantino, P.D., Synderman, C.H., Chow, L.C., Takagi, S., 2000. Reconstruction of the frontal sinus and frontofacial skeleton with hydroxyapatite cement. Arch. Facial Plast. Surg. 2, 124129. Fujikawa, K., Sugawara, A., Kusama, K., Nishiyama, M., Murai, S., Takagi, S., et al., 2002. Fluorescent labeling analysis and electron probe microanalysis for alveolar ridge augmentation using calcium phosphate cement. Dent. Mater. J. 21, 296305. Fujita, H., 2009. Dental caries in Japanese human skeletal remains. J. Oral Biosci 51, 105114. Gaberthüel, T.W., Strub, J.R., 1977. Treatment of periodontal pockets with tricalcium phosphate in man. A preliminary report. SSO Schweiz Monatsschr. Zahnheilkd. 87, 809814. Gaetti-Jardim, E.C., Santiago-Junior, J.F., Goiato, M.C., Pellizer, E.P., Magro-Filho, O., Jardim Jr., E.G., 2011. Dental implants in patients with osteoporosis: a clinical reality? J. Craniofac. Surg. 22, 11111113. Gaffar, A., Blake-Haskins, J., Mellberg, J., 1993. In vivo studies with a dicalcium phosphate dihydrate/MFP system for caries prevention. Int. Dent. J 43 (Suppl. 1), 8188. Gambarini, G., Tagger, M., 1996. Sealing ability of a new hydroxyapatite-containing endodontic sealer using lateral condensation and thermatic compaction of gutta-percha, in vitro. J. Endod. 22, 165167. Garcı́a, D., Garcı́a, L., Pérez, M.P., Suarez, M., Delgado, J.A., Garcı́a, R., et al., 2001. Filling of post-extraction dental socket with hydroxyapatite granules APAFILL-Gt. Key Eng. Mater. 192195, 925928. Geiger, S., Matalon, S., Blasbalg, J., Tung, M.S., Eichmiller, F.C., 2003. The clinical effect of amorphous calcium phosphate (ACP) on root surface hypersensitivity. Oper. Dent. 28, 496500. Gineste, L., Gineste, M., Ranz, X., Ellefterion, A., Guilhem, A., Rouquet, N., et al., 1999. Degradation of hydroxylapatite, fluorapatite, and fluorhydroxyapatite coatings of dental implants in dogs. J. Biomed. Mater. Res. 48, 224234. Giniger, M., MacDonald, J., Spaid, M., Felix, H., 2005a. A 180-day clinical investigation of the tooth whitening efficacy of a bleaching gel with added amorphous calcium phosphate. J. Clin. Dent. 16, 1116. Giniger, M., MacDonald, J., Ziemba, S., Felix, H., 2005b. The clinical performance of professionally dispensed bleaching gel with added amorphous calcium phosphate. J. Am. Dent. Assoc. 136, 383392. Giordano, M., Macchi, A., Ostinelli, E., Tagliabue, A., 1996. Effetto protettivo sullo smalto della idrossilapatite ultramicronizzata addizionata ad un cemento composito ortodontico. Studio in vivo. [Protective effect on enamel of ultra-micronized hydroxyapatite in combination with orthodontic composite cement. In vivo study]. Minerva Stomatol. 45, 2935. Goh, B.T., Chanchareonsook, N., Tideman, H., Teoh, S.H., Chow, J.K., Jansen, J.A., 2014. The use of a polycaprolactone-tricalcium phosphate scaffold for bone regeneration of tooth socket facial wall defects and simultaneous immediate dental implant placement in Macaca fascicularis. J. Biomed. Mater. Res. A 102A, 13791388. Gomes, Y.S.L., Alexandrino, L.D., Alencar, C.M., Alves, E.B., Faial, K.C., Silva, C.M., 2017. In situ effect of nanohydroxyapatite paste in enamel teeth bleaching. J. Contemp. Dent. Pract. 18, 9961003. González-Cabezas, C., 2010. The chemistry of caries: remineralization and demineralization events with direct clinical relevance. Dent. Clin. North. Am. 54, 469478.
Calcium orthophosphates as a dental regenerative material 425 Gu, H., Mijares, D., Zhao, Z., Boylan, R., Ling, J., LeGeros, R.Z., 2013. Experimental antibacterial and mineralizing calcium phosphate-based treatment for dentin surfaces. J. Biomater. Appl. 27, 783790. Gupta, R., Prakash, V., 2011. CPP-ACP complex as a new adjunctive agent for remineralisation: a review. Oral Health Prev. Dent. 9, 151165. Gurunathan, D., Somasundaram, S., 2015. Prevention of white spot lesion in orthodontic patients using casein phosphopeptide-stabilized amorphous calcium phosphate a systematic review. Int. J. Pharma Bio Sci. 6, B702B706. Haghgoo, R., Rezvani, M.B., Zeinabadi, M.S., 2014. Comparison of nano-hydroxyapatite and sodium fluoride mouthrinse for remineralization of incipient carious lesions. J. Dent. (Tehran) 11, 406410. Hallman, M., Hedin, M., Sennerby, L., Lundgren, S., 2002. A prospective 1-year clinical and radiographic study of implants placed after maxillary sinus floor augmentation with bovine hydroxyapatite and autogenous bone. J. Oral Maxillofac. Surg. 60, 277284. Hamba, H., Nikaido, T., Inoue, G., Sadr, A., Tagami, J., 2011. Effects of CPP-ACP with sodium fluoride on inhibition of bovine enamel demineralization: a quantitative assessment using micro-computed tomography. J. Dent. 39, 405413. Hammad, S.M., El Banna, M.S., Elsaka, S.E., 2013. Twelve-month bracket failure rate with amorphous calcium phosphate bonding system. Eur. J. Orthod. 35, 622627. Hannig, C., Basche, S., Burghardt, T., Al-Ahmad, A., Hannig, M., 2013. Influence of a mouthwash containing hydroxyapatite microclusters on bacterial adherence in situ. Clin. Oral Investig. 17, 805814. Hara, Y., Murakami, T., Kajiyama, K., Maeda, K., Akamine, A., Nagamine, N., et al., 1989. Application of calcium phosphate ceramics to periodontal therapy. 8. Effects of orthodontic force on repaired bone with hydroxyapatite. Nihon Shishubyo Gakkai Kaishi 31, 224234. Harbert, H., 1991. Generic tricalcium phosphate plugs: an adjunct in endodontics. J. Endod. 17, 131134. Hasnamudhia, F., Bachtiar, E.W., Sahlan, M., Soekanto, S.A., 2017. The effect of CPP-ACPpropolis chewing gum on calcium and phosphate ion release on caries-active subjects’ saliva and the formation of Streptococcus mutans biofilm. J. Phys. Conf. Ser. 884, 012137. Hayashi, Y., Imai, M., Yanagiguchi, K., Viloria, I.L., Ikeda, T., 1999. Hydroxyapatite applied as direct pulp capping medicine substituted for osteodentin. J. Endod. 25, 225229. Hayashi, C., Kinoshita, A., Oda, S., Mizutani, K., Shirakata, Y., Ishikawa, I., 2006. Injectable calcium phosphate bone cement provides favorable space and a scaffold for periodontal regeneration in dogs. J. Periodontol. 77, 940946. Head, J.A., 1912. A study of saliva and its action on tooth enamel in reference to its hardening and softening. J. Am. Med. Assoc. 59, 21182122. Hegde, M., Moany, A., 2012. Remineralization of enamel subsurface lesions with casein phosphopeptide-amorphous calcium phosphate: a quantitative energy dispersive X-ray analysis using scanning electron microscopy: an in vitro study. J. Conserv. Dent. 15, 6167. Heinz, B., Kasaj, A., Teich, M., Jepsen, S., 2010. Clinical effects of nanocrystalline hydroxyapatite paste in the treatment of intrabony periodontal defects: a randomized controlled clinical study. Clin. Oral Investig. 14, 525531. Heller, A.I., Koenigs, J.F., Brilliant, J.D., Melfi, R.C., Driskell, T.D., 1975. Direct pulp capping of permanent teeth in primates using a resorbable form of tricalcium phosphate ceramic. J. Endod. 1, 95101.
426 Advanced Dental Biomaterials Hemmerle, J., Leize, M., Voegel, J.C., 1995. Long-term behaviour of a hydroxyapatite/collagen-glycosaminoglycan biomaterial used for oral surgery: a case report. J. Mater. Sci. Mater. Med. 6, 360366. Hicks, M.J., Flaitz, C.M., 2000. Enamel caries formation and lesion progression with a fluoride dentifrice and a calcium-phosphate containing fluoride dentifrice: a polarized light microscopic study. ASDC J. Dent. Child. 67, 2128. Higashi, T., 1996. Influence of particle size of calcium phosphate ceramics as a capping agent on the formation of a hard tissue barrier in amputated dental pulp. J. Endod. 22, 281283. Hill, R.G., Gillam, D.G., Chen, X., 2015. The ability of a nano hydroxyapatite toothpaste and oral rinse containing fluoride to protect enamel during an acid challenge using 19F solid state NMR spectroscopy. Mater. Lett. 156, 6971. Hiller, K.A., Buchalla, W., Grillmeier, I., Neubauer, C., Schmalz, G., 2018. In vitro effects of hydroxyapatite containing toothpastes on dentin permeability after multiple applications and ageing. Sci. Rep. 8, 4888. Himel, V., Brady, J., Weir, J., 1985. Evaluation of repair of mechanical perforations of the pulp chamber floor using biodegradable tricalcium phosphate or calcium hydroxide. J. Endod. 11, 161165. Höland, W., Schweiger, M., Watzke, R., Peschke, A., Kappert, H., 2008. Ceramics as biomaterials for dental restoration. Expert Rev. Med. Dev 5, 729745. Holmes, R.E., Hagler, H.K., 1987. Porous hydroxylapatite as a bone graft substitute in mandibular contour augmentation: a histometric study. J. Oral Maxillofac. Surg. 45, 421429. Hontsu, S., Kato, N., Yamamoto, E., Nishikawa, H., Kusunoki, M., Hayami, T., et al., 2012. Regeneration of tooth enamel by flexible hydroxyapatite sheet. Key Eng. Mater. 493494, 615619. Horch, H.H., Sader, R., Pautke, C., Neff, A., Deppe, H., Kolk, A., 2006. Synthetic, purephase beta-tricalcium phosphate ceramic granules (Cerasorbs) for bone regeneration in the reconstructive surgery of the jaws. Int. J. Oral Maxillofac. Surg. 35, 708713. Horst, O.V., Chavez, M.G., Jheon, A.H., Desai, T., Klein, O.D., 2012. Stem cell and biomaterials research in dental tissue engineering and regeneration. Dent. Clin. North Am. 56, 495520. Horváth, A., Stavropoulos, A., Windisch, P., Lukács, L., Gera, I., Sculean, A., 2013. Histological evaluation of human intrabony periodontal defects treated with an unsintered nanocrystalline hydroxyapatite paste. Clin. Oral Investig. 17, 423430. Hu, B., Nadiri, A., Kuchler-Bopp, S., Perrin-Schmitt, F., Peters, H., Lesot, H., 2006. Tissue engineering of tooth crown, root, and periodontium. Tissue Eng. 12, 20692075. Huang, S.B., Gao, S.S., Yu, H.Y., 2009. Effect of nano-hydroxyapatite concentration on remineralization of initial enamel lesion in vitro. Biomed. Mater. 4, 034104. Huang, M.S., Wu, H.D., Teng, N.C., Peng, B.Y., Wu, J.Y., Chang, W.J., et al., 2010a. In vivo evaluation of poorly crystalline hydroxyapatite-based biphasic calcium phosphate bone substitutes for treating dental bony defects. J. Dent. Sci. 5, 100108. Huang, S., Gao, S., Cheng, L., Yu, H., 2010b. Combined effects of nano-hydroxyapatite and Galla chinensis on remineralisation of initial enamel lesion in vitro. J. Dent. 38, 811819. Huang, S., Gao, S., Cheng, L., Yu, H., 2011. Remineralization potential of nanohydroxyapatite on initial enamel lesions: an in vitro study. Caries Res. 45, 460468. Hulshoff, J.E.G., Jansen, J.A., 1997. Initial interfacial healing events around calcium phosphate (Ca-P) coated oral implants. Clin. Oral Implant. Res. 8, 393400.
Calcium orthophosphates as a dental regenerative material 427 Hüttemann, R.W., Dönges, H., 1987. Untersuchungen zur Therapie überempfindlicher Zahnhälse mit Hydroxylapatit. Dtsch. Zahnärztl. Z 42, 486488. Iezzi, G., Scarano, A., Petrone, G., Piattelli, A., 2007. Two human hydroxyapatite-coated dental implants retrieved after a 14-year loading period: a histologic and histomorphometric case report. J. Periodontol. 78, 940947. Ignatius, A.A., Ohnmacht, M., Claes, L.E., Kreidler, J., Palm, F., 2001. A composite polymer/tricalcium phosphate membrane for guided bone regeneration in maxillofacial surgery. J. Biomed. Mater. Res. 58, 564569. Iijima, Y., Cai, F., Shen, P., Walker, G., Reynolds, C., Reynolds, E.C., 2004. Acid resistance of enamel subsurface lesions remineralized by a sugar-free chewing gum containing casein phosphopeptide-amorphous calcium phosphate. Caries Res. 38, 551556. Ikeda, E., Morita, R., Nakao, K., Ishida, K., Nakamura, T., Takano-Yamamoto, T., et al., 2009. Fully functional bioengineered tooth replacement as an organ replacement therapy. Proc. Natl. Acad. Sci. U.S.A. 106, 1347513480. Ilie, O., van Loosdrecht, M.C.M., Picioreanu, C., 2012. Mathematical modelling of tooth demineralisation and pH profiles in dental plaque. J. Theor. Biol. 309, 159175. Imamura, Y., Tanaka, Y., Nagai, A., Yamashita, K., Takagi, Y., 2010. Self-sealing ability of OCPmediated cement as a deciduous root canal filling material. Dent. Mater. J. 29, 582588. Ishida, H., Nahara, Y., Hamada, T., 1992. Dimensional accuracy of castable apatite ceramic crowns: the influence of heat treatment on dimensional changes and distortion of crowns. J. Prost. Dent. 68, 279283. Itthagarun, A., King, N.M., Yiu, C., Dawes, C., 2005. The effect of chewing gums containing calcium phosphates on the remineralization of artificial caries-like lesions in situ. Caries Res. 39, 251254. Jaber, L., Mascres, C., Donohue, W.B., 1991. Electron microscope characteristics of dentin repair after hydroxyapatite direct pulp capping in rats. J. Oral. Pathol. Med. 20, 502508. Jaber, L., Mascrès, C., Donohue, W.B., 1992. Reaction of the dental pulp to hydroxyapatite. Oral Surg. Oral Med. Oral Pathol. 73, 9298. Jandt, K.D., 2006. Probing the future in functional soft drinks on the nanometre scale towards tooth friendly soft drinks. Trends Food Sci. Technol. 17, 263271. Jayakumar, A., Rajababu, P., Rohini, S., Butchibabu, K., Naveen, A., Reddy, P.K., et al., 2011. Multi-centre, randomized clinical trial on the efficacy and safety of recombinant human platelet-derived growth factor with β-tricalcium phosphate in human intraosseous periodontal defects. J. Clin. Periodontol. 38, 163172. Jean, A.H., Pouezat, J.A., Daculsi, G., 1993. Pulpal response to calcium phosphate materials. In vivo study of calcium phosphate materials in endodontics. Cell. Mater. 3, 193200. Jones, J.D., Saigusa, M., van Sickels, J.E., Tiner, B.D., Gardner, W.A., 1997. Clinical evaluation of hydroxyapatite-coated titanium plasma-sprayed and titanium plasma-sprayed cylinder dental implants. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 84, 137141. Jones, J.D., Lupori, J., van Sickels, J.E., Gardner, W., 1999. A 5-year comparison of hydroxyapatite-coated titanium plasma-sprayed and titanium plasma-sprayed cylinder dental implants. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 87, 649652. Jeong, S.H., Jang, S.O., Kim, K.N., Kwon, H.K., Park, Y.D., Kim, B.I., 2006. Remineralization potential of new toothpaste containing nano-hydroxyapatite. Key Eng. Mater. 309311, 537540. Jeong, S.H., Hong, S.J., Choi, C.H., Kim, B.I., 2007. Effect of new dentifrice containing nanosized carbonated apatite on enamel remineralization. Key Eng. Mater. 330332, 291294.
428 Advanced Dental Biomaterials Joo, H.J., Park, Y.G., 2007. Friction of calcium phosphate brackets to stainless steel wire. Korean J. Orthod. 37, 376385. Jung, U.W., Lee, J.S., Park, W.Y., Cha, J.K., Hwang, J.W., Park, J.C., et al., 2011. Periodontal regenerative effect of a bovine hydroxyapatite/collagen block in one-wall intrabony defects in dogs: a histometric analysis. J. Periodont. Implant Sci. 41, 285292. Jung, U.W., Hwang, J.W., Choi, D.Y., Hu, K.S., Kwon, M.K., Choi, S.H., et al., 2012. Surface characteristics of a novel hydroxyapatite-coated dental implant. J. Periodont. Implant Sci 42, 5963. Junker, R., Manders, P.J.D., Wolke, J., Borisov, Y., Braceras, I., Jansen, J.A., 2011. Bone reaction adjacent to microplasma-sprayed calcium phosphate-coated oral implants subjected to an occlusal load, an experimental study in the dog. Clin. Oral Implant. Res. 22, 135142. Kamakura, S., Sasano, Y., Nakamura, M., Suzuki, O., Ohki, H., Kagayama, M., et al., 1996. Initiation of alveolar ridge augmentation in the rat mandible by subperiosteal implantation of octacalcium phosphate. Arch. Oral Biol. 41, 10291038. Kani, T., Kani, M., Isozaki, A., Kato, H., Fukuoka, Y., Ohashi, T., et al., 1988. The effect to apatite-containing dentifrices on artificial caries lesions. J. Dent. Health 38, 364365. Kani, T., Kani, M., Isozaki, A., Shimatani, H., Ohashi, T., Tokumoto, T., 1989. Effect of apatite-containing dentifrices on dental caries in school children. J. Dent. Health 39, 104109. Kano, T., Yamamoto, R., Miyashita, A., Komatsu, K., Hayakawa, T., Sato, M., et al., 2012. Regeneration of periodontal ligament for apatite-coated tooth-shaped titanium implants with and without occlusion using rat molar model. J. Hard Tissue Biol. 21, 189202. Karlinsey, R.L., Pfarrer, A.M., 2012. Fluoride plus functionalized β-TCP: a promising combination for robust remineralization. Adv. Dent. Res. 24, 4852. Karlinsey, R.L., Mackey, A.C., Walker, E.R., Frederick, K.E., 2010a. Surfactant-modified β-TCP: structure, properties, and in vitro remineralization of subsurface enamel lesions. J. Mater. Sci. Mater. Med. 21, 20092020. Karlinsey, R.L., Mackey, A.C., Walker, E.R., Frederick, K.E., 2010b. Preparation, characterization and in vitro efficacy of an acid-modified β-TCP material for dental hard-tissue remineralization. Acta Biomater. 6, 969978. Kato, N., Yamamoto, E., Isai, A., Nishikawa, H., Kusunoki, M., Yoshikawa, K., et al., 2013. Ultrathin amorphous calcium phosphate freestanding sheet for dentin tubule sealing. Bioceram. Dev. Appl. S1, 007. Kato, N., Isai, A., Yamamoto, E., Nishikawa, H., Kusunoki, M., Yoshikawa, K., et al., 2015. Evaluation of dentin tubule sealing rate improved by attaching ultrathin amorphous calcium phosphate sheet. Key Eng. Mater. 631, 258261. Kaushick, B.T., Jayakumar, N.D., Padmalatha, O., Varghese, S., 2011. Treatment of human periodontal infrabony defects with hydroxyapatite 1 β tricalcium phosphate bone graft alone and in combination with platelet rich plasma: a randomized clinical trial. Indian J. Dent. Res. 22, 505510. Kawai, T., Matsui, K., Ezoe, Y., Kajii, F., Suzuki, O., Takahashi, T., et al., 2018. Efficacy of octacalcium phosphate collagen composite for titanium dental implants in dogs. Materials 11, 229. Kawase, T., Okuda, K., Kogami, H., Nakayama, H., Nagata, M., Sato, T., et al., 2010. Human periosteum-derived cells combined with superporous hydroxyapatite blocks used as an osteogenic bone substitute for periodontal regenerative therapy: an animal implantation study using nude mice. J. Periodontol. 81, 420427.
Calcium orthophosphates as a dental regenerative material 429 Keçik, D., Çehreli, S.B., Şar, Ç., Ünver, B., 2008. Effect of acidulated phosphate fluoride and casein phosphopeptide-amorphous calcium phosphate application on shear bond strength of orthodontic brackets. Angle Orthod. 78, 129133. Kensche, A., Pötschke, S., Hannig, C., Richter, G., Hoth-Hannig, W., Hannig, M., 2016. Influence of calcium phosphate and apatite containing products on enamel erosion. Sci. World. J. 2016, 7959273 (12 pp.). Kensche, A., Holder, C., Basche, S., Tahan, N., Hannig, C., Hannig, M., 2017. Efficacy of a mouthrinse based on hydroxyapatite to reduce initial bacterial colonisation in situ. Arch. Oral. Biol. 80, 1826. Kent, J.N., Quinn, J.H., Zide, M.F., 1983. Alveolar ridge augmentation using nonresorbable hydroxylapatite with or without autogenous cancellous bone. J. Oral Maxillofac. Surg. 41, 629642. Kerr, N.W., 1990. The prevalence and pattern of distribution of root caries in a Scottish medieval population. J. Dent. Res. 69, 857860. Khashaba, R.M., Chutkan, N.B., Borke, J.L., 2009. Comparative study of biocompatibility of newly developed calcium phosphate-based root canal sealers on fibroblasts derived from primary human gingiva and a mouse L929 cell line. Int. Endod. J. 42, 711718. Khashaba, R.M., Moussa, M.M., Chutkan, N.B., Borke, J.L., 2011. The response of subcutaneous connective tissue to newly developed calcium phosphate-based root canal sealers. Int. Endod. J. 44, 342352. Khoury, F., Antoun, H., Missika, P., 2007. Bone Augmentation in Oral Implantology. Quintessence Publishing, Hanover Park, IL, 435 pp. Kiba, W., Imazato, S., Takahashi, Y., Yoshioka, S., Ebisu, S., Nakano, T., 2010. Efficacy of polyphasic calcium phosphates as a direct pulp capping material. J. Dent. 38, 828837. Kim, J.S., Baek, S.H., Bae, K.S., 2004. In vitro study of the biocompatibility of newly developed calcium phosphate-based root canal sealers. J. Endod. 30, 708711. Kim, B.I., Jeong, S.H., Jang, S.O., Kim, K.N., Kwon, H.K., Park, Y.D., 2006. Tooth whitening effect of toothpastes containing nano-hydroxyapatite. Key Eng. Mater. 309311, 541544. Kim, M.Y., Kwon, H.K., Choi, C.H., Kim, B.I., 2007. Combined effects of nanohydroxyapatite and NaF on remineralization of early caries lesion. Key Eng. Mater. 330332, 13471350. Kim, S.H., Park, J.B., Lee, C.W., Koo, K.T., Kim, T.I., Seol, Y.J., et al., 2009. The clinical effects of a hydroxyapatite containing toothpaste for dentine hypersensitivity. J. Korean Acad. Periodontol. 39, 8794. Kim, S.G., Hahn, B.D., Park, D.S., Lee, Y.C., Choi, E.J., Chae, W.S., et al., 2011a. Aerosol deposition of hydroxyapatite and 4-hexylresorcinol coatings on titanium alloys for dental implants. J. Oral Maxillofac. Surg. 69, e354e363. Kim, S., Jung, U.W., Lee, Y.K., Choi, S.H., 2011b. Effects of biphasic calcium phosphate bone substitute on circumferential bone defects around dental implants in dogs. Int. J. Oral Maxillofac. Implants 26, 265273. Kitamura, C., Nishihara, T., Terashita, M., Tabata, Y., Washio, A., 2012. Local regeneration of dentin-pulp complex using controlled release of FGF-2 and naturally derived spongelike scaffolds. Int. J. Dent. 2012, 190561 (8 pp.). Klijn, R.J., Hoekstra, J.W.M., van den Beucken, J.J.J.P., Meijer, G.J., Jansen, J.A., 2012. Maxillary sinus augmentation with microstructured tricalcium phosphate ceramic in sheep. Clin. Oral Implant. Res. 23, 274280. Koch, K.A., Brave, D.G., Nasseh, A.A., 2010. Bioceramic technology: closing the endorestorative circle, Part I. Dent. Today 29, 100105.
430 Advanced Dental Biomaterials Kodaka, T., Kobori, M., Hirayama, A., Masayuki, A., 1999. Abrasion of human enamel by brushing with a commercial dentifrice containing hydroxyapatite crystals in vitro. J. Electron Microsc. 48, 167172. Koenigs, J.F., Heller, A.L., Brilliant, J.D., Melfi, R.C., Driskell, T.D., 1975. Induced apical closure of permanent teeth in adult primates using a resorbable form of tricalcium phosphate ceramic. J. Endod. 1, 102106. Kouassi, M., Michaı̈lesco, P., Lacoste-Armynot, A., Boudeville, P., 2003. Antibacterial effect of a hydraulic calcium phosphate cement for dental applications. J. Endod. 29, 100103. Koulourides, T., Cueto, H., Pigman, W., 1961. Rehardening of softened enamel surfaces of human teeth by solutions of calcium phosphates. Nature 189, 226227. Kovtun, A., Kozlova, D., Ganesan, K., Biewald, C., Seipold, N., Gaengler, P., et al., 2012. Chlorhexidine-loaded calcium phosphate nanoparticles for dental maintenance treatment: combination of mineralising and antibacterial effects. RSC Adv. 2, 870875. Krell, K.V., Madison, S., 1985. Comparison of apical leakage in teeth obturated with a calcium phosphate cement or Grossman’s cement using lateral condensation. J. Endod. 11, 336339. Krell, K.V., Wefel, J.S., 1984. A calcium phosphate cement root canal sealer scanning electron microscopic analysis. J. Endod. 10, 571576. Kuemmerle, J.M., Oberle, A., Oechslin, C., Bohner, M., Frei, C., Boecken, I., et al., 2005. Assessment of the suitability of a new brushite calcium phosphate cement for cranioplasty an experimental study in sheep. J. Cran. Maxillofac. Surg. 33, 3744. Kumar, V.L.N., Itthagarun, A., King, N.M., 2008. The effect of casein phosphopeptideamorphous calcium phosphate on remineralization of artificial caries-like lesions: an in vitro study. Aust. Dent. J. 53, 3440. Lanfranco, L.P., Eggers, S., 2012. Caries through time: an anthropological overview. In: Li, M.Y. (Ed.), Contemporary Approach to Dental Caries. InTech, Rijeka, Croatia, pp. 334. Langhorst, S.E., O’Donnell, J.N.R., Skrtic, D., 2009. In vitro remineralization of enamel by polymeric amorphous calcium phosphate composite: quantitative microradiographic study. Dent. Mater. 25, 884891. Larsen, H.D., Finger, I.M., Guerra, L.R., Kent, J.N., 1983. Prosthodontic management of the hydroxylapatite denture patient: a preliminary report. J. Prost. Dent. 49, 461470. Larsen, H.D., Guerra, L.R., Finger, I.M., 1984. Hydroxylapatite: prosthodontic clinical considerations. Compend. Contin. Educ. Dent. 5, 786790. Lavenus, S., Louarn, G., Layrolle, P., 2010. Nanotechnology and dental implants. Int. J. Biomater. 2010, 915327 (9 pp.). Lee, Y.K., Lim, B.S., Kim, C.W., 2003. Mechanical properties of calcium phosphate based dental filling and regeneration materials. J. Oral Rehabil. 30, 418425. Lee, H.J., Min, J.H., Choi, C.H., Kwon, H.G., Kim, B.I., 2007. Remineralization potential of sports drink containing nano-sized hydroxyapatite. Key Eng. Mater. 330-332, 275278. Lee, D.W., Kim, J.Y., Lew, D.H., 2010a. Use of rapidly hardening hydroxyapatite cement for facial contouring surgery. J. Craniofac. Surg. 21, 10841088. Lee, S.K., Lee, S.K., Lee, S.I., Park, J.H., Jang, J.H., Kim, H.W., et al., 2010b. Effect of calcium phosphate cements on growth and odontoblastic differentiation in human dental pulp cells. J. Endod. 36, 15371542. Lee, J.S., Wikesjö, U.M.E., Jung, U.W., Choi, S.H., Pippig, S., Siedler, M., et al., 2010c. Periodontal wound healing/regeneration following implantation of recombinant human
Calcium orthophosphates as a dental regenerative material 431 growth/differentiation factor-5 in a β-tricalcium phosphate carrier into one-wall intrabony defects in dogs. J. Clin. Periodontol. 37, 382389. LeGeros, R.Z., 1988. Calcium phosphate materials in restorative dentistry: a review. Adv. Dent. Res. 2, 164180. LeGeros, R.Z., 1991. Calcium Phosphates in Oral Biology and Medicine. Karger, Basel, Switzerland, 210 pp. LeGeros, R.Z., 1999. Calcium phosphates in demineralization/remineralization processes. J. Clin. Dent. 10, 6573. Levin, M.P., Getter, L., Cutright, D.E., Bhaskar, S.N., 1974. Biodegradable ceramic in periodontal defects. Oral Surg. Oral Med. Oral Pathol. 38, 344351. Lew, D., Farrell, B., Bardach, J., Keller, J., 1997. Repair of craniofacial defects with hydroxyapatite cement. J. Oral Maxillofac. Surg. 55, 14411449. discussion 14491451. Li, T., Akao, M., Takagi, M., 1998. Tissue reaction of hydroxyapatite sol to rat molar pulp. J. Mater. Sci. Mater. Med. 9, 631642. Li, B.G., Wang, J.P., Zhao, Z.Y., Sui, Y.F., Zhang, Y.X., 2007. Mineralizing of nanohydroxyapatite powders on artificial caries. Rare Metal. Mat. Eng. 36, 128130. Li, L., Pan, H.H., Tao, J.H., Xu, X.R., Mao, C.Y., Gu, X.H., et al., 2008. Repair of enamel by using hydroxyapatite nanoparticles as the building blocks. J. Mater. Chem. 18, 40794084. Li, X., Wang, J., Joiner, A., Chang, J., 2014a. The remineralisation of enamel: a review of the literature. J. Dent 42 (Suppl. 1), S12S20. Li, J., Xie, X., Wang, Y., Yin, W., Antoun, J.S., Farella, M., et al., 2014b. Long-term remineralizing effect of casein phosphopeptide-amorphous calcium phosphate (CPP-ACP) on early caries lesions in vivo: a systematic review. J. Dent. 42, 769777. Liang, X., Tang, S.Q., Lu, D., Zhao, Z.H., Chao, Y.L., Wang, H., 1998. Study on hydroxyapatite-coated titanium implants used as orthodontic anchorage an experimental investigation of implant stability and peri-implant neck tissue in dogs. Chin. J. Dent. Res. 1, 5761. Liang, K., Zhou, H., Weir, M.D., Bao, C., Reynolds, M.A., Zhou, X., et al., 2017a. Poly (amido amine) and calcium phosphate nanocomposite remineralization of dentin in acidic solution without calcium phosphate ions. Dent. Mater. 33, 818829. Liang, K., Weir, M.D., Reynolds, M.A., Zhou, X., Li, J., Xu, H.H.K., 2017b. Poly (amido amine) and nano-calcium phosphate bonding agent to remineralize tooth dentin in cyclic artificial saliva/lactic acid. Mater. Sci. Eng. C 72, 717. Liang, K., Xiao, S., Wu, J., Li, J., Weir, M.D., Cheng, L., et al., 2018. Long-term dentin remineralization by poly(amido amine) and rechargeable calcium phosphate nanocomposite after fluid challenges. Dent. Mater. 34, 607618. Liao, F., Chen, Y., Li, Z., Wang, Y., Shi, B., Gong, Z., et al., 2010. A novel bioactive threedimensional β-tricalcium phosphate/chitosan scaffold for periodontal tissue engineering. J. Mater. Sci. Mater. Med. 21, 489496. Lim, K.T., Suh, J.D., Kim, J., Choung, P.H., Chung, J.H., 2011. Calcium phosphate bioceramics fabricated from extracted human teeth for tooth tissue engineering. J. Biomed. Mater. Res. B: Appl. Biomater. 99B, 399411. Lin, J., Zhu, J., Gu, X., Wen, W., Li, Q., Fischer-Brandies, H., et al., 2011. Effects of incorporation of nano-fluorapatite or nano-fluorohydroxyapatite on a resin-modified glass ionomer cement. Acta Biomater. 7, 13461353. Lippert, F., Parker, D.M., Jandt, K.D., 2004a. In situ remineralisation of surface softened human enamel studied with AFM nanoindentation. Surf. Sci. 553, 105114.
432 Advanced Dental Biomaterials Lippert, F., Parker, D.M., Jandt, K.D., 2004b. In vitro demineralization / remineralization cycles at human tooth enamel surfaces investigated by AFM and nanoindentation. J. Colloid Interface Sci. 280, 442448. Lippert, F., Parker, D.M., Jandt, K.D., 2004c. Toothbrush abrasion of surface softened enamel studied with tapping mode AFM and AFM nanoindentation. Caries Res. 38, 464472. Liu, H.C., E, L.L., Wang, D.S., Su, F., Wu, X., Shi, Z.P., et al., 2011. Reconstruction of alveolar bone defects using bone morphogenetic protein 2 mediated rabbit dental pulp stem cells seeded on nano-hydroxyapatite/collagen/poly(L-lactide). Tissue Eng. A 17, 24172433. Liu, Y., Zhang, L., Niu, L.N., Yu, T., Xu, H.H.K., Weir, M.D., et al., 2018a. Antibacterial and remineralizing orthodontic adhesive containing quaternary ammonium resin monomer and amorphous calcium phosphate nanoparticles. J. Dent. 72, 5363. Liu, Y., Zhang, L., Niu, L.N., Yu, T., Xu, H.H.K., Weir, M.D., et al., 2018b. Antibacterial and remineralizing orthodontic adhesive containing quaternary ammonium resin monomer and amorphous calcium phosphate nanoparticles. J. Dent. 72, 5363. Llena, C., Forner, L., Baca, P., 2009. Anticariogenicity of casein phosphopeptide-amorphous calcium phosphate: a review of the literature. J. Contemp. Dent. Pract. 10, 19. Low, S.B., Allen, E.P., Kontogiorgos, E.D., 2015. Reduction in dental hypersensitivity with nano-hydroxyapatite, potassium nitrate, sodium monoflurophosphate and antioxidants. Open Dent. J 9, 9297. Lu, K., Meng, X., Zhang, J., Li, X., Zhou, M., 2007. Inhibitory effect of synthetic nanohydroxyapatite on dental caries. Key Eng. Mater. 336338, 15381541. Luaces-Rey, R., Garcı́a-Rozado, A., Crespo-Escudero, J.L., Seijas, B.P., Arenaz-Búa, J., López-Cedrún, J.L., 2009. Use of carbonated calcium phosphate bone cement and resorbable plates for the treatment of frontal sinus fractures: two case reports. J. Plast. Reconstr. Aesthet. Surg. 62, 272273. Lucas, M.E., Arita, K., Nishino, M., 2003. Toughness, bonding and fluoride release properties of hydroxyapatite-added glass ionomer cement. Biomaterials 24, 37873794. Luo, J., Ning, T., Cao, Y., Zhu, X., Xu, X., Tang, X., et al., 2012. Biomimic enamel remineralization by hybridization calcium- and phosphate-loaded liposomes with amelogenininspired peptide. Key Eng. Mater. 512-515, 17271730. Lussi, A. (Ed.), 2006. Dental Erosion: From Diagnosis to Therapy. Karger, Basel, Switzerland, 219 pp. Lv, K., Zhang, J., Meng, X., Li, X., 2007. Remineralization effect of the nano-HA toothpaste on artificial caries. Key Eng. Mater. 330332, 267270. Lv, K.L., Yuan, H.W., Meng, X.C., Li, X.Y., 2010. Remineralized evaluation of nanohydroxyapatite to artificial caries. Adv. Mater. Res. 105106, 576579. MacDonald, A., Moore, B.K., Newton, C.W., Brown Jr., C.E., 1994. Evaluation of an apatite cement as a root end filling material. J. Endod. 20, 598604. Machot, E.A., Hoffmann, T., Lorenz, K., Khalili, I., Noack, B., 2014. Clinical outcomes after treatment of periodontal intrabony defects with nanocrystalline hydroxyapatite (Ostim) or enamel matrix derivatives (Emdogain): a randomized controlled clinical trial. BioMed Res. Int. 2014, 786353 (9 pp.). Madhumathi, K., Kumar, T.S.S., 2014. Regenerative potential and anti-bacterial activity of tetracycline loaded apatitic nanocarriers for the treatment of periodontitis. Biomed. Mater. (Bristol) 9, Article number 035002. Madhumathi, K., Rekha, L.J., Kumar, T.S.S., 2018. Tailoring antibiotic release for the treatment of periodontal infrabony defects using bioactive gelatin-alginate/apatite nanocomposite films. J. Drug Deliv. Sci. Technol. 43, 5764. Maeda, Y., Okada, M., Okuno, Y., Soga, K., Yamamoto, H., Okazaki, M., 1984. Clinical application of implant stabilizers: combined use of single-crystal sapphire endodontic implants with hydroxyapatite particles. J. Osaka Univ. Dent. Sch. 24, 131144.
Calcium orthophosphates as a dental regenerative material 433 Maksimovskii, I.M., Zemskova, M.I., 1994. Primenenie kal’tsii-fosfatnoi keramiki pri lechenii glubokogo kariesa. [The use of a calcium phosphate ceramic in treating deep caries.]. Stomatologiia 73, 1417. Makvandi, P., Corcione, C.E., Paladini, F., Gallo, A.L., Montagna, F., Jamaledin, R., et al., 2018. Antimicrobial modified hydroxyapatite composite dental bite by stereolithography. Polym. Adv. Technol. 29, 364371. Mangano, C., Venini, E., Venini, G., 1984. Il DAC blu come sostituto dell’osso in chirurgia orale. [Dense apatite ceramic (DAC) as a bone substitute in oral surgery]. Dent. Cadmos 52, 97. 100101, 104105. Mangin, C., Yesilsoy, C., Nissan, R., Stevens, R., 2003. The comparative sealing ability of hydroxyapatite cement, mineral trioxide aggregate, and super ethoxybenzoic acid as root-end filling materials. J. Endod. 29, 261264. Markiewicz, M., Dzierzak-Mietla, M., Frankiewicz, A., Zielinska, P., Koclega, A., Kruszelnicka, M., et al., 2012. Treating oral mucositis with a supersaturated calcium phosphate rinse: comparison with control in patients undergoing allogeneic hematopoietic stem cell transplantation. Support. Care Cancer 20, 22232229. Markopoulou, C.E., Dereka, X.E., Vavouraki, H.N., Pepelassi, E.E., Mamalis, A.A., Karoussis, I.K., et al., 2011. Effect of rhTGF-β1 combined with bone grafts on human periodontal cell differentiation. Growth Factors 29, 1420. Marković, D., Živojinović, V., Koković, V., Jokanović, V., 2004. Hydroxyapatite as a root canal system filling material: cytotoxicity testing. Mater. Sci. Forum 453454, 555560. Marković, D., Petrović, V., Krstić, N., Lazarević-Macanović, M., Nikolić, Z., 2007. Radiological assessment of apex formation following use of hydroxyapatite. Acta Veter. (Beograd) 57, 275287. Marukawa, K., Ueki, K., Okabe, K., Nakagawa, K., Yamamoto, E., 2011. Use of self-setting α-tricalcium phosphate for maxillary sinus augmentation in rabbit. Clin. Oral Implant. Res. 22, 606612. Masudi, S.M., Luddin, N., Mohamad, D., Alkashakhshir, J.J., Adnan, R., Ramli, R.A., 2010. In vitro study on apical sealing ability of nano-hydroxyapatite-filled epoxy resin based endodontic sealer. AIP Conf. Proceed. 1217, 467471. Matsuse, K., Hashimoto, Y., Kakinoki, S., Yamaoka, T., Morita, S., 2018. Periodontal regeneration induced by porous alpha-tricalcium phosphate with immobilized basic fibroblast growth factor in a canine model of 2-wall periodontal defects. Med. Mol. Morphol. 51, 4856. Matsuura, T., Akizuki, T., Hoshi, S., Ikawa, T., Kinoshita, A., Sunaga, M., et al., 2015. Effect of a tunnel-structured β-tricalcium phosphate graft material on periodontal regeneration: a pilot study in a canine one-wall intrabony defect model. J. Periodont. Res. 50, 347355. Mayne, R.J., Cochrane, N.J., Cai, F., Woods, M.G., Reynolds, E.C., 2011. In-vitro study of the effect of casein phosphopeptide amorphous calcium fluoride phosphate on iatrogenic damage to enamel during orthodontic adhesive removal. Am. J. Orthod. Dentofac. Orthoped. 139, e543e551. McClendon, J.F., Carpousis, A., 1945. Prevention of dental caries by brushing the teeth with powdered fluorapatite. J. Dent. Res. 24, 199. Medvecky, L., Stulajterova, R., Giretova, M., Mincik, J., Vojtko, M., Balko, J., et al., 2018. Effect of tetracalcium phosphate/monetite toothpaste on dentin remineralization and tubule occlusion in vitro. Dent. Mater. 34, 442451. Meffert, R.M., Thomas, J.R., Hamilton, K.M., Brownstein, C.N., 1985. Hydroxylapatite as an alloplastic graft in the treatment of human periodontal osseous defects. J. Periodontol. 56, 6373.
434 Advanced Dental Biomaterials Meguro, D., Hayakawa, T., Kawasaki, M., Kasai, K., 2006a. Shear bond strength of calcium phosphate ceramic brackets to human enamel. Angle Orthod. 76, 301305. Meguro, D., Hayakawa, T., Kasai, K., 2006b. Efficacy of using orthodontic adhesive resin in bonding and debonding characteristics of a calcium phosphate ceramic bracket. Orthod. Waves 65, 148154. Mehdawi, I., Neel, E.A., Valappil, S.P., Palmer, G., Salih, V., Pratten, J., et al., 2009. Development of remineralizing, antibacterial dental materials. Acta Biomater. 5, 25252539. Mehta, D., Gowda, V., Finger, W.J., Sasaki, K., 2015. Randomized, placebo-controlled study of the efficacy of a calcium phosphate containing paste on dentin hypersensitivity. Dent. Mater. 31, 12981303. Mei, H.L., Chen, L.Y., Zhang, D., Zhang, P.L., Liu, B., Zhao, W., et al., 2009. Effects of casein phosphopeptide-stabilized amorphous calcium phosphate solution on enamel remineralization. J. Clin. Rehabil. Tissue Eng. Res. 13, 48254828. Mellgren, T., Qin, T., Öhman-Mägi, C., Zhang, Y., Wu, B., Xia, W., et al., 2018. Calcium phosphate microspheres as a delivery vehicle for tooth-bleaching agents. J. Dent. Res. 97, 283288. Melo, M.A., Cheng, L., Weir, M.D., Hsia, R.C., Rodrigues, L.K., Xu, H.H.K., 2013a. Novel dental adhesive containing antibacterial agents and calcium phosphate nanoparticles. J. Biomed. Mater. Res. B: Appl. Biomater. 101B, 620629. Melo, M.A., Cheng, L., Zhang, K., Weir, M.D., Rodrigues, L.K., Xu, H.H.K., 2013b. Novel dental adhesives containing nanoparticles of silver and amorphous calcium phosphate. Dent. Mater. 9, 199210. Mendes, A.C., Restrepo, M., Bussaneli, D., Zuanon, A.C., 2018. Use of casein amorphous calcium phosphate (CPP-ACP) on white-spot lesions: randomised clinical trial. Oral Health Prev. Dent. 16, 2731. Meng, X., Lv, K., Zhang, J., Qu, D., 2007. Caries inhibitory activity of the nano-HA in vitro. Key Eng. Mater. 330332, 251254. Mercier, P., 1995. Failures in ridge reconstruction with hydroxyapatite. Analysis of cases and methods for surgical revision. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 80, 389393. Mercier, P., Bellavance, F., Cholewa, J., Djokovic, S., 1996. Long-term stability of atrophic ridges reconstructed with hydroxylapatite: a prospective study. J. Oral Maxillofac. Surg. 54, 960969. Michaı̈lesco, P., Kouassi, M., Briak, H.E., Armynot, A., Boudeville, P., 2005. Antimicrobial activity and tightness of a DCPD CaO-based hydraulic calcium phosphate cement for root canal filling. J. Biomed. Mater. Res. B: Appl. Biomater. 74B, 760-767. Mielczarek, A., Michalik, J., 2014. The effect of nano-hydroxyapatite toothpaste on enamel surface remineralization. An in vitro study. Am. J. Dent. 27, 287290. Miloslavich, E.L., 1925. Calcium metabolism in its relation to dental pathology calciprival odontopathia. Int. J. Orthodontia Oral Surg. Radiography 11, 111123. Min, J.H., Kwon, H.K., Kim, B.I., 2011. The addition of nano-sized hydroxyapatite to a sports drink to inhibit dental erosion in vitro study using bovine enamel. J. Dent. 39, 629635. Mishra, S., Singh, R.K., Mohammad, S., Pradhan, R., Pal, U.S., 2010. A comparative evaluation of decalcified freeze dried bone allograft, hydroxyapatite and their combination in osseous defects of the jaws. J. Maxillofac. Oral Surg. 9, 236240. Miura, K., Matsui, K., Kawai, T., Kato, Y., Matsui, A., Suzuki, O., et al., 2012. Octacalcium phosphate collagen composites with titanium mesh facilitate alveolar augmentation in canine mandibular bone defects. Int. J. Oral Maxillofac. Surg. 41, 11611169.
Calcium orthophosphates as a dental regenerative material 435 Miyamoto, C.T., Wobb, J., Micaily, B., Li, S., Achary, M.P., 2012. A retrospective match controlled study of supersaturated calcium phosphate oral rinse vs. supportive care for radiation induced oral mucositis. J. Cancer Ther. 3, 630636. Mjör, I.A., Toffenetti, F., 2000. Secondary caries: a literature review with case reports. Quintessence Int. 31, 165179. Moenning, J.E., Wolford, L.M., 1989. Coralline porous hydroxyapatite as a bone graft substitute in orthognathic surgery: 24-month follow-up results. Int. J. Adult Orthod. Orthognathic Surg. 4, 105117. Monroe, E.A., Votava, W., Bass, D.B., Mcmullen, J., 1971. New calcium phosphate ceramic material for bone and tooth implants. J. Dent. Res. 50, 860862. Moogi, P., Rupesh, M.C., Mehta, D., Finger, W.J., Sasaki, K., 2017. A novel desensitizer paste containing calcium phosphate: randomized, placebo-controlled, double-blinded and 6 months trial. World J. Dent. 8, 440444. Mora, F., Ouhayoun, J.P., 1995. Clinical evaluation of natural coral and porous hydroxyapatite implants in periodontal bone lesions: results of a 1-year follow-up. J. Clin. Periodontol. 22, 877884. Moreau, J.L., Sun, L., Chow, L.C., Xu, H.H.K., 2011. Mechanical and acid neutralizing properties and bacteria inhibition of amorphous calcium phosphate dental nanocomposite. J. Biomed. Mater. Res. B: Appl. Biomater. 98B, 8088. Morgan, M.V., Adams, G.G., Bailey, D.L., Tsao, C.E., Fischman, S.L., Reynolds, E.C., 2008. The anticariogenic effect of sugar-free gum containing CPP-ACP nanocomplexes on approximal caries determined using digital bitewing radiography. Caries Res. 42, 171184. Morris, H.F., Ochi, S., Spray, J.R., Olson, J.W., 2000. Periodontal-type measurements associated with hydroxyapatite-coated and non-HA-coated implants: uncovering to 36 months. Ann. Periodontol. 5, 5667. Moshaverinia, A., Ansari, S., Moshaverinia, M., Roohpour, N., Darr, J.A., Rehman, I., 2008a. Effects of incorporation of hydroxyapatite and fluoroapatite nanobioceramics into conventional glass ionomer cements (GIC). Acta Biomater. 4, 432440. Moshaverinia, A., Ansari, S., Movasaghi, Z., Billington, R.W., Darr, J.A., Rehman, I.U., 2008b. Modification of conventional glass-ionomer cements with N-vinylpyrrolidone containing polyacids, nano-hydroxy and fluoroapatite to improve mechanical properties. Dent. Mater. 24, 13811390. Müller, N., 1989. Alveolarfortsatzerhohung mit Hydroxylapatit. Klinische Erfahrungen aus prothetischer Sicht. [Augmentation of alveolar process with hydroxylapatite. Clinical orthodontic experience]. Dtsch. Zahnarztl. Z 44, 596599. Najibfard, K., Chedjieu, I., Ramalingam, K., Amaechi, B.T., 2011. Remineralization of early caries by a nano-hydroxyapatite dentifrice. J. Clin. Dent. 22, 139143. Nakagawa, K.I., 1983. Clinico-pathological studies of dental biomaterials in endodontic therapy, with special reference to the biocompatibility of the hydroxyapatite on exposed vital human pulp. Shikwa Gakuho 83, 501527. Nakao, K., Morita, R., Saji, Y., Ishida, K., Tomita, Y., Ogawa, M., et al., 2007. The development of a bioengineered organ germ method. Nat. Methods 4, 227230. Nakata, T., Kitasako, Y., Sadr, A., Nakashima, S., Tagami, J., 2018. Effect of a calcium phosphate and fluoride paste on prevention of enamel demineralization. Dent. Mater. J. 37, 6570. Nelson, D.R., von Gonten, A.S., 1988. Prosthodontic management of the hydroxylapatiteaugmented ridge. Gen. Dent. 36, 315319. Nery, E.B., Lynch, K.L., Hirthe, W.M., Mueller, K.H., 1975. Bioceramic implants in surgically produced infrabony defects. J. Periodontol. 46, 328347.
436 Advanced Dental Biomaterials Nery, E.B., Lynch, K.L., Rooney, G.E., 1978. Alveolar ridge augmentation with tricalcium phosphate ceramic. J. Prost. Dent. 40, 668675. Nery, E.B., Eslami, A., van Swol, R.L., 1990a. Biphasic calcium phosphate ceramic combined with fibrillar collagen with and without citric acid conditioning in the treatment of periodontal osseous defects. J. Periodontol. 61, 166172. Nery, E.B., Lee, K.K., Czajkowski, S., Dooner, J.J., Duggan, M., Ellinger, R.F., et al., 1990b. A Veterans Administration Cooperative Study of biphasic calcium phosphate ceramic in periodontal osseous defects. J. Periodontol. 61, 737744. Nery, E.B., LeGeros, R.Z., Lynch, K.L., Lee, K., 1992. Tissue response to biphasic calcium phosphate ceramic with different ratios of HA/βTCP in periodontal osseous defects. J. Periodontol. 63, 729735. Neto, D.M.A., Carvalho, E.V., Rodrigues, E.A., Feitosa, V.P., Sauro, S., Mele, G., et al., 2016. Novel hydroxyapatite nanorods improve anti-caries efficacy of enamel infiltrants. Dent. Mater. 32, 784793. Nevins, A.J., Cymerman, J.J., 2015. Revitalization of open apex teeth with apical periodontitis using a collagen-hydroxyapatite scaffold. J. Endod. 41, 966973. Nevins, A.J., LaPorta, R.F., Borden, B.G., Spangberg, L.S., 1980. Pulpotomy and partial pulpectomy procedures in monkey teeth using cross-linked collagen-calcium phosphate gel. Oral Surg. Oral Med. Oral Pathol. 49, 360365. Ngoc, N.T.B., Mukohyama, H., Hlaing, S., Kondo, H., Inoue, T., Taniguchi, H., et al., 1997. Prosthodontic treatment for patients with large mandibular defects; porous hydroxyapatite grafts. J. Med. Dent. Sci. 44, 9398. Ngoc, T.N.D., Tra, T.N., Nguyen, T.H., Huynh, C.K., Van, T.V., 2018. Preparation and characterization of nano-sized biphasic calcium phosphate (BCP) for demineralized dentin infiltration in hypersensitivity treatment. IFMBE Proc. 63, 677680. Nicholson, J.W., Hawkins, S.J., Smith, J.E., 1993. The incorporation of hydroxyapatite into glass-polyalkenoate (“glass-ionomer”) cements: a preliminary study. J. Mater. Sci. Mater. Med. 4, 418421. Ning, T.Y., Xu, X.H., Zhu, L.F., Zhu, X.P., Chu, C.H., Liu, L.K., et al., 2012. Biomimetic mineralization of dentin induced by agarose gel loaded with calcium phosphate. J. Biomed. Mater. Res. B: Appl. Biomater. 100, 138144. Niu, L.N., Zhang, W., Pashley, D.H., Breschi, L., Mao, J., Chen, J.H., et al., 2014. Biomimetic remineralization of dentin. Dent. Mater. 30, 7796. Niwa, M., Sato, T., Li, W., Aoki, H., Aoki, H., Daisaku, T., 2001. Polishing and whitening properties of toothpaste containing hydroxyapatite. J. Mater. Sci. Mater. Med. 12, 277281. Niwa, K., Ogawa, K., Miyazawa, K., Aoki, T., Kawai, T., Goto, S., 2009. Application of α-tricalcium phosphate coatings on titanium subperiosteal orthodontic implants reduces the time for absolute anchorage: a study using rabbit femora. Dent. Mater. J. 28, 477486. Noetzel, J., Özer, K., Reisshauer, B.H., Anil, A., Rössler, R., Neumann, K., et al., 2006. Tissue responses to an experimental calcium phosphate cement and mineral trioxide aggregate as materials for furcation perforation repair, a histological study in dogs. Clin. Oral Investig. 10, 7783. Nyman, S., Lindhe, J., Karring, T., Rylander, H., 1982. New attachment following surgical treatment of human periodontal disease. J. Clin. Periodontol. 9, 290296. O’Donnell, J.N.R., Schumacher, G.E., Antonucci, J.M., Skrtic, D., 2009. Adhesion of amorphous calcium phosphate composites bonded to dentin: a study in failure modality. J. Biomed. Mater. Res. B: Appl. Biomater. 90B, 238249.
Calcium orthophosphates as a dental regenerative material 437 Oduncu, B.S., Yucel, S., Aydin, I., Sener, I.D., Yamaner, G., 2010. Polymerisation shrinkage of light-cured hydroxyapatite (HA) reinforced dental composites. World Acad. Sci. Eng. Technol. 40, 286291. Ogawa, K., Miyaji, H., Kato, A., Kosen, Y., Momose, T., Yoshida, T., et al., 2016. Periodontal tissue engineering by nano beta-tricalcium phosphate scaffold and fibroblast growth factor-2 in one-wall infrabony defects of dogs. J. Periodont. Res. 51, 758767. Ogiso, M., Tabata, T., Kuo, P.T., Borgese, D., 1994. A histologic comparison of the functional loading capacity of an occluded dense apatite implant and the natural dentition. J. Prost. Dent. 71, 581588. Ohara, T., Itaya, T., Usami, K., Ando, Y., Sakurai, H., Honda, M.J., et al., 2010. Evaluation of scaffold materials for tooth tissue engineering. J. Biomed. Mater. Res. A 94A, 800805. Okashi, T., Kani, T., Isozaki, A., Nishida, A., Shintani, H., Tokumoto, T., et al., 1991. Remineralization of artificial caries lesions by hydroxyapatite. J. Dent. Health 41, 214223. Okubo, N., Fujita, T., Ishii, Y., Ota, M., Shibukawa, Y., Yamada, S., 2013. Coverage of gingival recession defects using acellular dermal matrix allograft with or without betatricalcium phosphate. J. Biomater. Appl. 27, 627637. Okuda, K., Tai, H., Tanabe, K., Suzuki, H., Sato, T., Kawase, T., et al., 2005. Platelet-rich plasma combined with a porous hydroxyapatite graft for the treatment of intrabony periodontal defects in humans: a comparative controlled clinical study. J. Periodontol. 76, 890898. Okuda, K., Yamamiya, K., Kawase, T., Mizuno, H., Ueda, M., Yoshie, H., 2009. Treatment of human infrabony periodontal defects by grafting human cultured periosteum sheets combined with platelet-rich plasma and porous hydroxyapatite granules: case series. J. Int. Acad. Periodontol. 11, 206213. Ong, J.L., Chan, D.C.N., 1999. Hydroxyapatite and their use as coatings in dental implants: a review. Crit. Rev. Biomed. Eng. 28, 667707. Onuma, K., Yamagishi, K., Oyane, A., 2005. Nucleation and growth of hydroxyapatite nanocrystals for nondestructive repair of early caries lesions. J. Cryst. Growth 282, 199207. Oshiro, M., Yamaguchi, K., Takamizawa, T., Inage, H., Watanabe, T., Irokawa, A., et al., 2007. Effect of CPP-ACP paste on tooth mineralization: an FE-SEM study. J. Oral Sci. 49, 115120. Owadally, I.D., Ford, T.R.P., 1994. Effect of addition of hydroxyapatite on the physical properties of IRM. Int. Endod. J. 27, 227232. Owadally, I.D., Chong, B.S., Ford, T.R.P., Wilson, R.F., 1994. Biological properties of IRM with the addition of hydroxyapatite as a retrograde root filling material. Endod. Dent. Traumatol. 10, 228232. Palarie, V., Bicer, C., Lehmann, K.M., Zahalka, M., Draenert, F.G., Kämmerer, P.W., 2012. Early outcome of an implant system with a resorbable adhesive calcium-phosphate coating—a prospective clinical study in partially dentate patients. Clin. Oral Investig. 16, 10391048. Panahi, F., Rabiee, S.M., Shidpour, R., 2017. Synergic effect of chitosan and dicalcium phosphate on tricalcium silicate-based nanocomposite for root-end dental application. Mater. Sci. Eng. C 80, 631-641. Pandit, N., Gupta, R., Gupta, S., 2010. A comparative evaluation of biphasic calcium phosphate material and bioglass in the treatment of periodontal osseous defects: a clinical and radiological study. J. Contemp. Dent. Pract. 11, 2532.
438 Advanced Dental Biomaterials Panich, M., Poolthong, S., 2009. The effect of casein phosphopeptide-amorphous calcium phosphate and a cola soft drink on in vitro enamel hardness. J. Am. Dent. Assoc. 140, 455460. Papas, A., 2008. Calcium phosphate mouth rinse for preventing oral mucositis. Community Oncol. 5, 171172. Park, S.Y., Cha, J.Y., Kim, K.N., Hwang, C.J., 2013. The effect of casein phosphopeptide amorphous calcium phosphate on the in vitro shear bond strength of orthodontic brackets. Korean J. Orthod. 43, 2328. Pendleton, L.W., 1873. The lacto-phosphate of lime. Trans. Maine Med. Assoc. 4, 313318. Pepelassi, E.M., Bissada, N.F., Greenwell, H., Farah, C.F., 1991. Doxycycline-tricalcium phosphate composite graft facilitates osseous healing in advanced periodontal furcation defects. J. Periodontol. 62, 106115. Pickel, F.D., Bilotti, A., 1965. The effects of a chewing gum containing dicalcium phosphate on salivary calcium and phosphate. Alabama J. Med. Sci. 2, 286287. Pietruska, M., Pietruski, J., Nagy, K., Brecx, M., Arweiler, N.B., Sculean, A., 2012a. Fouryear results following treatment of intrabony periodontal defects with an enamel matrix derivative alone or combined with a biphasic calcium phosphate. Clin. Oral Investig. 16, 11911197. Pietruska, M., Skurska, A., Pietruski, J., Dolińska, E., Arweiler, N., Milewski, R., et al., 2012b. Clinical and radiographic evaluation of intrabony periodontal defect treatment by open flap debridement alone or in combination with nanocrystalline hydroxyapatite bone substitute. Ann. Anat. 194, 533537. Porciani, P.F., Chazine, M., Grandini, S., 2014. A clinical study of the efficacy of a new chewing gum containing calcium hydroxyapatite in reducing dentin hypersensitivity. J. Clin. Dent. 25, 3236. Portella, F.F., Collares, F.M., dos Santos, L.A., dos Santos, B.P., Camassola, M., Leitune, V.C.B., et al., 2015. Glycerol salicylate-based containing α-tricalcium phosphate as a bioactive root canal sealer. J. Biomed. Mater. Res. B: Appl. Biomater. 103B, 16631669. Prabhakar, A.R., Sharma, D., Sugandhan, S., 2012. Comparative evaluation of the remineralising effects and surface microhardness of glass ionomer cement containing grape seed extract and casein phosphopeptide amorphous calcium phosphate: an in vitro study. Eur. Arch. Paediatr. Dent. 13, 138143. Pradeep, A.R., Bajaj, P., Rao, N.S., Agarwal, E., Naik, S.B., 2017. Platelet-rich fibrin combined with a porous hydroxyapatite graft for the treatment of 3-wall intrabony defects in chronic periodontitis: a randomized controlled clinical trial. J. Periodontol. 88, 12881296. Quinn, B., 2013. Efficacy of a supersaturated calcium phosphate oral rinse for the prevention and treatment of oral mucositis in patients receiving high-dose cancer therapy: a review of current data. Eur. J. Cancer Care 22, 564579. Ramalingam, L., Messer, L.B., Reynolds, E.C., 2005. Adding casein phosphopeptideamorphous calcium phosphate to sports drinks to eliminate in vitro erosion. Pediatr. Dent. 27, 6167. Ranjitkar, S., Rodriguez, J.M., Kaidonis, J.A., Richards, L.C., Townsend, G.C., Bartlett, D. W., 2009a. The effect of casein phosphopeptide-amorphous calcium phosphate on erosive enamel and dentine wear by toothbrush abrasion. J. Dent. 37, 250254. Ranjitkar, S., Narayana, T., Kaidonis, J.A., Hughes, T.E., Richards, L.C., Townsend, G.C., 2009b. The effect of casein phosphopeptide-amorphous calcium phosphate on erosive dentine wear. Aust. Dent. J. 54, 101107.
Calcium orthophosphates as a dental regenerative material 439 Reddi, S.P., Stevens, M.R., Kline, S.N., Villanueva, P., 1999. Hydroxyapatite cement in craniofacial trauma surgery, indications and early experience. J. Cran. Maxillofac. Trauma 5, 712. Redondo, L.M., Garcı́a Cantera, J.M., Hernández, A.V., Puerta, C.V., 1995. Effect of particulate porous hydroxyapatite on osteoinduction of demineralized bone autografts in experimental reconstruction of the rat mandible. Int. J. Oral Maxillofac. Surg. 24, 445448. Reema, S.D., Lahiri, P.K., Roy, S.S., 2014. Review of casein phosphopeptides-amorphous calcium phosphate. Chin. J. Dent. Res. 17, 714. Reis, E.C.C., Borges, A.P.B., Araújo, M.V.F., Mendes, V.C., Guan, L., Davies, J.E., 2011. Periodontal regeneration using a bilayered PLGA/calcium phosphate construct. Biomaterials 32, 92449253. Reynolds, E.C., 1997. Remineralization of enamel subsurface lesions by casein phosphopeptide-stabilized calcium phosphate solutions. J. Dent. Res. 76, 15871595. Reynolds, E.C., 1998. Anticariogenic complexes of amorphous calcium phosphate stabilized by casein phosphopeptides: a review. Spec. Care Dent. 18, 816. Reynolds, E.C., 2008. Calcium phosphate-based remineralization systems: scientific evidence? Aust. Dent. J. 53, 268273. Reynolds, E.C., Cai, F., Cochrane, N.J., Shen, P., Walker, G.D., Morgan, M.V., et al., 2008. Fluoride and casein phosphopeptide-amorphous calcium phosphate. J. Dent. Res. 87, 344348. Rezvani, M., Atai, M., Rouhollahi, H., Malekhoseini, K., Rezai, H., Hamze, F., 2015. Effect of nano-tricalcium phosphate and nanohydroxyapatite on staining susceptibility of bleached enamel. Int. Sch. Res. Notices 2015, 935264 (7 pp.). Richardson, A.S., Hole, L.W., McCombie, F., Kolthammer, J., 1972. Anticariogenic effect of dicalcium phosphate dihydrate chewing gum: results after two years. J. Can. Dent. Assoc. 38, 213218. Ridgway, H.K., Mellonig, J.T., Cochran, D.L., 2008. Human histologic and clinical evaluation of recombinant human platelet-derived growth factor and beta-tricalcium phosphate for the treatment of periodontal intraosseous defects. Int. J. Periodontics Restorative Dent. 28, 171179. Rimondini, L., Palazzo, B., Iafisco, M., Canegallo, L., Demarosi, F., Merlo, M., et al., 2007. The remineralizing effect of carbonate-hydroxyapatite nanocrystals on dentine. Mater. Sci. Forum 539543, 602605. Roane, J.B., Benenati, F.W., 1987. Successful management of a perforated mandibular molar using amalgam and hydroxyapatite. J. Endod. 13, 400404. Rosen, H.M., Ackerman, J.L., 1991. Porous block hydroxyapatite in orthognathic surgery. Angle Orthod. 61, 185191. discussion 192. Roveri, N., Battistella, E., Foltran, I., Foresti, E., Iafisco, M., Lelli, M., et al., 2008. Synthetic biomimetic carbonate-hydroxyapatite nanocrystals for enamel remineralization. Adv. Mater. Res. 4750, 821824. Roveri, N., Foresti, E., Lelli, M., Lesci, I.G., 2009a. Recent advancements in preventing teeth health hazard: the daily use of hydroxyapatite instead of fluoride. Rec. Pat. Biomed. Eng. 2, 197215. Roveri, N., Battistella, E., Bianchi, C.L., Foltran, I., Foresti, E., Iafisco, M., et al., 2009b. Surface enamel remineralization: biomimetic apatite nanocrystals and fluoride ions different effects. J. Nanomater. 2009, 746383 (9 pp.). Saffar, J.L., Colombier, M.L., Detienville, R., 1990. Bone formation in tricalcium phosphatefilled periodontal intrabony lesions. Histological observations in humans. J. Periodontol. 61, 209216.
440 Advanced Dental Biomaterials Saito, E., Saito, A., Kuboki, Y., Kimura, M., Honma, Y., Takahashi, T., et al., 2012. Periodontal repair following implantation of beta-tricalcium phosphate with different pore structures in class III furcation defects in dogs. Dent. Mater. J. 31, 681688. Saito, A., Saito, E., Ueda, Y., Shibukawa, Y., Honma, Y., Takahashi, T., et al., 2014. Effect of tunnel structure of β-TCP on periodontal repair in class III furcation defects in dogs. Bioceram. Dev. Appl 4, 073. Salyer, K.E., Hall, C.D., 1989. Porous hydroxyapatite as an onlay bone-graft substitute for maxillofacial surgery. Plast. Reconstr. Surg. 84, 236244. Sanada, Y., Fujinaka, T., Yoshimine, T., Kato, A., 2011. Optimal reconstruction of the bony defect after frontotemporal craniotomy with hydroxyapatite cement. J. Clin. Neurosci. 18, 280282. Sato, I., Akizuki, T., Oda, S., Tsuchioka, H., Hayashi, C., Takasaki, A.A., et al., 2009. Histological evaluation of alveolar ridge augmentation using injectable calcium phosphate bone cement in dogs. J. Oral Rehabil 36, 762769. Scarano, A., Degidi, M., Perrotti, V., Piattelli, A., Iezzi, G., 2012. Sinus augmentation with phycogene hydroxyapatite: histological and histomorphometrical results after 6 months in humans. A case series. Oral Maxillofac. Surg. 16, 4145. Schemehorn, B.R., Orban, J.C., Wood, G.D., Fischer, G.M., Winston, A.E., 1999. Remineralization by fluoride enhanced with calcium and phosphate ingredients. J. Clin. Dent. 10, 1316. Schlafer, S., Birkedal, H., Olsen, J., Skovgaard, J., Sutherland, D.S., Wejse, P.L., et al., 2016. Calcium-phosphate-osteopontin particles for caries control. Biofouling 32, 349357. Schliephake, H., Scharnweber, D., Roesseler, S., Dard, M., Sewing, A., Aref, A., 2006. Biomimetic calcium phosphate composite coating of dental implants. Int. J. Oral Maxillofac. Implants 21, 738746. Schneider, B., Diedrich, P., 1989. Interaktion von kieferorthopadischer Zahnbewegung und Hydroxylapatit-Keramik. [Interaction between orthodontic tooth movement and hydroxyapatite ceramics]. Dtsch. Zahnarztl. Z 44, 282285. Schreiber, C.T., Shern, R.J., Chow, L.C., Kingman, A., 1988. Effects of rinses with an acidic calcium phosphate solution on fluoride uptake, caries, and in situ plaque pH in rats. J. Dent. Res. 67, 959963. Sculean, A., Windisch, P., Szendröi-Kiss, D., Horváth, A., Rosta, P., Becker, J., et al., 2008. Clinical and histologic evaluation of an enamel matrix derivative combined with a biphasic calcium phosphate for the treatment of human intrabony periodontal defects. J. Periodontol. 79, 19911999. Seifi, M., Arayesh, A., Shamloo, N., Hamedi, R., 2015. Effect of nanocrystalline hydroxyapatite socket preservation on orthodontically induced inflammatory root resorption. Cell J. 16, 514527. Sena, M., Yamashita, Y., Nakano, Y., Ohgaki, M., Nakamura, S., Yamashita, K., et al., 2004. Octacalcium phosphate-based cement as a pulp-capping agent in rats. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 97, 749-755. Seong, K.C., Cho, K.S., Daculsi, C., Seris, E., Daculsi, G., 2014. Eight-year clinical followup of sinus grafts with micro-macroporous biphasic calcium phosphate granules. Key Eng. Mater. 587, 321324. Sepantafar, M., Mohammadi, H., Maheronnaghsh, R., Tayebi, L., Baharvand, H., 2018. Single phased silicate-containing calcium phosphate bioceramics: promising biomaterials for periodontal repair. Ceram. Int. (early view).
Calcium orthophosphates as a dental regenerative material 441 Shammukha, G., Santhosh, B.P., Preeti, J., Naveen, B., Inderjith, G., Santhosh, B., et al., 2012. Evaluation of changes in salivary concentration of calcium by CPP-ACP containing chewing gum a clinical trial. Int. J. Adv. Res. Oral Sci. 1, 17. Sharma, H., Gupta, C., Thakur, S., Srivastava, S., 2017. Comparative evaluation of calcium phosphate-based varnish and resin-modified glass ionomer-based varnish in reducing dentinal hypersensitivity: a randomized controlled clinical trial. Eur. J. Dent. 11, 491495. Shayegan, A., Petein, M., Abbeele, A.V., 2008. Beta-tricalcium phosphate, white mineral trioxide aggregate, white Portland cement, ferric sulfate, and formocresol used as pulpotomy agents in primary pig teeth. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 105, 536542. Shayegan, A., Atash, R., Petein, M., Abbeele, A.V., 2010. Nanohydroxyapatite used as a pulpotomy and direct pulp capping agent in primary pig teeth. J. Dent. Child. (Chic.) 77, 7783. Shayesteh, Y.S., Khojasteh, A., Soleimani, M., Alikhasi, M., Khoshzaban, A., Ahmadbeigi, N., 2008. Sinus augmentation using human mesenchymal stem cells loaded into a β-tricalcium phosphate/hydroxyapatite scaffold. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 106, 203209. Shen, P., Cai, F., Nowicki, A., Vincent, J., Reynolds, E.C., 2001. Remineralization of enamel subsurface lesions by sugar-free chewing gum containing casein phosphopeptideamorphous calcium phosphate. J. Dent. Res. 80, 20662070. Shen, Q., Sun, J., Wu, J., Liu, C., Chen, F., 2010. An in vitro investigation of the mechanical-chemical and biological properties of calcium phosphate/calcium silicate/ bismutite cement for dental pulp capping. J. Biomed. Mater. Res. B: Appl. Biomater. 94B, 141148. Shern, R.J., Couet, K.M., Chow, L.C., Brown, W.E., 1979. Effects of sequential calcium phosphate-fluoride rinses on fluoride uptake in rats. J. Dent. Res. 58 (Spec. Iss. B), 1023. Shern, R.J., Chow, L.C., Couet, K.M., Kingman, A., Brown, W.E., 1984. Effects of sequential calcium phosphate-fluoride rinses on dental plaque, staining, fluoride uptake, and caries in rats. J. Dent. Res. 63, 13551359. Shetty, S., Kohad, R., Yeltiwar, R., 2010. Hydroxyapatite as an in-office agent for tooth hypersensitivity: a clinical and scanning electron microscopic study. J. Periodontol. 81, 17811789. Shetty, R., Bhat, A.N., Mehta, D., Finger, W.J., 2017. Effect of a calcium phosphate desensitizer on pre- and postcementation sensitivity of teeth prepared for full-coverage restorations: a randomized, placebo-controlled clinical study. Int. J. Prosthodont. 30, 3842. Shi, H., Ma, J., Zhao, N., Chen, Y., Liao, Y., 2008. Periodontal regeneration in experimentally-induced alveolar bone dehiscence by an improved porous biphasic calcium phosphate ceramic in beagle dogs. J. Mater. Sci. Mater. Med. 19, 35153524. Shieh, T.M., Hsu, S.M., Chang, K.C., Chen, W.C., Lin, D.J., 2017. Calcium phosphate cement with antimicrobial properties and radiopacity as an endodontic material. Materials 10, 1256. Shirai, Y., Okuda, K., Kubota, T., Wolff, L.F., Yoshie, H., 2012. The comparative effectiveness of granules or blocks of superporous hydroxyapatite for the treatment of intrabony periodontal defects. Open J. Stomatol. 2, 8187. Shirakata, Y., Oda, S., Kinoshita, A., Kikuchi, S., Tsuchioka, H., Ishikawa, I., 2002. Histocompatible healing of periodontal defects after application of injectable calcium phosphate bone cement. A preliminary study in dogs. J. Periodontol. 73, 10431053.
442 Advanced Dental Biomaterials Shirakata, Y., Setoguchi, T., Machigashira, M., Matsuyama, T., Furuichi, Y., Hasegawa, K., et al., 2008. Comparison of injectable calcium phosphate bone cement grafting and open flap debridement in periodontal intrabony defects: a randomized clinical trial. J. Periodontol. 79, 2532. Shirakata, Y., Taniyama, K., Yoshimoto, T., Takeuchi, N., Noguchi, K., 2012. Effect of bone swaging with calcium phosphate bone cement on periodontal regeneration in dogs. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. 114, 3542. Shirakawa, M., Nomura, T., Itoh, T., Sakai, N., Shizume, M., 1988. Clinical application of hydroxyapatite ceramics APS-7 in the field of oral surgery. Shigaku 76, 782815. Shon, W.J., Bae, K.S., Baek, S.H., Kum, K.Y., Han, A.R., Lee, W.C., 2012. Effects of calcium phosphate endodontic sealers on the behavior of human periodontal ligament fibroblasts and MG63 osteoblast-like cells. J. Biomed. Mater. Res. B: Appl. Biomater. 100B, 21412147. Shyam, R., Manjunath, B.C., Kumar, A., Narang, R., Ghanghas, M., Rani, G., 2017. Role of casein phosphopeptide-amorphous calcium phosphate (CPP-ACP) in prevention of dental caries: a review. Int. J. Adv. Res. 5, 9991004. Silva, M.F.D.A., Melo, E.V.D.S., Stewart, B., de Vizio, W., Sintes, J.L., Petrone, M.E., et al., 2001. The enhanced anticaries efficacy of a sodium fluoride and dicalcium phosphate dihydrate dentifrice in a dual-chambered tube. A 2-year caries clinical study on children in Brazil. Am. J. Dent. 14 (Spec. Iss. 5), 19A23A. Silva, K.G., Pedrini, D., Delbem, A.C.B., Ferreira, L., Cannon, M., 2010. In situ evaluation of the remineralizing capacity of pit and fissure sealants containing amorphous calcium phosphate and/or fluoride. Acta Odontol. Scand. 68, 1118. Silverstone, L.M., 1972. Remineralization of human enamel in vitro. Proc. R. Soc. Med. 65, 906908. Silverstone, L.M., Johnson, N.W., 1971. The effect on sound human enamel of exposure to calcifying fluids in vitro. Caries Res. 5, 323342. Sinai, I.H., Romea, D.J., Glassman, G., Morse, D.R., Fantasia, J., Furst, M.L., 1989. An evaluation of tricalcium phosphate as a treatment for endodontic perforations. J. Endod. 15, 399403. Singh, M.L., Papas, A.S., 2009. Long-term clinical observation of dental caries in salivary hypofunction patients using a supersaturated calcium-phosphate remmeralizing rinse. J. Clin. Dent. 20, 8792. Singh, K.A., Burstein, F.D., Williams, J.K., 2010. Use of hydroxyapatite cement in pediatric craniofacial reconstructive surgery: strategies for avoiding complications. J. Craniofac. Surg. 21, 11301135. Sintes, J.L., Elı́as-Boneta, A., Stewart, B., Volpe, A.R., Lovett, J., 2002. Anticaries efficacy of a sodium monofluorophosphate dentifrice containing xylitol in a dicalcium phosphate dihydrate base. A 30-month caries clinical study in Costa Rica. Am. J. Dent. 15, 215219. Skrtic, D., Antonucci, J.M., 2005. Matrix resin effects on selected physicochemical properties of amorphous calcium phosphate composites. J. Bioact. Compat. Polym. 20, 2949. Skrtic, D., Antonucci, J.M., 2007. Dental composites based on amorphous calcium phosphate resin composition/physicochemical properties study. J. Biomater. Appl. 21, 375393. Skrtic, D., Antonucci, J.M., 2011. Bioactive polymeric composites for tooth mineral regeneration: physicochemical and cellular aspects. J. Funct. Biomater. 2, 271307. Skrtic, D., Antonucci, J.M., 2016. Polymeric dental composites based on remineralizing amorphous calcium phosphate fillers. Curr. Trends Polym. Sci. 17, 131.
Calcium orthophosphates as a dental regenerative material 443 Skrtic, D., Antonucci, J.M., Eanes, E.D., 1996a. Improved properties of amorphous calcium phosphate fillers in remineralizing resin composites. Dent. Mater. 12, 295301. Skrtic, D., Hailer, A.W., Takagi, S., Antonucci, J.M., Eanes, E.D., 1996b. Quantitative assessment of the efficacy of amorphous calcium phosphate/methacrylate composites in remineralizing caries-like lesions artificially produced in bovine enamel. J. Dent. Res. 75, 16791686. Skrtic, D., Antonucci, J.M., Eanes, E.D., 2001. Effect of the monomer and filler system on the remineralizing potential of bioactive dental composites based on amorphous calcium phosphate. Polym. Adv. Technol. 12, 369379. Skrtic, D., Antonucci, J.M., Eanes, E.D., 2003. Amorphous calcium phosphate-based bioactive polymeric composites for mineralized tissue regeneration. J. Res. Natl. Inst. Stand. Technol. 108, 167182. Skrtic, D., Antonucci, J.M., Eanes, E.D., Eidelman, N., 2004. Dental composites based on hybrid and surface-modified amorphous calcium phosphates. Biomaterials 25, 11411150. Smartt, J.M., Karmacharya, J., Gannon, F.H., Ong, G., Jackson, O., Bartlett, S.P., et al., 2005. Repair of the immature and mature craniofacial skeleton with a carbonated calcium phosphate cement: assessment of biocompatibility, osteoconductivity and remodeling capacity. Plast. Reconstr. Surg. 115, 16421650. Sodagar, A., Akhavan, A., Hashemi, E., Arab, S., Pourhajibagher, M., Sodagar, K., et al., 2016. Evaluation of the antibacterial activity of a conventional orthodontic composite containing silver/hydroxyapatite nanoparticles. Prog. Orthod. 17, 40. Sorensen, R.G., Wikesjö, U.M.E., Kinoshita, A., Wozney, J.M., 2004. Periodontal repair in dogs: Evaluation of a bioresorbable calcium phosphate cement (Ceredext) as a carrier for rhBMP-2. J. Clin. Periodontol. 31, 796804. Souza, B.M., Comar, L.P., Vertuan, M., Neto, C.F., Buzalaf, M.A., Magalhães, A.C., 2015. Effect of an experimental paste with hydroxyapatite nanoparticles and fluoride on dental demineralisation and remineralisation in situ. Caries Res. 49, 499507. Stahl, S.S., Froum, S.J., 1987. Histologic and clinical responses to porous hydroxylapatite implants in human periodontal defects. Three to twelve months postimplantation. J. Periodontol. 58, 689695. Stavropoulos, A., Windisch, P., Szendröi-Kiss, D., Peter, R., Gera, I., Sculean, A., 2010. Clinical and histologic evaluation of granular beta-tricalcium phosphate for the treatment of human intrabony periodontal defects: a report on five cases. J. Periodontol. 81, 325334. Stefanic, M., Krnel, K., Pribosic, I., Kosmac, T., 2012. Rapid biomimetic deposition of octacalcium phosphate coatings on zirconia ceramics (Y-TZP) for dental implant applications. Appl. Surf. Sci. 258, 46494656. Strub, J.R., Gaberthüel, T.W., 1978. Trikalziumphosphat und dessen biologisch abbaubare Keramik in der parodontalen Knochenchirurgie Eine Literaturubersicht. [Tricalcium phosphate and its biodegradable ceramics in periodontal bone surgery. A review of the literature]. SSO Schweiz Monatsschr. Zahnheilkd. 88, 798803. Struillou, X., Boutigny, H., Badran, Z., Fellah, B.H., Gauthier, O., Sourice, S., et al., 2011. Treatment of periodontal defects in dogs using an injectable composite hydrogel/ biphasic calcium phosphate. J. Mater. Sci. Mater. Med. 22, 17071717. Su, B., Su, J., Ran, J., Su, B., 2008. Biological performance of dental biphasic calcium phosphate ceramics modified by cold plasma. Key Eng. Mater. 368372, 12641267. Sugawara, A., Chow, L.C., Takagi, S., Nishiyama, M., Ohashi, M., 1989. An in vitro study of dentin hypersensitivity using calcium phosphate cement. Shika Zairyo Kikai 8, 282294.
444 Advanced Dental Biomaterials Sugawara, A., Chow, L.C., Takagi, S., Chohayeb, H., 1990. In vitro evaluation of the sealing ability of a calcium phosphate cement when used as a root canal sealer-filler. J. Endod. 16, 162165. Sugawara, A., Fujikawa, K., Kusama, K., Nishiyama, M., Murai, S., Takagi, S., et al., 2002. Histopathologic reaction of calcium phosphate cement for alveolar ridge augmentation. J. Biomed. Mater. Res. 61, 4752. Sugawara, A., Fujikawa, K., Takagi, S., Chow, L.C., 2008. Histological analysis of calcium phosphate bone grafts for surgically created periodontal bone defects in dogs. Dent. Mater. J. 27, 787-794. Suge, T., Ishikawa, K., Kawasaki, A., Suzuki, K., Matsuo, T., Noiri, Y., et al., 2002. Calcium phosphate precipitation method for the treatment of dentin hypersensitivity. Am. J. Dent. 15, 220226. Sullivan, R.J., Charig, A., Blake-Haskins, J., Zhang, Y.P., Miller, S.M., Strannick, M., et al., 1997. In vivo detection of calcium from dicalcium phosphate dihydrate dentifrices in demineralized human enamel and plaque. Adv. Dent. Res. 11, 380387. Sullivan, R.J., Masters, J., Cantore, R., Roberson, A., Petrou, I., Stranick, M., et al., 2001. Development of an enhanced anticaries efficacy dual component dentifrice containing sodium fluoride and dicalcium phosphate dihydrate. Am. J. Dent. 14 (Spec. Iss. 5), 3A11A. Sultan, S., Chaitra, T.R., Chaudhary, S., Manuja, N., Kaur, H., Amit, S.A., et al., 2016. Effect of ACP-CPP chewing gum and natural chewable products on plaque pH: calcium and phosphate concentration. J. Clin. Diagn. Res. 10, ZC13ZC17. Sun, W., Zhang, F., Guo, J., Wu, J., Wu, W., 2008. Effects of amorphous calcium phosphate on periodontal ligament cell adhesion and proliferation in vitro. J. Med. Biol. Eng. 28, 107112. Sun, Y., Li, X., Deng, Y., Sun, J.N., Tao, D., Chen, H., et al., 2014. Mode of action studies on the formation of enamel minerals from a novel toothpaste containing calcium silicate and sodium phosphate salts. J. Dent. 42 (Suppl. 1), S30S38. Sung, Y.M., Shin, Y.K., Ryu, J.J., 2007. Preparation of hydroxyapatite/zirconia bioceramic nanocomposites for orthopaedic and dental prosthesis applications. Nanotechnology 18, 065602. Svanberg, A., Ohrn, K., Birgegard, G., 2015. Caphosols mouthwash gives no additional protection against oral mucositis compared to cryotherapy alone in stem cell transplantation. A pilot study. Eur. J. Oncol. Nurs. 19, 5053. Sverzut, A.T., Rodrigues, D.C., Lauria, A., Armando, R.S., de Oliveira, P.T., Moreira, R.W.F., 2015. Clinical, radiographic, and histological analyses of calcium phosphate cement as filling material in maxillary sinus lift surgery. Clin. Oral Implant. Res. 26, 633638. Tabrizi, A., Cakirer, B., 2011. A comparative evaluation of casein phosphopeptideamorphous calcium phosphate and fluoride on the shear bond strength of orthodontic brackets. Eur. J. Orthod. 33, 282287. Tamimi, F., Torres, J., Cabarcos, E.L., Bassett, D.C., Habibovic, P., Luceron, E., et al., 2009. Minimally invasive maxillofacial vertical bone augmentation using brushite based cements. Biomaterials 30, 208216. Tanaka, O., Hirai, T., Murase, H., 1988. Prosthodontic analysis in mandibular ridge augmentation with hydroxyapatite particle. Part I. Evaluation of alveolar ridge form. Nippon Hotetsu Shika Gakkai Zasshi 32, 13451357. Tanaka, O., Hirai, T., Murase, H., 1989. Prosthodontic analysis in mandibular ridge augmentation with hydroxyapatite particle. 2. Evaluation of masticatory function and overall assessment. Nippon Hotetsu Shika Gakkai Zasshi 33, 14661476.
Calcium orthophosphates as a dental regenerative material 445 ten Cate, J.M., Arends, J., 1977. Remineralization of artificial enamel lesions in vitro. Caries Res 11, 277286. ten Cate, J.M., Arends, J., 1978. Remineralization of artificial enamel lesions in vitro. II. Determination of activation energy and reaction order. Caries Res. 12, 213222. ten Cate, J.M., Arends, J., 1980. Remineralization of artificial enamel lesions in vitro. III. A study of the deposition mechanism. Caries Res. 14, 351358. ten Cate, J.M., Jongebloed, W.L., Arends, J., 1981. Remineralization of artificial enamel lesions in vitro. IV. Influence of fluorides and diphosphonates on short- and long-term remineralization. Caries Res. 15, 6069. Teodorović, N., Martinović, Ž., 2005. Significance of crown-down root canal preparation technique in endodontic therapy by using the hydroxylapatite sealer. Vojnosanit. Pregl. 62, 447452. Thaweboon, S., Nakornchai, S., Miyake, Y., Yanagisawa, T., Thaweboon, B., Sooampon, S., et al., 2009. Remineralization of enamel subsurface lesions by xylitol chewing gum containing funoran and calcium hydrogenphosphate. Southeast Asian J. Trop. Med. Public Health 40, 345353. Thein-Han, W., Liu, J., Xu, H.H.K., 2012. Calcium phosphate cement with biofunctional agents and stem cell seeding for dental and craniofacial bone repair. Dent. Mater. 28, 10591070. Thompson, J.Y., Stoner, B.R., Piascik, J.R., 2007. Ceramics for restorative dentistry: critical aspects for fracture and fatigue resistance. Mater. Sci. Eng. C 27, 565569. Tian, K., Peng, M., Wu, P., Liao, C.H., Huang, F.Y., 2012. Biomineralization of the hydroxyapatite with 3D-structure for enamel reconstruction. Adv. Mater. Res. 391392, 633637. Tinsley, D., Watson, C.J., Russell, J.L., 2001. A comparison of hydroxylapatite coated implant retained fixed and removable mandibular prostheses over 4 to 6 years. Clin. Oral Implant. Res. 12, 159166. Tiwari, S., Nandlal, B., 2012. Role of Synthetic Hydroxyapatite in Dentistry. Lap Lambert Academic Publishing, Saarbrucken, Deutschland, 90 pp. Tomizuka, K., 1990. Experimental study on apatite ceramics used as endodontic endosseous implant. Kokubyo Gakkai Zasshi 57, 201226. Tonomura, A., Mizuno, D., Hisada, A., Kuno, N., Ando, Y., Sumita, Y., et al., 2010. Differential effect of scaffold shape on dentin regeneration. Ann. Biomed. Eng. 38, 16641671. Trombelli, L., Simonelli, A., Pramstraller, M., Wikesjö, U.M.E., Farina, R., 2010. Single flap approach with and without guided tissue regeneration and a hydroxyapatite biomaterial in the management of intraosseous periodontal defects. J. Periodontol. 81, 12561263. Tschoppe, P., Zandim, D.L., Martus, P., Kielbassa, A.M., 2011. Enamel and dentine remineralization by nano-hydroxyapatite toothpastes. J. Dent. 39, 430437. Tung, M.S., Eichmiller, F.C., 1999. Dental applications of amorphous calcium phosphates. J. Clin. Dent. 10, 16. Uysal, T., Amasyali, M., Koyuturk, A.E., Sagdic, D., 2009a. Efficiency of amorphous calcium phosphate-containing orthodontic composite and resin modified glass ionomer on demineralization evaluated by a new laser fluorescence device. Eur. J. Dent. 3, 127134. Uysal, T., Ulker, M., Baysal, A., Usumez, S., 2009b. Microleakage between composite-wire and composite-enamel interfaces of flexible spiral wire retainers. Part 2: Comparison of
446 Advanced Dental Biomaterials amorphous calcium phosphate-containing adhesive with conventional lingual retainer composite. Eur. J. Orthod. 31, 652657. Uysal, T., Ulker, M., Akdogan, G., Ramoglu, S.I., Yilmaz, E., 2009c. Bond strength of amorphous calcium phosphate-containing orthodontic composite used as a lingual retainer adhesive. Angle Orthod. 79, 117121. Uysal, T., Ustdal, A., Nur, M., Catalbas, B., 2010a. Bond strength of ceramic brackets bonded to enamel with amorphous calcium phosphate-containing orthodontic composite. Eur. J. Orthod. 32, 281284. Uysal, T., Amasyali, M., Ozcan, S., Koyuturk, A.E., Akyol, M., Sagdic, D., 2010b. In vivo effects of amorphous calcium phosphate-containing orthodontic composite on enamel demineralization around orthodontic brackets. Aust. Dent. J. 55, 285291. Uysal, T., Amasyali, M., Koyuturk, A.E., Ozcan, S., Sagdic, D., 2010c. Amorphous calcium phosphate-containing orthodontic composites. Do they prevent demineralisation around orthodontic brackets? Aust. Orthod. J 26, 1015. Uysal, T., Baysal, A., Uysal, B., Aydinbelge, M., Al-Qunaian, T., 2011. Do fluoride and casein phosphopeptide-amorphous calcium phosphate affect shear bond strength of orthodontic brackets bonded to a demineralized enamel surface? Angle Orthod. 81, 490495. Vaca-Cornejo, F., Reyes, H.M., Jiménez, S.H.D., Velázquez, R.A.L., Jiménez, J.M.D., 2017. Pilot study using a chitosan-hydroxyapatite implant for guided alveolar bone growth in patients with chronic periodontitis. J. Funct. Biomater. 8, 29. Vaishnavi, C., Mohan, B., Narayanan, L.L., 2011. Treatment of endodontically induced periapical lesions using hydroxyapatite, platelet-rich plasma, and a combination of both: an in vivo study. J. Conserv. Dent. 14, 140146. Vanichvatana, S., Auychai, P., 2013. Efficacy of two calcium phosphate pastes on the remineralization of artificial caries: a randomized controlled double-blind in situ study. Int. J. Oral Sci. 5, 224228. Vano, M., Derchi, G., Barone, A., Pinna, R., Usai, P., Covani, U., 2018. Reducing dentine hypersensitivity with nano-hydroxyapatite toothpaste: a double-blind randomized controlled trial. Clin. Oral. Investig. 22, 313320. van Oirschot, B.A.J.A., Bronkhorst, E.M., van den Beucken, J.J.J.P., Meijer, G.J., Jansen, J. A., Junker, R., 2016a. A systematic review on the long-term success of calcium phosphate plasma-spray-coated dental implants. Odontology 104, 347356. van Oirschot, B.A.J.A., Bronkhorst, E.M., van den Beucken, J.J.J.P., Meijer, G.J., Jansen, J. A., Junker, R., 2016b. A systematic review on the long-term success of calcium phosphate plasma-spray-coated dental implants. Odontology 104, 347356. Vargas, S., Estevez, M., Hernandez, A., Laiz, J.C., Brostow, W., Lobland, H.E.H., et al., 2013. Hydroxyapatite based hybrid dental materials with controlled porosity and improved tribological and mechanical properties. Mater. Res. Innov. 17, 154160. Viale-Bouroncle, S., Bey, B., Reichert, T.E., Schmalz, G., Morsczeck, C., 2011. β-Tricalcium-phosphate stimulates the differentiation of dental follicle cells. J. Mater. Sci. Mater. Med. 22, 17191724. Vogel, G.L., Zhang, Z., Carey, C.M., Ly, A., Chow, L.C., Proskin, H.M., 1998. Composition of plaque and saliva following a sucrose challenge and use of an α-tricalcium-phosphate-containing chewing gum. J. Dent. Res. 77, 518524. erratum: p. 1575. Vogel, G.L., Zhang, Z., Carey, C.M., Ly, A., Chow, L.C., Proskin, H.M., 2000. Composition of plaque and saliva following use of an α-tricalcium-phosphate-containing chewing gum and a subsequent sucrose challenge. J. Dent. Res. 79, 5862.
Calcium orthophosphates as a dental regenerative material 447 Vyavhare, S., Sharma, D.S., Kulkarni, V.K., 2015. Effect of three different pastes on remineralization of initial enamel lesion: an in vitro study. J. Clin. Pediatr. Dent. 39, 149160. Wagner, W., Wiltfang, J., Pistner, H., Yildirim, M., Ploder, B., Chapman, M., et al., 2012. Bone formation with a biphasic calcium phosphate combined with fibrin sealant in maxillary sinus floor elevation for delayed dental implant. Clin. Oral Implants Res. 23, 11121117. Walker, G.D., Cai, F., Shen, P., Reynolds, C., Ward, B., Fone, C., et al., 2006. Increased remineralization of tooth enamel by milk containing added casein phosphopeptideamorphous calcium phosphate. J. Dairy Res. 73, 7478. Walker, G.D., Cai, F., Shen, P., Bailey, D.L., Yuan, Y., Cochrane, N.J., et al., 2009. Consumption of milk with added casein phosphopeptide-amorphous calcium phosphate remineralizes enamel subsurface lesions in situ. Aust. Dent. J. 54, 245249. Walker, G.D., Cai, F., Shen, P., Adams, G.G., Reynolds, C., Reynolds, E.C., 2010. Casein phosphopeptide-amorphous calcium phosphate incorporated into sugar confections inhibits the progression of enamel subsurface lesions in situ. Caries Res. 44, 3340. Wang, Y., Wang, Q.S., 2010. Application of nano-hydroxyapatite and its composite biomaterials in stomatology. J. Clin. Rehabil. Tissue Eng. Res. 14, 14261428. Wang, J.P., Geogi, G., 2014. Antimicrobial and sealant properties of nanohydroxyapatite as endodontic sealer. Chin. J. Tissue Eng. Res. 18, 33503354. Wang, W., Zeng, X.L., Liu, W., 2008. Dental caries in ancient Chinese in Xia dynasty. Zhonghua Kou Qiang Yi Xue Za Zhi 43, 308310 (in Chinese). Wang, C.J., Zhang, Y.F., Wei, J., Wei, S.C., 2011a. Repair of artificial enamel lesions by nano fluorapatite paste containing fluorin. J. Clin. Rehabil. Tissue Eng. Res. 15, 63466350. Wang, H.Y., Sun, K.N., Shan, T., Yang, X.Q., Zhao, Y., Liang, Y.J., 2011b. Biomimetic synthesis of fluorapatite coating. Adv. Mater. Res. 306307, 6371. Wang, L., Li, J., Xie, Y., Yang, P., Liao, Y., Guo, G., 2012. Effect of nano biphasic calcium phosphate bioceramics on periodontal regeneration in the treatment of alveolar defects. Adv. Mater. Res. 486, 422425. Wang, Q., Liu, S., Gao, X., Wei, Y., Deng, X., Chen, H., et al., 2015a. Remineralizing efficacy of fluorohydroxyapatite gel on artificial dentinal caries lesion. J. Nanomater. 2015, Article number 380326. Wang, X., Suzawa, T., Miyauchi, T., Zhao, B., Yasuhara, R., Anada, T., et al., 2015b. Synthetic octacalcium phosphate-enhanced reparative dentine formation via induction of odontoblast differentiation. J. Tissue Eng. Regen. Med. 9, 13101320. Wang, L., Xie, X., Li, C., Liu, H., Zhang, K., Zhou, Y., et al., 2017. Novel bioactive root canal sealer to inhibit endodontic multispecies biofilms with remineralizing calcium phosphate ions. J. Dent. 60, 2535. Wefel, J.S., Harless, J.D., 1987. The use of saturated DCPD in remineralization of artificial caries lesions in vitro. J. Dent. Res. 66, 16401643. Wegehaupt, F.J., Attin, T., 2010. The role of fluoride and casein phosphopeptide/amorphous calcium phosphate in the prevention of erosive/abrasive wear in an in vitro model using hydrochloric acid. Caries Res. 44, 358363. Wei, J., Wang, J., Shan, W., Liu, X., Ma, J., Liu, C., et al., 2011a. Development of fluorapatite cement for dental enamel defects repair. J. Mater. Sci. Mater. Med. 22, 1607-1614. Wei, J., Wang, J., Liu, X., Ma, J., Liu, C., Fang, J., et al., 2011b. Preparation of fluoride substituted apatite cements as the building blocks for tooth enamel restoration. Appl. Surf. Sci. 257, 78877892.
448 Advanced Dental Biomaterials Weir, M.D., Chow, L.C., Xu, H.H.K., 2012. Remineralization of demineralized enamel via calcium phosphate nanocomposite. J. Dent. Res. 91, 979984. Weir, M.D., Ruan, J., Zhang, N., Chow, L.C., Zhang, K., Chang, X., et al., 2017. Effect of calcium phosphate nanocomposite on in vitro remineralization of human dentin lesions. Dent. Mater. 33, 10331044. Weiss, P., Layrolle, P., Clergeau, L.P., Enckel, B., Pilet, P., Amouriq, Y., et al., 2007. The safety and efficacy of an injectable bone substitute in dental sockets demonstrated in a human clinical trial. Biomaterials 28, 32953305. Wen, Z., Chen, J., Wang, H., Zhong, S., Hu, Y., Wang, Z., et al., 2016. Abalone watersoluble matrix for self-healing biomineralization of tooth defects. Mater. Sci. Eng. C 67, 182187. West, N.X., Joiner, A., 2014. Enamel mineral loss. J. Dent 42 (Suppl. 1), S2S11. White, J.M., Goodis, H., 1991. In vitro evaluation of an hydroxyapatite root canal system filling material. J. Endod. 17, 561566. Widström, E., Birn, H., Haugejorden, O., Sundberg, H., 1992. Fear of amalgam: dentists’ experiences in the Nordic countries. Int. Dent. J. 42, 6570. Willershausen, B., Schulz-Dobrick, B., Gleissner, C., 2009. In vitro evaluation of enamel remineralisation by a casein phosphopeptide-amorphous calcium phosphate paste. Oral Health Prev. Dent. 7, 1321. Wilson, C.J., 1975. The effect of calcium sucrose phosphates chewing gum on caries incidence in children. J. Wis. Dent. Assoc. 51, 521525. Wiltfang, J., Schlegel, K.A., Schultze-Mosgau, S., Nkenke, E., Zimmermann, R., Kessler, P., 2003. Sinus floor augmentation with β-tricalciumphosphate (β-TCP): does platelet-rich plasma promote its osseous integration and degradation? Clin. Oral Implant. Res. 14, 213218. Wiltfang, J., Kessler, P., Buchfelder, M., Merten, H.A., Neukam, F.W., Rupprecht, S., 2004. Reconstruction of skull bone defects using the hydroxyapatite cement with calvarial split transplants. J. Oral Maxillofac. Surg. 62, 2935. Witjaksono, W., Naing, L., Mulyawati, E., Samsudin, A.R., Oo, M.M.T., 2007. Sealing ability of hydroxyapatite as a root canal sealer: in vitro study. Dent. J 40, 101105. Wittkampf, A.R.M., 1988. Augmentation of the maxillary alveolar ridge with hydroxylapatite and fibrin glue. J. Oral Maxillofac. Surg. 46, 10191021. Wolford, L.M., Freitas, R.Z., 1999. Porous block hydroxyapatite as a bone graft substitute in the correction of jaw and craniofacial deformities. BUMC Proc. 12, 243246. Wolford, L.M., Wardrop, R.W., Hartog, J.M., 1987. Coralline porous hydroxylapatite as a bone graft substitute in orthognathic surgery. J. Oral Maxillofac. Surg. 45, 10341042. Wu, D., Yang, J., Li, J., Chen, L., Tang, B., Chen, X., et al., 2013. Hydroxyapatite-anchored dendrimer for in situ remineralization of human tooth enamel. Biomaterials 34, 50365047. Wu, X.T., Mei, M.L., Li, Q.L., Cao, C.Y., Chen, J.L., Xia, R., et al., 2015. A direct electric field-aided biomimetic mineralization system for inducing the remineralization of dentin collagen matrix. Materials 8, 78897899. Xiao, Y., Yin, Q., Wang, L., Bao, C., 2015. Macro-porous calcium phosphate scaffold with collagen and growth factors for periodontal bone regeneration in dogs. Ceram. Int. 41, 9951003. Xiao, Z., Que, K., Wang, H., An, R., Chen, Z., Qiu, Z., et al., 2017. Rapid biomimetic remineralization of the demineralized enamel surface using nano-particles of amorphous calcium phosphate guided by chimaeric peptides. Dent. Mater. 33, 12171228.
Calcium orthophosphates as a dental regenerative material 449 Xie, C., Lu, H., Li, W., Chen, F.M., Zhao, Y.M., 2012. The use of calcium phosphate-based biomaterials in implant dentistry. J. Mater. Sci. Mater. Med. 23, 853862. Xie, X.J., Xing, D., Wang, L., Zhou, H., Weir, M.D., Bai, Y.X., et al., 2018. Novel rechargeable calcium phosphate nanoparticle-containing orthodontic cement. Int. J. Oral Sci. 9, 2432. Xu, H.H.K., Moreau, J.L., 2010. Dental glass-reinforced composite for caries inhibition: calcium phosphate ion release and mechanical properties. J. Biomed. Mater. Res. B: Appl. Biomater. 92B, 332340. Xu, H.H.K., Takagi, S., Sun, L., Hussain, L., Chow, L.C., Guthrie, W.F., et al., 2006a. Development of a nonrigid, durable calcium phosphate cement for use in periodontal bone repair. J. Am. Dent. Assoc. 137, 11311138. Xu, H.H.K., Sun, L., Weir, M.D., Antonucci, J.M., Yakagi, S., Chow, L.C., et al., 2006b. Nano DCPA-whisker composites with high strength and Ca and PO4 release. J. Dent. Res. 85, 722727. Xu, H.H.K., Sun, L., Weir, M.D., Takagi, S., Chow, L.C., Hockey, B., 2007a. Effects of incorporating nanosized calcium phosphate particles on of properties of whiskerreinforced dental composites. J. Biomed. Mater. Res. B: Appl. Biomater. 81B, 116125. Xu, H.H.K., Weir, M.D., Sun, L., Takagi, S., Chow, L.C., 2007b. Effects of calcium phosphate nanoparticles on Ca-PO4 composite. J. Dent. Res. 86, 378383. Xu, H.H.K., Weir, M.D., Sun, L., 2007c. Nanocomposites with Ca and PO4 release: effects of reinforcement, dicalcium phosphate particle size and silanization. Dent. Mater. 23, 14821491. Xu, H.H.K., Weir, M.D., Sun, L., 2009. Calcium and phosphate ion releasing composite: effect of pH on release and mechanical properties. Dent. Mater. 25, 535542. Xu, H.H.K., Moreau, J.L., Sun, L., Chow, L.C., 2011. Nanocomposite containing amorphous calcium phosphate nanoparticles for caries inhibition. Dent. Mater. 27, 762769. Yamagishi, K., Onuma, K., Suzuki, T., Okada, F., Tagami, J., Otsuki, M., et al., 2005. Materials chemistry: a synthetic enamel for rapid tooth repair. Nature 433, 819. Yamaguchi, K., 1989. Biological evaluation of new hydroxyapatite endodontic cement in vivo. Histopathological and clinico-pathological observation. Dent. Sci. Rep. 89, 761792. Yamamoto, E., Kato, N., Nishikawa, H., Kusunoki, M., Hayami, T., Yoshikawa, K., et al., 2013a. Adhesive strength between flexible hydroxyapatite sheet and tooth enamel. Key Eng. Mater. 529530, 522525. Yamamoto, E., Kato, N., Isai, A., Nishikawa, H., Kusunoki, M., Yoshikawa, K., et al., 2013b. Restoration of tooth enamel using a flexible hydroxyapatite sheet coated with tricalcium phosphate. Bioceram. Dev. Appl S1, 006. Yamamoto, E., Kato, N., Isai, A., Nishikawa, H., Hashimoto, Y., Yoshikawa, K., et al., 2015. A novel treatment for dentine cavities with intraoral laser ablation method using an Er: YAG laser. Key Eng. Mater. 631, 262266. Yang, S.E., Baek, S.H., Lee, W., Kum, K.Y., Bae, K.S., 2007. In vitro evaluation of the sealing ability of newly developed calcium phosphate-based root canal sealer. J. Endod. 33, 978981. Yang, X., Yang, F., Walboomers, X.F., Bian, Z., Fan, M., Jansen, J.A., 2010. The performance of dental pulp stem cells on nanofibrous PCL/gelatin/nHA scaffolds. J. Biomed. Mater. Res. A 93A, 247257. Yap, A.U., Pek, Y.S., Kumar, R.A., Cheang, P., Khor, K.A., 2002. Experimental studies on a new bioactive material: HAIonomer cements. Biomaterials 23, 955962.
450 Advanced Dental Biomaterials Yates, R., Owens, J., Jackson, R., Newcombe, R.G., Addy, M., 1998. A split-mouth placebocontrolled study to determine the effect of amorphous calcium phosphate in the treatment of dentine hypersensitivity. J. Clin. Periodontol. 25, 687692. Yelick, P.C., Vacanti, J.P., 2006. Bioengineered teeth from tooth bud cells. Dent. Clin. North Am. 50, 191203. Yengopal, V., Mickenautsch, S., 2009. Caries preventive effect of casein phosphopeptideamorphous calcium phosphate (CPP-ACP): a meta-analysis. Acta Odontol. Scand. 67, 321332. Yin, Y., Yun, S., Fang, J., Chen, H., 2009. Chemical regeneration of human tooth enamel under near-physiological conditions. Chem. Commun. 58925894. Ylinen, P., Suuronen, R., Taurio, R., Törmälä, P., Rokkanen, P., 2002. Use of hydroxylapatite/polymer-composite in facial bone augmentation. An experimental study. Int. J. Oral Maxillofac. Surg. 31, 405409. Yoshiba, K., Yoshiba, N., Iwaku, M., 1994. Histological observations of hard tissue barrier formation in amputated dental pulp capped with alpha-tricalcium phosphate containing calcium hydroxide. Endod. Dent. Traumatol. 10, 113120. Yoshikawa, M., Hayami, S., Tsuji, I., Toda, T., 1997. Histopathological study of a newly developed root canal sealer containing tetracalcium-dicalcium phosphates and 1.0% chondroitin sulfate. J. Endod. 23, 162166. Yoshinari, M., Oda, Y., Inoue, T., Matsuzaka, K., Shimono, M., 2002. Bone response to calcium phosphate-coated and bisphosphonate-immobilized titanium implants. Biomaterials 23, 28792885. Yoshino, K., Taniguchi, Y., Yoda, Y., Matsukubo, T., 2013. Autotransplantation of tooth by osteotome sinus floor elevation technique with beta-tricalcium phosphate (β-TCP). J. Oral Maxillofac. Surg. Med. Pathol. 25, 351354. Yoshinuma, N., Sato, S., Fukuyama, T., Murai, M., Ito, K., 2012. Ankylosis of nonresorbable hydroxyapatite graft material as a contributing factor in recurrent periodontitis. Int. J. Periodontics Restorative Dent. 32, 331336. Yu, J.M., Choe, H.C., 2018. Mg-containing hydroxyapatite coatings on Ti-6Al-4V alloy for dental materials. Appl. Surf. Sci. 432, 294299. Yu, L., Xu, B., Wu, B., 2003. Treatment of combined endodontic-periodontic lesions by intentional replantation and application of hydroxyapatites. Dent. Traumatol. 19, 6063. Yu, J.H., Wang, Y.J., Deng, Z.H., Tang, L., Li, Y.F., Shi, J.N., et al., 2007. Odontogenic capability: bone marrow stromal stem cells versus dental pulp stem cells. Biol. Cell 99, 465674. Yu, O.Y., Zhao, I.S., Mei, M.L., Lo, E.C.M., Chu, C.H., 2017. A review of the common models used in mechanistic studies on demineralization-remineralization for cariology research. Dent. J 5, 20. Yukna, R.A., Callan, D.P., Krauser, J.T., Evans, G.H., Aichelmann-Reidy, M.E., Moore, K., et al., 1998. Multi-center clinical evaluation of combination anorganic bovine-derived hydroxyapatite matrix (ABM)/cell binding peptide (P-15) as a bone replacement graft material in human periodontal osseous defects. 6-month results. J. Periodontol. 69, 655663. Zaharia, A., Muşat, V., Anghel, E.M., Atkinson, I., Mocioiu, O.C., Buşilă, M., et al., 2017. Biomimetic chitosan-hydroxyapatite hybrid biocoatings for enamel remineralization. Ceram. Int. 43, 1139011402. Zecha, P.J., Schortinghuis, J., van der Wal, J.E., Nagursky, H., van den Broek, K.C., Sauerbier, S., et al., 2011. Applicability of equine hydroxyapatite collagen (eHAC) bone
Calcium orthophosphates as a dental regenerative material 451 blocks for lateral augmentation of the alveolar crest. A histological and histomorphometric analysis in rats. Int. J. Oral Maxillofac. Surg. 40, 533542. Zeltser, C., Masella, R., Cholewa, J., Mercier, P., 1989. Surgical and prosthodontic residual ridge reconstruction with hydroxyapatite. J. Prost. Dent. 62, 441448. Zerbo, I.R., Zijderveld, S.A., de Boer, A., Bronckers, A.L.J.J., de Lange, G., ten Bruggenkate, C.M., et al., 2004. Histomorphometry of human sinus floor augmentation using a porous β-tricalcium phosphate: a prospective study. Clin. Oral Implant. Res. 15, 724732. Zhang, Y., Wang, Y., 2012. Hydroxyapatite effect on photopolymerization of self-etching adhesives with different aggressiveness. J. Dent. 40, 564570. Zhang, Y.P., Din, C.S., Miller, S., Nathoo, S.A., Gaffar, A., 1995. Intra-oral remineralization of enamel with a MFP/DCPD and MFP/silica dentifrice using surface microhardness. J. Clin. Dent. 6, 148153. Zhang, W., Walboomers, X.F., Jansen, J.A., 2008. The formation of tertiary dentin after pulp capping with a calcium phosphate cement, loaded with PLGA microparticles containing TGF-β1. J. Biomed. Mater. Res. A 85A, 439444. Zhang, J.C., Lu, H.Y., Lv, G.Y., Mo, A.C., Yan, Y.G., Huang, C., 2010. The repair of critical-size defects with porous hydroxyapatite/polyamide nanocomposite: an experimental study in rabbit mandibles. Int. J. Oral Maxillofac. Surg. 39, 469477. Zhang, Q., Zou, J., Yang, R., Zhou, X., 2011. Remineralization effects of casein phosphopeptide-amorphous calcium phosphate crème on artificial early enamel lesions of primary teeth. Int. J. Paediatr. Dent 21, 374381. Zhang, X., Li, Y., Sun, X., Kisheen, A., Deng, X., Yang, X., et al., 2014. Biomimetic remineralization of demineralized enamel with nano-complexes of phosphorylated chitosan and amorphous calcium phosphate. J. Mater. Sci. Mater. Med. 25, 26192628. Zhang, L., Weir, M.D., Chow, L.C., Antonucci, J.M., Chen, J., Xu, H.H.K., 2016a. Novel rechargeable calcium phosphate dental nanocomposite. Dent. Mater. 32, 285293. Zhang, L., Weir, M.D., Chow, L.C., Reynolds, M.A., Xu, H.H.K., 2016b. Rechargeable calcium phosphate orthodontic cement with sustained ion release and re-release. Sci. Rep. 6, 36476 (11 pp.). Zhang, K., Cheng, L., Weir, M.D., Bai, Y.X., Xu, H.H.K., 2016c. Effects of quaternary ammonium chain length on the antibacterial and remineralizing effects of a calcium phosphate nanocomposite. Int. J. Oral Sci. 8, 4553. Zhao, J., Liu, Y., Sun, W.B., Zhang, H., 2011. Amorphous calcium phosphate and its application in dentistry. Chem. Cent. J. 8, 40 (7 pp.). Zhao, J., Liu, Y., Sun, W.B., Yang, X., 2012. First detection, characterization, and application of amorphous calcium phosphate in dentistry. J. Dent. Sci. 7, 316323. Zhen, T., Hongkun, W., Anchun, M., Zhiqin, C., Yubao, L., 2007. Effect of apatite nanoparticles on remineralization of the demineralized human dentin. Key Eng. Mater. 330332, 13811384. Zheng, L., Yang, F., Shen, H., Hu, X., Mochizuki, C., Sato, M., et al., 2011. The effect of composition of calcium phosphate composite scaffolds on the formation of tooth tissue from human dental pulp stem cells. Biomaterials 32, 70537059. Zhou, C., Zhang, D., Bai, Y., Li, S., 2014. Casein phosphopeptide-amorphous calcium phosphate remineralization of primary teeth early enamel lesions. J. Dent. 42, 2129. Zhou, J., Chiba, A., Scheffel, D.L., Hebling, J., Agee, K., Niu, L.N., et al., 2016. Effects of a dicalcium and tetracalcium phosphate-based desensitizer on in vitro dentin permeability. PLoS One 11, e0158400. Zijderveld, S.A., Zerbo, I.R., van den Bergh, J.P.A., Schulten, E.A.J.M., ten Bruggenkate, C. M., 2005. Maxillary sinus floor augmentation using a β-tricalcium phosphate
452 Advanced Dental Biomaterials (Cerasorbs) alone compared to autogenous bone grafts. Int. J. Oral Maxillofac. Implants 20, 432440. Zuolin, J., Hong, Q., Jiali, T., 2010. Dental follicle cells combined with beta-tricalcium phosphate ceramic: a novel available therapeutic strategy to restore periodontal defects. Med. Hypotheses 75, 669670. https://en.wikipedia.org/wiki/Dental_public_health (accessed May 2018.). https://en.wikipedia.org/wiki/Dentifrice (accessed May 2018.). https://en.wikipedia.org/wiki/Dentin_hypersensitivity (accessed May 2018.). https://en.wikipedia.org/wiki/Dentistry (accessed May 2018.). https://en.wikipedia.org/wiki/Endodontics (accessed May 2018.). https://en.wikipedia.org/wiki/Mucositis (accessed May 2018.). https://en.wikipedia.org/wiki/Orthodontics (accessed May 2018.). https://en.wikipedia.org/wiki/Prosthodontics (accessed May 2018.). https://en.wikipedia.org/wiki/Specialty_dentistry (accessed May 2018.). Further reading Kamakura, S., Sasano, Y., Homma, H., Suzuki, O., Kagayama, M., Motegi, K., 2001. Experimental oral pathology: implantation of octacalcium phosphate nucleates isolated bone formation in rat skull defects. Oral Dis. 7, 259265. McDonagh, M.S., Kleijnen, J., Whiting, P.F., Wilson, P.M., Sutton, A.J., Chestnutt, I., et al., 2000. Systematic review of water fluoridation. Br. Med. J. 321, 855859.
Bioactive glasses—structure and applications 17 Imran Farooq1, Saqib Ali1, Shehriar Husain2, Erum Khan3,4 and Robert G. Hill5 1 Department of Biomedical Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Saudi Arabia, 2Department of Dental Materials Science, Jinnah Sindh Medical University, Karachi, Pakistan, 3Bhitai Dental and Medical College, Liaquat University of Medical and Health Sciences, Jamshoro, Pakistan, 4Faculty of Dentistry, King Abdulaziz University, Jeddah, Saudi Arabia, 5Dental Physical Sciences, Institute of Dentistry, Barts and The London School of Medicine and Dentistry, Queen Mary University of London, London, United Kingdom Chapter Outline 17.1 Introduction 454 17.2 Bioactivity of glasses 454 17.2.1 Mechanism of action 455 17.2.2 Solubility 455 17.3 Factors affecting apatite formation 456 17.4 Composition of different bioactive glasses 456 17.4.1 Silicate-based bioactive glasses 457 17.4.2 Borate-based bioactive glasses 458 17.5 Methods of synthesis 460 17.6 Clinical applications of bioactive glasses 17.6.1 17.6.2 17.6.3 17.6.4 17.6.5 17.6.6 17.6.7 17.6.8 17.6.9 460 Bone graft substitute 461 Bone regeneration 461 Drug delivery system 462 Coating of implants 463 Use in toothpastes 463 Antibacterial activity 465 Role in minimal invasive dentistry 465 Bioactive glass scaffolds 465 Particle size of bioactive glasses and its effect on various clinical applications 468 17.7 Future of bioactive glasses 17.8 Conclusion 470 References 470 Further reading 476 469 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00017-7 Copyright © 2019 Elsevier Ltd. All rights reserved.
454 17.1 Advanced Dental Biomaterials Introduction “A material is said to be bioactive, if it gives an appropriate biological response and results in the formation of a bond between material and the tissue” (Farooq et al., 2012). Bioactive glass (BG) is a biomaterial that was introduced by Prof. Larry Hench in 1969 (Hench, 2006), with the intent of developing a biocompatible material that forms an intimate bond with the bone. The first glass discovered is known as 45S5 (Bioglass), having a glass composition of 46.1 mol.% SiO2, 24.4 mol.% Na2O, 26.9 mol.% CaO, and 2.6 mol.% P2O5 (Kobayashi et al., 2010). BG has the capability to interact with the body tissues to form a resilient bond, and its controlled degradation over time is useful in releasing therapeutic ions, which can help the bone regeneration (Fuchs et al., 2015; Ali et al., 2014). Safety of these glasses was a concern, so various studies were performed, and these glasses were graded safe to be used by US Food and Drug Administration (FDA) (Paolinelis et al., 2008). BGs are synthetic and osteoconductive materials that form a hydroxycarbonate apatite (HCA) layer at the site, after the dissolution of the glass (Jones, 2013). This HCA is comparable to the bone mineral, and it interacts with collagen fibers of the bone to form an adherent interface between the material and the tissue (Jones, 2013; Nejatain et al., 2017). Recently, BGs have been used for different dental applications such as in toothpaste to enhance enamel remineralization and occlude open dentinal tubules (Farooq et al., 2012). Recent modifications in the structure of BG have resulted in various formulations being available, such as high phosphate-based glasses, borate-based glasses, and fluoride-containing BGs (Khalid et al., 2017), and these glasses are different from the conventional 45S5 glass. Borate and phosphate glasses have rapid solubility (which can be altered from several hours to several months based on the end application) (Fu et al., 2010a), whereas fluoride-containing glasses can form fluorapatite (FAP) layer, on the surface, which promotes remineralization and is more durable in the mouth (Farooq et al., 2013). This chapter focuses on the structure of different BGs and their clinical applications. Apart from these objectives, this chapter also reviews the methods of their synthesis and gives an overview of bioactivity of these glasses. 17.2 Bioactivity of glasses The BGs are highly bioactive and a previous in vitro study has demonstrated the formation of apatite by the new composition of glasses in Tris-buffer solution, in as little as 6 hours (Farooq et al., 2013). These glasses stimulate cell cycling in vitro, which results in stimulation of osteoblasts, as a result of which more mineralized tissue is formed in a shorter time (Xynos et al., 2000). Due to ion exchange, BGs can raise the pH of the solution in which they are present, thus exhibiting antibacterial effects (Gubler et al., 2008). Some groups of these glasses, such as silvercontaining BGs (Newby et al., 2011), release important ions, such as silver, which
Bioactive glasses—structure and applications 455 can augment these antibacterial effects (Kwakye-Awuah et al., 2008). In vivo studies are important as they help in predicting the clinical performance of a material. An in vivo survey carried out in rabbits showed better performance of 45S5 as compared to the synthetic hydroxyapatite (HAP) regarding bone regeneration (Oonishi et al., 2000). 17.2.1 Mechanism of action The bioactivity of these glasses is dependent on their mechanism and speed of action. Briefly, the reactions to form HAP involve ion exchange of Ca21 and Na1 ions for H1 ions from the solution, a consequent increase in the pH of the solution, the formation of silanol (Si-OH) bonds on the surface of the glass, and the resulting formation of a silica-rich layer, degradation of silica (due to increased pH), and then then the development of a layer of amorphous CaO P2O5 on the silica-rich layer, which then crystallizes as HAP due to the absorption of hydroxyl and carbonate ions (Jones, 2013). 17.2.2 Solubility The glass composition plays a vital role in determining the solubility and bioactivity of BGs. At this point, understanding some other terms such as glass transition temperature (Tg) and network connectivity (NC) becomes essential. The Tg is defined as a range of transformation when an amorphous solid is converted into a supercooled liquid on heating (Dudowicz et al., 2005). The degradation rate of the glass and its strength can be assessed through Tg (O’Donnell, 2011). There is a presence of a significant processing window between the Tg and peak crystallization temperature of glass warrants that the glass will not crystallize during quenching (Dimarzio and Gibbs, 1959). The presence of crystalline phases reduces the exchange of ions, and thus bioactivity is reduced (Ali et al., 2014). The NC can be defined as the mean number of bridging oxygen bonds per silicon atom (Hill and Brauer, 2011). It can be used to analyze the solubility and bioactivity of a glass, as a low NC implies that the glass has lower Tg, but a high solubility and bioactivity (Hill, 1996). The silica in the glass is considered a network former, which can hold the glass structure together (Srivastava et al., 2012). Therefore a lower content of silica can ensure faster dissolution and more rapid bioactivity (Jones, 2013). Fluoride is an essential ion when it comes to dentistry as it remineralizes tooth structure and prevents demineralization (Featherstone, 2000). The inclusion of fluoride in the BG composition decreases its Tg, which means that the glass will have a reduced hardness, but it will be more bioactive (Farooq et al., 2012). Sodium oxide (Na2O) can also affect the properties of BGs. The Na2O is regarded as a network disrupter as its addition expands the glass network; therefore it reduces Tg and the glass becomes more bioactive (Wallace et al., 1999). In the fluoride-containing BG compositions, when phosphate content is increased, it results in maintenance of the NC and formation of FAP (Brauer et al., 2010). In an in vitro study, it was reported that
456 Advanced Dental Biomaterials high phosphate-containing BGs were able to form apatite more rapidly in Trisbuffer solution and within 6 hours, as compared with low phosphate-containing BGs (which formed apatite after 3 days) (Mneimne et al., 2011). This means that by controlling the amount of phosphate in the BG composition, reactivity, and apatite-forming ability of the BGs can be controlled (Mneimne et al., 2011). 17.3 Factors affecting apatite formation Many factors, such as different ions, can accelerate or hinder apatite formation ability of BGs. As described in the previous section, phosphate content can significantly increase apatite formation capabilities of BGs (when NC is maintained) (O’Donnell et al., 2009). Similarly, controlled addition of fluoride (Brauer et al., 2010) and substitution of calcium by strontium can positively favor the apatite formation process (Brauer, 2015). On the other hand, magnesium ions have been shown to delay or inhibit apatite formation (Diba et al., 2012), probably due to the obstruction of active growth sites on the apatite crystal surfaces (Kanzaki et al., 2000). Generally, for dental applications, faster apatite formation is required, whereas for some other medical applications, such as cartilage repair, a slower apatite formation is desirable (Brauer, 2015). Therefore the addition of different ions into BG composition can be carefully planned according to the desired outcome. 17.4 Composition of different bioactive glasses The affinity of a biomaterial to bond to mineralized biological tissue is dependent upon its composition. A range of such glass systems has been developed over the years, which include compositional modifications with certain elemental additives and variations in synthesis techniques. The compositions of 45S5 and a few essential variants are shown in Table 17.1. Table 17.1 Composition of different bioactive glasses (Khalid et al., 2017). 45S5 13 93 6P53B 58S 70S30C 13 93B1 13 93B1 P50C35N15 Na2O K2O MgO CaO SiO2 P2O5 B2O3 24.5 6.0 10.3 0 0 5.8 5.5 9.3 0 12.0 2.8 0 0 11.7 11.1 0 0 5.0 10.2 0 0 4.9 4.6 0 24.5 20.0 18.0 32.6 28.6 19.5 18.5 19.7 45.0 53.0 52.7 58.2 71.4 34.4 0 0 6.0 4.0 6.0 9.2 0 3.8 3.7 71.0 0 0 0 0 0 19.9 56.6 0
Bioactive glasses—structure and applications 457 17.4.1 Silicate-based bioactive glasses Silicate glasses are considered as a reliable material when considering the design of novel medical devices boasting specific properties. A large surface area, a high degree of purity index, and porosity within the bulk of the material contribute significantly toward the high level of reactivity exhibited when employing this material for research purposes and as a component of a biomaterial (Aguiar et al., 2009). The open-ended structure of these BGs enables the manifestation of free space within the bulk of the material. This allows for the addition of Na1, K1, Ca21, and Mg21 cations into the glass matrix. These cationic species are termed as network modifiers. Network modifiers are known to induce network disruption—a crucial preliminary step if realization of nonbridging oxygen (NBO) groups is to occur within the glass structure (González et al., 2003; Sen and Youngman, 2003). The presence of NBOs has a strong influence on the connectivity of the material as a whole. The commercial variant 45S5 (based on silicate glass) has been the focus of multiple studies pertaining to biomedical applications (Hench, 2006). The inherent properties of this version of 45S5 allow for it to bond to mineralized bone tissue—a phenomenon which has been the subject of numerous investigations for the better part of four decades and counting. The silicate glass 45S5 structure can be best visualized as a three-dimensional (3D) SiO2 network with a central silicon ion surrounded by four coordinating oxygen ions arranged in a tetrahedron configuration. The SiO2 content in silicate BG compositions is in direct relation to their chemical stability. Moreover, glass network modifiers in variable ratios are the primary structural and compositional determinants of the bioactive nature of this material. Common glass network modifiers include Na2O and CaO, whereas magnesium (Mg21) and zinc (Zn21) fall on the border of modifiers and intermediate ions (which increase the processing window to avoid crystallization) (Dietzel, 1941). The role of these modifiers is to break some part of the Si O bonds, thereby creating NBOs and disturbing the tetrahedral silicate network (Benoit et al., 2001). In other words, these network modifiers are responsible for considerably lowering the connectivity value of the SiO2 network attributed to the formation of nonbridging silicon oxygen bonds. The resulting enhanced dissolution index translates into an increased probability of ionic exchange events at the surface of the material that would eventually contribute to the formation of a biomimetic HAP layer on the surface of the implanted material (Vichery and Nedelec, 2016; Hench, 1991). Phosphate ions can increase the bioactivity of the glass (Mercier et al., 2011), and a linear relationship in terms of increased bioactivity can be observed with the addition of phosphate in BG, owing to the inherent ability of phosphate to influence the formation of apatite in living mineralized tissues (Hill and Brauer, 2011). Some studies have aimed to decipher the bonding configurations and the presence of NBO groups by utilizing spectroscopic techniques such as infrared (IR) spectroscopy and X-ray photoelectron spectroscopy (Serra et al., 2003). The results from these analyses were found to be in agreement with one another. The inclusion of network modifiers to the silica network has a direct effect on electron density of the
458 Advanced Dental Biomaterials bonding states of the silicon and oxygen atomic structure (Serra et al., 2003). Moreover, by carrying out these analyses, the investigators were able to demonstrate a rise in the proportion of alkali-earth elements in the silica network, which leads to a concurrent cleavage of Si O Si (bridging oxygen) by Si O NBO groups (nonbridging silicon oxygen groups). This has far-reaching implications in terms of eliciting a reproducible and robust biological response at the material interface when submerged in body fluids (Peitl et al., 2001). The role of NBO groups during the initial stages of bioactivity of BGs is of much importance. IR spectroscopy has emerged as a powerful tool that yields useful information in terms of developing a quantifiable database pertaining to the concentration of Si O NBO groups and ultimately their bond strength, which may be ascertained from the relative intensity of the respective IR absorption bands. The ideal ratio between the Si O NBO and the Si O groups, estimated to be at $ 1 as derived from IR absorption band intensity, is critical for realization of an effective ion exchange and subsequent dissolution of the silica—ultimately contributing to the formation of an SiO2-rich layer on the material surface (Serra et al., 2002). Therefore a definite correlation exists between the nature and type of NBO functional groups and the formation of calcium phosphate rich layers. Hence spectroscopic techniques have established themselves as reliable techniques for gaining a deeper understanding of the pivotal role of network modifiers through alteration of the BG structure and subsequently their bioactive behavior. When the 45S5 BG is implanted, a chemical degradation reaction process releasing cationic species such as Na1 and Ca21 ensues. This paves the way for the formation of a carbonate-substituted HAP-like material. This leads to a subsequent release of silicon in the guise of silicic acid (Si(OH)4) (Lai et al., 2002). The low SiO2 and high Na2O and CaO content are important compositional features determining the bioactivity of 45S5 (Huang et al., 2006a). Many investigators still consider the original 45S5 version as the gold standard in the realm of BG for hard mineralized tissue regeneration. However, issues pertaining to the impact of time-dependent release of degradation products especially during the initial stages, when there is an uptake in the alkaline earth metal sourced cationic species concentrations of Na1 and Ca21 from the bulk of the material with accompanying pH changes (Tahriri et al., 2017), have yet to be fully elucidated in light of their biological niche, toxicity, and subsequent dispersion from the site of implantation. Another limitation of 45S5 glass is its incredible performance in high load-bearing applications such as bone scaffolds (Hench and Jones, 2015). Research is still going on to achieve tougher scaffolds that can serve as an appropriate scaffolding material. 17.4.2 Borate-based bioactive glasses The trace element boron is essential for maintaining bone health. The first borosilicate glass formulation to be considered for biomedical applications was conceived in 1990 (Brink, 1997). These glasses are considered as a reactive species, which accounts for their low chemical stability index. This translates to a more
Bioactive glasses—structure and applications 459 rapid conversion rate into HAP for this group of BG as compared to their silicabased counterparts (Yao et al., 2007). The substitution of SiO2 with B2O3 allowed a significant escalation in the conversion rate of borate glass to HAP postimmersion in aqueous phosphate solutions (Huang et al., 2006b). Although the mechanism by which borate-based BG undergoes conversion to a layer of apatite is similar to its silica-based equivalent, in that a borate-rich layer ensues in place of a silicate-rich layer, borate-based BGs are more amenable to precise tinkering with the degradation profile and sintering behavior, which can be achieved over a range of time lines in contrast to silicate-based BGs (Yang et al., 2012). Spanning the course of a few decades, a significant amount of research has been conducted, and consequently, key data have been generated from studies pertaining to the behavior and properties of the original formulation of 45S5 BG. On the other hand, probing research questions exploring borate-based compositions are only just beginning to surface in the literature (Rahaman et al., 2011; Kaur et al., 2014) An innovative take on developing BG powders was the development of sphericalshaped borate-based BG powders for replacement of bone tissue using spray pyrolysis (Cho and Kang, 2009). Evidence for the resulting crystalline structure of the powders prepared in this way was gathered using X-ray diffraction (XRD) analysis. The preparation technique had a profound influence on the phase transitions observed in the glass powder. Powders synthesized (using pyrolysis) below a temperature range of 1400 C predominantly exhibited a crystalline phase that dominated the composition. A low rate of turnover to the glass phase was attributed, in this instance, to a shorter “stay” time within the confines of the alumina reactor. The opposite is true in the case of increased preparation temperatures of 1500 C, in that the amorphous phase enjoyed a majority throughout the glass substructure summed up by a reduction in sharpness of the XRD peaks. In this way ideal temperature parameters for preparing spherical bioactive borate-based glass powders with a dense inner core were identified to stand at 1200 C. A standing criticism leveled against silicate-based BG stems from an incomplete conversion process for yielding a calcium phosphate material postimplantation in vivo. Indeed corresponding XRD patterns between different glass samples have in fact revealed the strongest intensity hydroxyapatite (HA) peaks, an indication of enhanced bioactive potential, in glass samples with increasing B2O3:SiO2 molar ratio postimmersion in a phosphate solution (Fu et al., 2007). These findings were in agreement with pH and weight loss studies—solutions containing glass samples having increased B2O3 content consistently showing increased pH values and more significant weight loss with immersion time compared to neat borate- and silicate-based glass samples (Huang et al., 2006b). The significant disparity in the reactivity, pH, and observable weight change rate between silicate and boratebased glass is primarily attributed to their overall network structure. The BO3 trihedron chains possess a threefold coordination number that hinders the formation of a 3D network structure when compared to their silicon-based counterpart (Cheng et al., 2009). This has a profound impact in lowering the chemical durability of the borate glass network structure translating to an accelerated dissolution rate.
460 Advanced Dental Biomaterials Moreover, in line with the suggestion put forth by prominently featured works, formation of a layer of HA on the surface of a biomaterial under certain parameters in vitro can be extrapolated as its bioactive potential in vivo (Hench, 1998; Zadpoor, 2014). The formation of a dense microstructured HAP as evidenced by scanning electron microscopy (SEM)-based surface morphological analysis on borate-based BG samples as reported previously (Yao et al., 2007) can be evaluated as evidence for bioactive potential in the same light. It is important to mention at this point the deviations in the mechanism that allows for the conversion of boratebased BG and silicate-based BG to carbonate-substituted HAP. Even though a large part of the dissolution precipitation reaction process for both glasses is similar for the most part, pure borate-based BG conversion to carbonate-substituted HAP differs considerably in that it does not involve the formation of the SiO2-rich layer at any point of the conversion and dissolution sequence when assessed postimmersion in a phosphate-rich solution (Huang et al., 2006b). Rather a complete ionic dissolution of the borate-based glass into the solution ensues due to a simultaneous infiltration and breakdown of the B O glass network structure by the phosphate solution. The process continues until the full conversion of the borate glass to carbonatesubstituted HAP postleaching and reaction of the Ca21 ions with the PO432 in solution. The rate at which borate-based BG undergoes conversion to carbonatedsubstituted HAP, a reliable scale for a measure of its degree of bioactivity, is heavily predicated on the B2O3:SiO2 ratio of the glass, with a higher B2O3 level translating to a higher borate glass conversion rate to HAP in this case. 17.5 Methods of synthesis The BGs require a high standard of raw material purity prior to their preparation. This is mainly because the quality of the materials heavily influences the quality of the end product at the starting point. Pure silica sand (quartz), reactive grade carbonates of sodium and/or potassium, etc. are some of the common ingredients required in weighted amounts. BGs are usually of a soft nature. This usually allows for their easy shaping and sizing. Two common methods of preparing BGs include melt quenching (at temperatures exceeding 1200 C) and sol gel method (converting the system from a liquid “sol” into a solid “gel”) (O’Donnell, 2012). However, going into the details of these methods is beyond the scope of this chapter. 17.6 Clinical applications of bioactive glasses The BGs are different from conventional glasses as they possess numerous features which ensure their wide-ranging clinical applications. These glasses are biocompatible, osteoconductive, and can bond firmly to the tissue (Toosi and Behravan, 2017). The conventional silicate-based glasses are composed of phosphate and calcium in somewhat similar proportion to that of the bone HAP. Due to these abilities, they
Bioactive glasses—structure and applications 461 can be utilized in a broad range of medical and dental applications. The BGs can be divided into different families having different compositions, which can be used for a specific function or to achieve a desired clinical outcome (Rao and Ravindranadh, 2016). These glasses can be effectively used for replacement, repair, or reconstruction of different body parts such as bone and teeth (Baino et al., 2014). Some of the most important applications of BGs are discussed in the following subsections. 17.6.1 Bone graft substitute When bone is lost due to an infection, disease process, or trauma, the bone can be replaced with a bone graft. Clinically, BG has been in use for more than a decade as a synthetic bone graft. In orthopedics, it is as a product called Novabone (Elshahat, 2006) and in maxillofacial surgeries as PerioGlas (Fetner et al., 1994). A very common BG which has been used as a bone graft is 45S5 (Rao and Ravindranadh, 2016). The FDA in 2005 permitted the use of 45S5 for osteostimulation (Hench, 1998). In another study it was reported that BGs have an osteostimulatory effect along with an osteoconductive function (Boccaccini et al., 2010), which other osteoconductive bioceramics usually lack (Gerhardt and Boccaccini, 2010). Also, their ability to bond to osseous tissues is much superior as compared with other alloplastic materials (Wilson et al., 1993). It is a reality that BGs do not have ideal mechanical properties, but the addition of Na2O into SiO2 CaO BG composition improves its biological absorbability and mechanical capability (Chen et al., 2010). A previous study reported that BG scaffolds can totally resorb in 6 months with little inflammatory response (Moimas et al., 2006), demonstrating their superiority over other bone graft materials which can cause adverse reactions as well (Wang and Yeung, 2017). An earlier study that was conducted on animal models where properties of BG and HAP were compared concluded that it is very easy to manipulate the composition of BG for specific uses, and BG takes less time in response generation as compared to HAP (Oonishi et al., 1997). The BGs when used as a bone graft have also shown admirable bone healing properties in numerous follow-ups of long-term studies (Khalid et al., 2017; Van-Gestel et al., 2015). 17.6.2 Bone regeneration Bone regeneration is another important clinical application, and BGs have a greater filler effect than that of an autogenous bone (Heikkilä et al., 1995). In a previous study by Macedo et al. (2004) two different compositions of BGs were used to study bone formation in tibiae of rats, and it was reported that both compositions of BGs promoted bone formation. It has been reported earlier that significant bone regeneration can be promoted by BGs in vitro as they have osteostimulatory effects (Hench, 2013). The presence of BGs in the treatment of large bony defects can result in infection-free bone regeneration (Stoor et al., 2017). In an animal study, periodontal defects were treated with particles of BG, which triggered bone mineralization (Felipe et al., 2009).
462 Advanced Dental Biomaterials One of the advancements in BG research is control of its degradation rate with the manipulation of its composition. Borate-based BGs are very useful in bone regeneration because of the variation of controlled degradations rates along with the ease of their fabrication. The compositional flexibility and easy manipulation of the glass composition can make it a source of different elements such as copper, fluoride, or boron which can encourage growth of the bone (Toosi and Behravan, 2017). 17.6.3 Drug delivery system Researchers have always been looking for an innovative drug delivery system to have superior medication control with a prolonged action. Assuming that a certain medication or molecule will reach a specific site without any secondary reactions and will perform the desired reaction, the drug delivery system becomes supremely important to the researchers. These glasses show larger flexibility in terms of compositional manipulation, making them independent of any specific stoichiometry. Thus ions of various concentrations having different therapeutic properties can be incorporated in conventional compositions. These ions can be released during the process of dissolution, and they can execute their desired therapeutic function in the human body (Hoppe et al., 2011). In a study on pulp capping agents containing BG, there was a greater occurrence of development of a properly positioned dentin bridge (Stanley et al., 2001). As discussed earlier, borate-based BGs are used in bone regeneration and can also be used in the treatment of infection of bone, where they act as a drugreleasing substrate (Liu et al., 2010; Jia et al., 2010). An earlier in vitro study comparing borate-based BGs with 45S5 reported that higher content of B2O3 improved the conversion rate to HAP, as compared to 45S5 particles, which stopped after only partial conversion of particles to HAP (Brown et al., 2009). Another study has also revealed that teicoplanin-loaded borate BG implants could be helpful in treating chronic osteomyelitis in animals (Zhang et al., 2010), so it can be predicted that they will be equally useful in humans as well. In a previous study conducted on animal models, BG porous blocks were used for delivering antibiotics in treating osteomyelitis (Kundu et al., 2011). After 2 years of treatment of infected arthroplasty, some outstanding results were witnessed, and osteogenesis was triggered by the implant material, which was evident on radiography where complete radiological replacement of the osseous defects was seen (Kawanabe et al., 1998). Excellent bone integration and biocompatibility were seen in BG implants infused with gentamicin sulfate, which released gentamicin into local osseous tissues, and during the resorption process, these implants also promoted the growth of the bone (Meseguer-Olmo et al., 2006; Arcos et al., 2001). In another study BG combined with tetracycline and BG combined with tetracycline:beta-cyclodextrin were able to demonstrate considerable bacteriostatic activity with little effect on the bioactivity of the glass itself (Domingues et al., 2004).
Bioactive glasses—structure and applications 463 17.6.4 Coating of implants The success of dental implants is dependent on several factors, of which the implant material is of utmost importance (Najeeb et al., 2015). Over time, the use of metallic materials has grown significantly due to their ideal mechanical properties (Roessler et al., 2002). Metals such as titanium, cobalt, and stainless steel (SS) grade 316L are used in the preparation of implants (Garcia et al., 2004). SS has been the choice in orthopedic implants as it is quite economical (Fathi et al., 2003). But this material, in long-term use, is very much prone to unwanted biological reactions, thus leading to failure of the implant mechanically. Titanium (Ti) and Ti-alloys have good mechanical and physical properties which makes them the material of choice for implant applications, and currently, they are the most widely used implant material (Najeeb et al., 2017). It is a challenge to get a perfect interface between the dental/orthopedic implant and the bone, because the development and preservation of viable bone opposing the biomaterial surfaces are very essential for the success and the stability of noncemented dental/orthopedic implants (Moimas et al., 2006). There has been the initiation of extensive research into coatings of metallic implants with BGs as BGs offer worthy bone bonding ability with controlled surface reactivity (Greenspan, 1999; Hench and Andersson, 1993; Ferraris et al., 1996). It was reported earlier that a bond is formed between implanted bioceramics and natural tissues as a result of an active biological layer that is formed on the surface of the implant by bioactive materials (Rahaman et al., 2011). In another study performed on the human jaw bone, there was a comparison of Ti-alloy dental implants coated with BG and HAP, and it was concluded that both materials were nontoxic and biocompatible, and BG demonstrated good osseointegration properties, comparable to that of HAP (Mistry et al., 2011). It is a common practice nowadays to use implants coated with BGs for achieving good osseointegration with the alveolar bone. Using BG in dental implants as coating material shows superior bone regeneration along with better adherence to the metal surface of the implant (Koller et al., 2007). 17.6.5 Use in toothpastes The use of BGs in toothpaste has been increased in the last two decades. They have been incorporated in various toothpastes because of their potential to treat dentin hypersensitivity and to remineralize tooth structure (Abbasi et al., 2015). BGs’ first commercial use in oral health was to treat dentin hypersensitivity, keeping in mind that they have the capability to occlude by a HAP layer the uncovered dentinal tubules that are the main cause of the pain and sensitivity (Burwell et al., 2009; Dababneh et al., 1999). The BG-based dentifrice has showed superior tubule occlusion properties on dentin discs of extracted human teeth when compared with a standard fluoride-based dentifrice, before and after a citric acid challenge, in an in vitro SEM study (Fig. 17.1) (Farooq et al., 2015).
464 Advanced Dental Biomaterials Figure 17.1 Scanning electron microscopy micrograph of a dentin disc that was brushed with bioactive glass containing toothpaste, post citric acid challenge at 6000 3 , scale bar: 20 mm. Source: Adapted from Farooq, I., Moheet, I.A., AlShwaimi, E., 2015. In vitro dentin tubule occlusion and remineralization competence of various toothpastes. Arch. Oral. Biol. 60 (9), 1246 1253. Lynch et al. (2012) also reported admirable capability of fluoride-containing BGs in occluding open dentinal tubules in an in vitro study. In recent times it has also been established that BGs can act as a remineralizing agent (Mehta et al., 2014; Reynolds, 2008). The role of fluoride in prevention and its application is of interest in dentistry. The addition of fluoride to BGs not only enhances remineralization, but it also prevents demineralization of enamel and dentin (O’Donnell, 2011). Farooq et al. (2018) performed a study to analyze remineralization potential of a novel dentifrice consisting of fluoride-containing BG (BiominF) with that of a dentifrice containing only BG (Novamin). Enamel blocks were demineralized with citric acid and then remineralized using toothpaste slurries (a mixture of toothpaste with artificial saliva). After 5 minutes and 24 hours, mean enamel volume changes were evaluated by microcomputed tomography, and mean surface loss or gain was investigated using a profilometer. It was demonstrated through the results of this study that BiominF specimens showed better remineralization potential, especially after 5 minutes. The possible reason for the better performance of BiominF could be that it contains fluoride in its BG composition with high PO432 content, which could serve as a source of the delivery of all essential ions (Ca21, PO432, and F2) together to form FAP, rather than fluorite (CaF2) (Mneimne et al., 2011). Another difference between Novamin and BiominF is the difference between their particle size (BiominF being smaller than Novamin), which can result in better tubule occlusion. The difference between the particle size of these two materials is shown in Table 17.2.
Bioactive glasses—structure and applications 465 Table 17.2 Showing particle size difference between Novamin and BiominF (http://www.biomin.co.uk/science/bioactive-glasses/biomintm-vs-novaminr). Particle size Novamin (µm) BiominF (µm) D10 D50 D90 0.177 14.47 45.55 0.62 5.92 0.62 17.6.6 Antibacterial activity Besides remineralization, these glasses can show antibacterial activity as well by raising the pH of an aqueous solution. It is very common to use antimicrobials in certain dental procedures used in the fields of periodontics and endodontics (Khalid et al., 2017). In a previous study the insertion of BG in periodontal defects has shown inhibition of bacterial colonization due to the rise in the pH and possibly because it provided calcium ions to the defective site (Allan et al., 2001). 17.6.7 Role in minimal invasive dentistry BGs have a shown significant potential in minimally invasive dentistry. BG powder can also be utilized in cutting cavities with air abrasion, causing less damage to tooth enamel as compared to conventional cavity preparation performed by a highspeed handpiece. Alumina powder is commonly used in air abrasion machines, as it has coarse particles (Fig. 17.2) (Hassan et al., 2017), but it is an inert material having no benefit other than quick cutting. Alumina can be replaced with BGs, as they possess apatite-forming ability, and also have coarse angular particles with sharp edges (Fig. 17.3) for efficient cutting. Farooq et al. (2013) synthesized a few new, different compositions of BGs with fluoride and reported the formation of apatite in vitro within 6 hours (which was quicker than traditional 45S5) and with comparable cutting efficiency to that of alumina (Fig. 17.4). 17.6.8 Bioactive glass scaffolds Tissue engineering has developed extensively in the past two decades as an approach for the repair and regeneration of tissues and organs which are lost or damaged due to traumatic injuries, diseases, or the aging process (Nerem, 1991). Autografts, bone allografts, synthetic biomaterials, and metallic implants have been reported in the literature for the rejuvenation of tissue and bony defects. The limitations of current treatments and the higher costs have encouraged interest in the engineering of new bone substitutes. The target of scaffold-based bone tissue engineering is to repair and regenerate bony defects with minimum side effects (Hutmacher et al., 2007). A scaffold is a porous structure which, preferably, should direct new tissue formation by providing a matrix with interconnected porosity and
466 Advanced Dental Biomaterials Figure 17.2 Scanning electron microscopy micrograph of alumina particles at 500 3 , scale bar 300 μm. Source: Adapted from Hassan, U., Farooq, I., Moheet, I.A., AlShwaimi, E., 2017. Cutting efficiency of different dental materials utilized in an air abrasion system. Int. J. Health. Sci. (Qassim) 11 (4), 23 27. Figure 17.3 Scanning electron microscopy micrograph of 45S5 at 1000 3 , scale bar 100 μm. tailored surface chemistry for the cell growth and proliferation and the transport of nutrients and metabolic waste (Hansbrough et al., 1994). Ideally a scaffold should mimic the bone morphologically, structurally, and functionally in order to augment integration with its surrounding tissues (Johnson and Herschler, 2011; Karageorgiou and Kaplan, 2005).
Bioactive glasses—structure and applications 467 Figure 17.4 FTIR spectra for a BG batch which shows formation of apatite at 6 h. BG, Bioactive glass; FTIR, Fourier-transform infrared spectroscopy. Source: Adapted from Farooq, I., Tylkowski, M., Müller, S., Janicki, T., Brauer, D.S., Hill, R.G., 2013. Influence of sodium content on the properties of bioactive glasses for use in air abrasion. Biomed. Mater. 8 (6), 065008. BGs have etching characteristics as a scaffold material for bone tissue engineering. These glasses undergo specific reactions that lead to the formation of amorphous calcium phosphate or crystalline HAP phase on the glass surface, which results in a stable bonding with the surrounding tissue (Jones et al., 2006). The activation of expression of osteogenic genes (Xynos et al., 2001) and stimulation of angiogenesis have been reported by BGs (Gorustovich et al., 2009). Although the low mechanical strength of BG scaffolds limits their usage for the repair of defects in load-bearing bones (Yunos et al., 2008), researchers have tried to overcome this issue by optimizing the composition, sintering conditions, and processing, and now BG scaffolds can be created with predesigned pore architecture and with strength comparable to human trabecular and cortical bones (Liu et al., 2011). In addition to strength and elastic modulus other mechanical properties such as reliability and fracture toughness are also of decisive importance for scaffolds implanted in loadbearing bone defects. As mentioned earlier, BG scaffolds can be created with the preferred compressive strength for the restoration of load-bearing bone defects. Still their practice in these applications may be restricted due to their intrinsic brittleness, which is also called fracture toughness. Adding a biocompatible polymer coating is suggested to improve the toughness of BG scaffolds by providing a crack bridging mechanism through the polymer layer for energy dissipation. Regardless of its innate brittleness, BG has numerous appealing characteristics to be used as a scaffold material in bone tissue engineering, especially novel innovative BGs
468 Advanced Dental Biomaterials established on borosilicate and borate compositions that have displayed the capability to develop new bone formation (Rahaman et al., 2011) with the least side effects. Literature reports various methods for the fabrication of BG scaffolds, which include sol gel, freeze casting, thermal bonding of particles, polymer foam replication, fibers or spheres, and solid free-form fabrication. The capability of BG scaffolds to support cell function and proliferation in laboratories and clinical tissue ingrowth has been publicized in many studies (Fu et al., 2010a; Goodridge et al., 2007; Fu et al., 2010b; Zhao et al., 2008). Regardless of brittleness, BGs have a distinctive set of properties such as formation of HAP layer, intimate bond formation with hard tissues, and release of ions during the degradation process, which are favorable for osteogenesis, angiogenesis, and chondrogenesis (Rahaman et al., 2011). Prospective research is expanding on the favorable properties of BGs, considering brittleness through innovative scaffold design and processing, predominantly when used for the repair of load-bearing bones. 17.6.9 Particle size of bioactive glasses and its effect on various clinical applications Before concluding this section, another key area to discuss is the importance of particle size of BGs and its impact on various clinical applications. In general, a smaller sized BG particle is preferred as it can produce better results. A few common BG-based materials along with their particle sizes and uses are summarized in Table 17.3. In dentistry an increased particle size of the glass in toothpaste can cause more abrasion of the enamel (Mahmood et al., 2014). Therefore one solution is to use smaller sized particles in dentifrices. To comminute the glass frit, percussion milling (ball milling) is usually performed (Mahmood et al., 2014). But grinding the particles to a smaller size (after milling) usually involves higher costs; therefore another alternative is to reduce the abrasivity of the existing glass by incorporating Table 17.3 Showing particle size of different bioactive glass based products used for various clinical application. Product Particle size Uses Novamin PerioGlas (D50 value) of 18 μm 90 710 μm Biogran 300 360 μm Used in toothpaste to treat hypersensitivity by blocking open dentinal tubules Used for bone regeneration around the tooth or bone repair in the jaw for anchoring implants Used as a bone graft in jaw defects Source: Values adapted from Jones, J.R., 2013. Review of bioactive glass: from Hench to hybrids. Acta Biomater. 9 (1), 4457 4486.
Bioactive glasses—structure and applications 469 ions such as fluoride, which can produce a softer glass that will form FAP (Mneimne et al., 2011). Wilson and Low (1992) reported the effect of different sizes of 45S5 particulates on the regeneration of bone in periodontal defects produced in a monkey model. The study demonstrated the ideal rate of bone repair when a range of 45S5 particle sizes were used. In another study Ajita et al. (2015) studied the effect of the size of nanostructured BG particles on mouse mesenchymal stem cell (MSC) proliferation. It was concluded from this study that smaller sized nano-BG particles were able to increase proliferation of MSCs, thus implicating that they could produce desirable results in various clinical applications. 17.7 Future of bioactive glasses The development of composite materials combining biodegradable polymers (synthetic and natural) with nanoscale BG particles or fibers is emerging as a robust approach toward third-generation bioactive materials. The biomedical applications of these novel materials are bound to expand, for example, as bone filler materials, temporary orthopedic implants, as 3D biocompatible scaffolds in the field of tissue engineering (Guarino et al., 2007), and in the dental industry for tooth remineralization, dentin regeneration, and reconstruction of bony defects. Composite materials add strength and bioactivity through an inorganic bioactive filler while polymers enhance flexibility and capacity to distort under loads (Boccaccini et al., 2010). The 45S5 particulate has been used in many oral care products for the treatment of tooth hypersensitivity as 45S5 particles stick to the dentin by forming a HAP layer that is comparable in composition to tooth enamel, and it blocks the dentinal tubules, thus relieving the pain for extended periods (Gillam et al., 2002). Dental care with 45S5 is not limited to toothpaste only. Sodium bicarbonate abrasives are used to remove stains by dentists, but the use of 45S5 through air polishing can stimulate remineralization; thus it reduces dentin hypersensitivity along with better stain removal and results in much whiter teeth (Banerjee et al., 2010). The use of BG, as or in a restorative material, is a debatable topic, as it is meant to degrade in an aqueous solution (saliva in case of the oral cavity). However, a recent study by Khvostenko et al. (2016) reported that the use of BG as filler for resin-based composite restorations could decrease biofilm penetration into marginal gaps of simulated tooth restorations, thus implicating that composite restorations containing BG can reduce the development and propagation of secondary tooth decay at the margins of the restoration. Another potential area (related to dentistry) for the BG is their use as pits and fissure sealant. Previously, Yang et al. (2016) reported in an in vitro study that BGcontaining sealants can inhibit the demineralization of the enamel surface within microgaps between the material and the tooth when disclosed to a cariogenic environment.
470 17.8 Advanced Dental Biomaterials Conclusion BGs make a firm bond with the host and have the ability to degrade and form apatite in physiological solutions. The easy manipulation of their composition makes them the material of choice for extensive clinical applications. With their current use in different medical and dental applications, an optimistic future for these glasses can easily be anticipated. References Abbasi, Z., Bahrololoom, M.E., Shariat, M.H., Bagheri, R., 2015. Bioactive glasses in dentistry: a review. J. Dent. Biomater. 2 (1), 1 9. Aguiar, H., Serra, J., González, P., León, B., 2009. Structural study of sol-gel silicate glasses by IR and Raman spectroscopies. J. Non-Cryst. Solids 355, 475 480. Ajita, J., Saravanan, S., Selvamurugan, N., 2015. Effect of size of bioactive glass nanoparticles on mesenchymal stem cell proliferation for dental and orthopedic applications. Mater. Sci. Eng. C Mater. Biol. Appl. 53, 142 149. Ali, S., Farooq, I., Iqbal, K., 2014. A review of the effect of various ions on the properties and the clinical applications of novel bioactive glasses in medicine and dentistry. Saudi Dent. J. 26 (1), 1 5. Allan, I., Newman, H., Wilson, M., 2001. Antibacterial activity of particulate Bioglasss against supra- and subgingival bacteria. Biomaterials 22 (12), 1683 1687. Arcos, D., Ragel, C.V., Vallet-Reg, M., 2001. Bioactivity in glass/PMMA composites used as drug delivery system. Biomaterials 22 (7), 701 708. Baino, F., Novajra, G., Vitale-Brovarone, C., 2014. Bioceramics and scaffolds: a winning combination for tissue engineering. Front. Bioeng. Biotechnol. 3, 202. Banerjee, A., Hajatdoost-Sani, M., Farrell, S., Thompson, I., 2010. A clinical evaluation and comparison of bioactive glass and sodium bicarbonate air-polishing powders. J. Dent. 38 (6), 475 479. Benoit, M., Ispas, S., Tuckerman, M.E., 2001. Structural properties of molten silicates from ab initio molecular-dynamics simulations: comparison between CaO-Al2O3-SiO2 and SiO2. Phys. Rev. B 64, 224205. Boccaccini, A.R., Erol, M., Stark, W.J., Mohn, D., Hong, Z., Mano, J.F., 2010. Polymer/bioactive glass nanocomposites for biomedical applications: a review. Compos. Sci. Technol. 70 (13), 1764 1776. Brauer, D.S., 2015. Bioactive glasses—structure and properties. Angew. Chem. Int. Ed. Engl. 54 (14), 4160 4181. Brauer, D.S., Karpukhina, N., O’Donnell, M.D., Law, R.V., Hill, R.G., 2010. Fluoridecontaining bioactive glasses: effect of glass design and structure on degradation, pH and apatite formation in simulated body fluid. Acta Biomater. 6 (8), 3275 3282. Brink, M., 1997. The influence of alkali and alkaline earths on the working range for bioactive glasses. J. Biomed. Mater. Res. 36, 109 117. Brown, R.F., Rahaman, M.N., Dwilewicz, A.B., et al., 2009. Effect of borate glass composition on its conversion to hydroxyapatite and on the proliferation of MC3T3-E1cells. J. Biomed. Mater. Res. A 88 (2), 392 400.
Bioactive glasses—structure and applications 471 Burwell, A.K., Litkowski, L.J., Greenspan, D.C., 2009. Calcium sodium phosphosilicate (NovaMins): remineralization potential. Adv. Dent. Res. 21 (1), 35 39. Chen, Q.Z., Li, Y., Jin, L.Y., Quinn, J.M., Komesaroff, P.A., 2010. A new sol-gel process for producing Na(2)O-containing bioactive glass ceramics. Acta Biomater. 6 (10), 4143 4153. Cheng, Y., Xiao, H., Shuguang, C., Tang, B., 2009. Structure and crystallization of B2O3Al2O3-SiO2 glasses. Physica B: Condens. Matter 404, 1230 1234. Cho, J.S., Kang, Y.C., 2009. Synthesis of spherical shape borate-based bioactive glass powders prepared by ultrasonic spray pyrolysis. Ceram. Int. 35, 2103 2109. Dababneh, R.H., Khouri, A.T., Addy, M., 1999. Dentine hypersensitivity—an enigma? A review of terminology, mechanisms, aetiology and management. Br. Dent. J. 187 (11), 606 611. Diba, M., Tapia, F., Boccaccini, A.R., Stro, L.A., 2012. Magnesium-containing bioactive glasses for biomedical applications. Int. J. Appl. Glass Sci. 3, 221 253. Dietzel, A., 1941. Structural chemistry of glasses. Naturwissenschaften 29, 537 547. Dimarzio, E.A., Gibbs, J.H., 1959. Glass temperature of copolymers. J. Polym. Sci. 40 (136), 121 131. Domingues, Z.R., Cortés, M.E., Gomes, T.A., et al., 2004. Bioactive glass as a drug delivery system of tetracycline and tetracycline associated with β-cyclodextrin. Biomaterials 25 (2), 327 333. Dudowicz, J., Freed, K.F., Douglas, J.F., 2005. The glass transition temperature of polymer melts. J. Phys. Chem. B 109 (45), 21285 21292. Elshahat, A.H., 2006. Correction of craniofacial skeleton contour defects using bioactive glass particles. Egypt J. Plast. Reconstr. Surg. 30 (2), 113 119. Farooq, I., Imran, Z., Farooq, U., Leghari, A., Ali, H., 2012. Bioactive glass: a material for the future. World J. Dent. 3 (2), 199 201. Farooq, I., Tylkowski, M., Müller, S., Janicki, T., Brauer, D.S., Hill, R.G., 2013. Influence of sodium content on the properties of bioactive glasses for use in air abrasion. Biomed. Mater. 8 (6), 065008. Farooq, I., Moheet, I.A., AlShwaimi, E., 2015. In vitro dentin tubule occlusion and remineralization competence of various toothpastes. Arch. Oral. Biol. 60 (9), 1246 1253. Farooq, I., Majeed, A., AlShwaimi, E., Almas, K., 2018. Efficacy of a novel fluoride containing bioactive glass based dentifrice in remineralizing artificially induced demineralization in human enamel. Fluoride (in press). Fathi, M.H., Salehi, M.A., Saatchi, A., Mortazavi, V., Moosavi, S.B., 2003. In vitro corrosion behavior of bioceramic, metallic, and bioceramic metallic coated stainless steel dental implants. Dent. Mater. 19 (3), 188 198. Featherstone, J.D., 2000. The science and practice of caries prevention. J. Am. Dent. Assoc. 131 (7), 887 899. Felipe, M.E., Andrade, P.F., Novaes, A.B., Grisi, M.F., Souza, S.L., Taba, M., et al., 2009. Potential of bioactive glass particles of different size ranges to affect bone formation in interproximal periodontal defects in dogs. J. Periodontol. 80 (5), 808 815. Ferraris, M., Rabajoli, P., Paracchini, L., Brossa, F., 1996. Vacuum plasma spray deposition of titanium particle/glass-ceramic matrix biocomposites. J. Am. Ceram. Soc. 79 (6), 1515 1520. Fetner, A.E., Hartigan, M.S., Low, S.B., 1994. Periodontal repair using PerioGlas in nonhuman primates: clinical and histologic observations. Compendium 15 (7), 932 935. Fu, Q., Rahaman, M.N., Bal, B.S., Huang, W., Day, D.E., 2007. Preparation and bioactive characteristics of a porous 13-93 glass, and fabrication into the articulating surface of a proximal tibia. J. Biomed. Mater. Res. A 82, 222 229.
472 Advanced Dental Biomaterials Fu, Q., Rahaman, M.N., Fu, H., Liu, X., 2010a. Silicate, borosilicate, and borate bioactive glass scaffolds with controllable degradation rate for bone tissue engineering applications. J. Biomed. Mater. Res. A 95 (1), 164 171. Fu, Q., Rahaman, M.N., Bal, B.S., Kuroki, K., Brown, R.F., 2010b. In vivo evaluation of 1393 bioactive glass scaffolds with trabecular and oriented microstructures in a subcutaneous rat implantation model. J. Biomed. Mater. Res. A 95 (1), 235 244. Fuchs, M., Gentleman, E., Shahid, S., Hill, R., Brauer, D., 2015. Therapeutic ion-releasing bioactive glass ionomer cements with improved mechanical strength and radiopacity. Front. Mater. 2, 63. Garcia, C., Cere, S., Duran, A., 2004. Bioactive coatings prepared by sol gel on stainless steel 316L. J. Non-Cryst. Solids 348, 218 224. Gerhardt, L.C., Boccaccini, A.R., 2010. Bioactive glass and glass-ceramic scaffolds for bone tissue engineering. Materials (Basel) 3 (7), 3867 3910. Gillam, D., Tang, J., Mordan, N., Newman, H., 2002. The effects of a novel bioglass dentifrice on dentine sensitivity: a scanning electron microscopy investigation. J. Oral Rehabil. 29 (4), 305 313. González, P., Serra, J., Liste, S., Chiussi, S., León, B., Pérez-Amor, M., et al., 2003. New biomorphic SiC ceramics coated with bioactive glass for biomedical applications. Biomaterials 24, 4827 4832. Goodridge, R.D., Woodm, D.J., Ohtsuki, C., Dalgarno, K.W., 2007. Biological evaluation of an apatite mullite glass-ceramic produced via selective laser sintering. Acta Biomater. 3 (2), 221 231. Gorustovich, A.A., Roether, J.A., Boccaccini, A.R., 2009. Effect of bioactive glasses on angiogenesis: a review of in vitro and in vivo evidences. Tissue Eng. Part B Rev. 16 (2), 199 207. Greenspan, D.C., 1999. Bioactive glass: mechanisms of bone bonding. Tandlakartidningen 91, 31 35. Guarino, V., Causa, F., Ambrosio, L., 2007. Bioactive scaffolds for bone and ligament tissue. Expert Rev. Med. Devices 4 (3), 405 418. Gubler, M., Brunner, T.J., Zehnder, M., Waltimo, T., Sener, B., Stark, W.J., 2008. Do bioactive glasses convey a disinfecting mechanism beyond a mere increase in pH? Int. Endod. J. 41 (8), 670 678. Hansbrough, J.F., Morgan, J., Greenleaf, G., Parikh, M., Nolte, C., Wilkins, L., 1994. Evaluation of Graftskin composite grafts on full-thickness wounds on athymic mice. J. Burn Care Rehabil. 15 (4), 346 353. Hassan, U., Farooq, I., Moheet, I.A., AlShwaimi, E., 2017. Cutting efficiency of different dental materials utilized in an air abrasion system. Int. J. Health Sci. (Qassim) 11 (4), 23 27. Heikkilä, J.T., Aho, H.J., Yli-Urpo, A., Happonen, R.P., Aho, A.J., 1995. Bone formation in rabbit cancellous bone defects filled with bioactive glass granules. Acta Orthop. Scand. 66 (5), 463 467. Hench, L.L., 1991. Bioceramics: from concept to clinic. J. Am. Ceram. Soc. 74, 1487 1510. Hench, L.L., 1998. Biomaterials: a forecast for the future. Biomaterials 19, 1419 1423. Hench, L.L., 2006. The story of Bioglasss. J. Mater. Sci. Mater. Med. 17, 967 978. Hench, L.L., 2013. Chronology of bioactive glass development and clinical applications. New J. Glass Ceram. 3 (2), 67 73. Hench, L.L., Andersson, O.H., 1993. An introduction to bioceramics. In: Hench, L.L., Wilson, J. (Eds.), Advanced Series in Ceramics., 1. Scientific Publishing Co Ltd, Singapore, pp. 41 46.
Bioactive glasses—structure and applications 473 Hench, L.L., Jones, J.R., 2015. Bioactive glasses: frontiers and challenges. Front. Bioeng. Biotechnol. 3, 194. Hill, R.G., 1996. An alternative view of degradation of Bioglasss. J. Mater. Sci. Lett. 15, 1122 1125. Hill, R.G., Brauer, D.S., 2011. Predicting the bioactivity of glasses using the network connectivity or split network models. J. Non-Cryst. Solids 357, 3884 3887. Hoppe, A., Güldal, N.S., Boccaccini, A.R., 2011. A review of the biological response to ionic dissolution products from bioactive glasses and glass-ceramics. Biomaterials 32 (11), 2757 2774. Available from: http://www.biomin.co.uk/science/bioactive-glasses/biomintm-vs-novaminr (accessed 02.02.18.). Huang, W., Rahaman, M.N., Day, D.E., Li, Y., 2006a. Mechanisms for converting bioactive silicate, borate, and borosilicate glasses to hydroxyapatite in dilute phosphate solution. Phys. Chem. Glasses 47, 647 658. Huang, W., Day, D.E., Kittiratanapiboon, K., et al., 2006b. Kinetics and mechanisms of the conversion of silicate (45S5), borate, and borosilicate glasses to hydroxyapatite in dilute phosphate solutions. J. Mater. Sci. Mater. Med. 17, 583 596. Hutmacher, D.W., Schantz, J.T., Lam, C.X.F., Tan, K.C., Lim, T.C., 2007. State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective. J. Tissue Eng. Regen. Med. 1 (4), 245 260. Jia, W.T., Zhang, X., Luo, S.H., et al., 2010. Novel borate glass/chitosan composite as a delivery vehicle for teicoplanin in the treatment of chronic osteomyelitis. Acta Biomater. 6 (3), 812 819. Johnson, A.J.W., Herschler, B.A., 2011. A review of the mechanical behavior of CaP and CaP/polymer composites for applications in bone replacement and repair. Acta Biomater. 7 (1), 16 30. Jones, J.R., 2013. Review of bioactive glass: from hench to hybrids. Acta Biomater. 9 (1), 4457 4486. Jones, J.R., Ehrenfried, L.M., Hench, L.L., 2006. Optimising bioactive glass scaffolds for bone tissue engineering. Biomaterials 27 (7), 964 973. Kanzaki, N., Onuma, K., Treboux, G., Tsutsumi, S., Ito, A., 2000. Inhibitory effect of magnesium and zinc on crystallization kinetics of hydroxyapatite (0001) face. J. Phys. Chem. B 104, 4189 4194. Karageorgiou, V., Kaplan, D., 2005. Porosity of 3D biomaterial scaffolds and osteogenesis. Biomaterials 26 (27), 5474 5491. Kaur, G., Pandey, O.P., Singh, K., Homa, D., Scott, B., Pickrell, G., 2014. A review of bioactive glasses: their structure, properties, fabrication and apatite formation. J. Biomed. Mater. Res. A 102, 254 274. Kawanabe, K., Okada, Y., Matsusue, Y., Iida, H., Nakamura, T., 1998. Treatment of osteomyelitis with antibiotic-soaked porous glass ceramic. J. Bone Joint Surg. Br. 80 (3), 527 530. Khalid, M.D., Khurshid, Z., Zafar, M.S., et al., 2017. Bioactive glasses and their applications in dentistry. J. Pak. Dent. Assoc. 26 (1), 32 38. Khvostenko, D., Hilton, T.J., Ferracane, J.L., Mitchell, J.C., Kruzic, J.J., 2016. Bioactive glass fillers reduce bacterial penetration into marginal gaps for composite restorations. Dent. Mater. 32 (1), 73 81. Kobayashi, M., Saito, H., Mase, T., et al., 2010. Polarization of hybridized calcium phosphoaluminosilicates with 45S5-type bioglasses. Biomed. Mater. 5 (2), 25001. Koller, G., Cook, R.J., Thompson, I.D., Watson, T.F., Di-Silvio, L., 2007. Surface modification of titanium implants using bioactive glasses with air abrasion technologies. J. Mater. Sci. Mater. Med. 18 (12), 2291 2296.
474 Advanced Dental Biomaterials Kundu, B., Nandi, S.K., Dasgupta, S., et al., 2011. Macro-to-micro porous special bioactive glass and ceftriaxone sulbactam composite drug delivery system for treatment of chronic osteomyelitis: an investigation through in vitro and in vivo animal trial. J. Mater. Sci. Mater. Med. 22 (3), 705 720. Kwakye-Awuah, B., Williams, C., Kenward, M.A., Radecka, I., 2008. Antimicrobial action and efficiency of silver-loaded zeolite X. J. Appl. Microbiol. 104 (5), 1516 1524. Lai, W., Garino, J., Ducheyne, P., 2002. Silicon excretion from bioactive glass implanted in rabbit bone. Biomaterials 23, 213 217. Liu, X., Xie, Z., Zhang, C., Pan, H., Rahaman, M.N., Zhang, X., et al., 2010. Bioactive borate glass scaffolds: in vitro and in vivo evaluation for use as a drug delivery system in the treatment of bone infection. J. Mater. Sci. Mater. Med. 21 (2), 575 582. Liu, X., Rahaman, M.N., Fu, Q., 2011. Oriented bioactive glass (13-93) scaffolds with controllable pore size by unidirectional freezing of camphene-based suspensions: microstructure and mechanical response. Acta Biomater. 7 (1), 406 416. Lynch, E., Brauer, D.S., Karpukhina, N., Gillam, D.G., Hill, R.G., 2012. Multi-component bioactive glasses of varying fluoride content for treating dentin hypersensitivity. Dent. Mater. 28 (2), 168 178. Macedo, N.L.D., Matuda Fda, S., Macedo, L.G., et al., 2004. Bone defect regeneration with bioactive glass implantation in rats. J. Appl. Oral Sci. 12, 137 143. Mahmood, A., Mneimne, M., Zou, L.F., Hill, R.G., Gillam, D.G., 2014. Abrasive wear of enamel by bioactive glass-based toothpastes. Am. J. Dent. 27 (5), 263 267. Mehta, A.B., Kumari, V., Jose, R., Izadikhah, V., 2014. Remineralization potential of bioactive glass and casein phosphopeptide-amorphous calcium phosphate on initial carious lesion: an in-vitro pH-cycling study. J. Conserv. Dent. 17 (1), 3 7. Mercier, C., Follet-Houttemane, C., Pardini, A., Revel, B., 2011. Influence of P2O5 content on the structure of SiO2-Na2O-CaO-P2O5 bioglasses by 29Si and 31P MAS-NMR. J. Non-Cryst. Solids 357, 3901 3909. Meseguer-Olmo, L., Ros-Nicolás, M.J., Vicente-Ortega, V., et al., 2006. A bioactive sol-gel glass implant for in vivo gentamicin release. Experimental model in rabbit. J. Orthop. Res. 24 (3), 454 460. Mistry, S., Kundu, D., Datta, S., Basu, D., 2011. Comparison of bioactive glass coated and hydroxyapatite coated titanium dental implants in the human jaw bone. Aust. Dent. J. 56 (1), 68 75. Mneimne, M., Hill, R.G., Bushby, A.J., Brauer, D.S., 2011. High phosphate content significantly increases apatite formation of fluoride-containing bioactive glasses. Acta Biomater. 7 (4), 1827 1834. Moimas, L., Biasotto, M., Di-Lenarda, R., Olivo, A., Schmid, C., 2006. Rabbit pilot study on the resorbability of three-dimensional bioactive glass fibre scaffolds. Acta Biomater. 2 (2), 191 199. Najeeb, S., Khurshid, Z., Matinlinna, J.P., et al., 2015. Nanomodified peek dental implants: bioactive composites and surface modification—a review. Int. J. Dent. 2015, 381759. Najeeb, S., Zafar, M.S., Khurshid, Z., et al., 2017. Bisphosphonate releasing dental implant surface coatings and osseointegration: a systematic review. J. Taibah Univ. Med. Sci. 12 (5), 369 375. Nejatain, T., Khurshid, Z., Zafar, M.S., et al., 2017. Dental biocomposites. Biomaterials for Oral and Dental Tissue Engineering. Woodhead Publishing, pp. 65 84. Nerem, R.M., 1991. Cellular engineering. Ann. Biomed. Eng. 19 (5), 529 545. Newby, P.J., El-Gendy, R., Kirkham, J., Yang, X.B., Thompson, I.D., Boccaccini, A.R., 2011. Ag-doped 45S5 Bioglasss-based bone scaffolds by molten salt ion exchange: processing and characterisation. J. Mater. Sci. Mater. Med. 22 (3), 557 569.
Bioactive glasses—structure and applications 475 O’Donnell, M.D., 2011. Predicting bioactive glass properties from the molecular chemical composition: glass transition temperature. Acta Biomater. 7, 2264 2269. O’Donnell, M.D., 2012. Melt-derived bioactive glass. In: Jones, J.R., Clare, A.G. (Eds.), BioGlasses. Wiley, Chichester, UK, pp. 13 28. O’Donnell, M.D., Watts, S.J., Hill, R.G., Law, R.V., 2009. The effect of phosphate content on the bioactivity of soda-lime-phosphosilicate glasses. J. Mater. Sci. Mater. Med. 20 (8), 1611 1618. Oonishi, H., Kushitani, S., Yasukawa, E., Iwaki, H., Hench, L.L., Wilson, J., et al., 1997. Particulate Bioglasss compared with hydroxyapatite as a bone graft substitute. Clin. Orthop. Relat. Res. 334, 316 325. Oonishi, H., Hench, L.L., Wilson, J., et al., 2000. Quantitative comparison of bone growth behavior in granules of Bioglasss, A-W glass-ceramic, and hydroxyapatite. J. Biomed. Mater. Res. 51 (1), 37 46. Paolinelis, G., Banerjee, A., Watson, T.F., 2008. An in vitro investigation of the effect and retention of bioactive glass air-abrasive on sound and carious dentine. J. Dent. 36 (3), 214 218. Peitl, O., Dutra, E., Hench, L.L., 2001. Highly bioactive P2O5 6 Na2O 6 CaO 6 SiO2 glassceramics. J. Non-Cryst. Solids 292, 115 126. Rahaman, M.N., Day, D.E., Sonny, B.B., et al., 2011. Bioactive glass in tissue engineering. Acta Biomater. 7, 2355 2373. Rao, M.C., Ravindranadh, K., 2016. Bioactive glasses for technological and clinical applications. Int. J. Chem. Sci. 14 (3), 1339 1348. Reynolds, E.C., 2008. Calcium phosphate-based remineralization systems: scientific evidence? Aust. Dent. J. 53 (3), 268 273. Roessler, S., Zimmermann, R., Scharnweber, D., Werner, C., Worch, H., 2002. Characterization of oxide layers on Ti6Al4V and titanium by streaming potential and streaming current measurements. Colloids Surf., B: Biointerfaces 26 (4), 387 395. Sen, S., Youngman, R.E., 2003. NMR study of Q-speciation and connectivity in K2O SiO2 glasses with high silica content. J. Non-Cryst. Solids 331, 100 107. Serra, J., González, P., Liste, S., Chiussi, S., León, B., Pérez-Amor, M., et al., 2002. Influence of the non-bridging oxygen groups on the bioactivity of silicate glasses. J. Mater. Sci. Mater. Med. 13, 1221 1225. Serra, J., González, P., Liste, S., Serra, C., Chiussi, S., León, B., et al., 2003. FTIR and XPS studies of bioactive silica based glasses. J. Non-Cryst. Solids 332, 20 27. Srivastava, A.K., Pyare, R., Singh, S.P., 2012. In vitro bioactivity and physical—mechanical properties of Fe2O3 substituted 45S5 bioactive glasses and glass—ceramics. Int. J. Sci. Eng. Res. 3 (2), 1 15. Stanley, H.R., Clark, A.E., Pameijer, C.H., Louw, N.P., 2001. Pulp capping with a modified Bioglasss formula (# A68-modified). Am. J. Dent. 14 (4), 227 232. Stoor, P., Apajalahti, S., Kontio, R., 2017. Regeneration of cystic bone cavities and bone defects with bioactive glass S53P4 in the upper and lower jaws. J. Craniofac. Surg. 28 (5), 1197 1205. Tahriri, M., Bader, R., Yao, W., Dehghani, S., Khoshroo, K., et al., 2017. Bioactive glasses and calcium phosphates. Biomaterials for Oral and Dental Tissue Engineering. Woodhead Publishing, pp. 7 24. Toosi, S., Behravan, J., 2017. An update on bioactive glass in bone tissue engineering. Bioceram. Dev. Appl. 7, e111. Van-Gestel, N.A., Geurts, J., Hulsen, D.J., Van-Rietbergen, B., Hofmann, S., Arts, J.J., 2015. Clinical applications of S53P4 bioactive glass in bone healing and osteomyelitic treatment: a literature review. Biomed. Res. Int. 2015, 684826.
476 Advanced Dental Biomaterials Vichery, C., Nedelec, J.-M., 2016. Bioactive glass nanoparticles: from synthesis to materials design for biomedical applications. Materials (Basel) 9 (4), 288. Wallace, K.E., Hill, R.G., Pembroke, J.T., Brown, C.J., Hatton, P.V., 1999. Influence of sodium oxide content on bioactive glass properties. J. Mater. Sci. Mater. Med. 10 (12), 697 701. Wang, W., Yeung, K.W.K., 2017. Bone grafts and biomaterials substitutes for bone defect repair: a review. Bioact. Mater. 2 (4), 224 247. Wilson, J., Low, S.B., 1992. Bioactive ceramics for periodontal treatment: comparative studies in the patus monkey. J. Appl. Biomater. 3 (2), 123 129. Wilson, J., Clark, A.E., Hall, M., et al., 1993. Tissue response to Bioglasss endosseous ridge maintenance implants. J. Oral. Implantol. 19, 295 302. Xynos, I.D., Hukkanen, M.V., Batten, J.J., Buttery, L.D., Hench, L.L., Polak, J.M., 2000. Bioglasss 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: implications and applications for bone tissue engineering. Calcif. Tissue Int. 67 (4), 321 329. Xynos, I.D., Edgar, A.J., Buttery, L.D., Hench, L.L., Polak, J.M., 2001. Gene-expression profiling of human osteoblasts following treatment with the ionic products of Bioglasss 45S5 dissolution. J. Biomed. Mater. Res. A 55 (2), 151 157. Yang, S.Y., Kwon, J.S., Kim, K.N., Kim, K.M., 2016. Enamel surface with pit and fissure sealant containing 45S5 bioactive glass. J. Dent. Res. 95 (5), 550 557. Yang, X., Zhang, L., Chen, X., Sun, X., Yang, G., et al., 2012. Incorporation of B2O3 in CaO-SiO2-P2O5 bioactive glass system for improving strength of low-temperature co-fired porous glass ceramics. J. Non-Cryst. Solids 358, 1171 1179. Yao, A., Wang, D., Huang, W., Fu, Q., Rahaman, M.N., Day, D.E., 2007. In vitro bioactive characteristics of borate-based glasses with controllable degradation behavior. J. Am. Ceram. Soc. 90, 303 306. Yunos, D.M., Bretcanu, O., Boccaccini, A.R., 2008. Polymer-bioceramic composites for tissue engineering scaffolds. J. Mater. Sci. 43 (13), 4433. Zadpoor, A.A., 2014. Relationship between in vitro apatite-forming ability measured using simulated body fluid and in vivo bioactivity of biomaterials. Mater. Sci. Eng. C 35, 134 143. Zhang, X., Jia, W., Gu, Y., et al., 2010. Teicoplanin-loaded borate bioactive glass implants for treating chronic bone infection in a rabbit tibia osteomyelitis model. Biomaterials 31 (22), 5865 5874. Zhao, D., Moritz, N., Vedel, E., Hupa, L., Aro, H.T., 2008. Mechanical verification of softtissue attachment on bioactive glasses and titanium implants. Acta Biomater. 4 (4), 1118 1122. Further reading Hench, L.L., 1981. Bioceramics. J. Am. Ceram. Soc. 81 (7), 1705 1728. Owens, G.J., Singh, R.K., Forouton, F., et al., 2016. Sol gel based materials for biomedical applications. Prog. Mater. Sci. 77, 1 79.
Nanotechnology and nanomaterials in dentistry 18 Muhammad S. Zafar1,2, Ahmad A. Alnazzawi3, Mothanna Alrahabi1, Muhammad A. Fareed4, Shariq Najeeb5 and Zohaib Khurshid6 1 Department of Restorative Dentistry, College of Dentistry, Taibah University, Medina, Saudi Arabia, 2Department of Dental Materials, Islamic International Dental College, Riphah International University, Islamabad, Pakistan, 3Department of Substitutive Dental Sciences, College of Dentistry, Taibah University, Medina, Saudi Arabia, 4Adult Restorative Dentistry, Dental Biomaterials and Prosthodontics Oman Dental College, Muscat, Sultanate of Oman, 5National Center for Proteomics, University of Karachi, Pakistan, 6Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia Chapter Outline 18.1 18.2 18.3 18.4 Introduction 477 Natural biomaterials and nanoscience 478 General properties of nanomaterials 480 Dental applications of nanobiomaterials 481 18.4.1 18.4.2 18.4.3 18.4.4 18.4.5 18.4.6 18.4.7 18.4.8 Nanobiomaterials for preventive dentistry 482 Nanomaterials for periodontics 483 Nanomaterials for dental implants 484 Restorative nanobiomaterials 486 Endodontic nanobiomaterials 488 Nanomaterials and endodontic regeneration 490 Nanomaterials and tissue engineering 490 Electrospun nanomaterials 491 18.5 Potential of nanomaterials 18.6 Conclusive remarks 494 References 494 18.1 492 Introduction In the last few decades, clinical applications of nanotechnology and nanomaterials have expanded in biomedical and dental care. Plenty of research has been conducted to investigate the excellent properties of various nanoscale biomaterials including carbon-based nanoparticles; nanoceramics (Najeeb et al., 2016a; Elgendy and Abo Shady, 2015; Vallet-Regı́ and Arcos, 2008), graphene oxide (Sun et al., 2008; Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00018-9 Copyright © 2019 Elsevier Ltd. All rights reserved.
478 Advanced Dental Biomaterials Gahlot et al., 2014; Rodrı́guez-Lozano et al., 2014), carbon nanotubes (Silva et al., 2014; Kang et al., 2009; Kostarelos et al., 2009), nanodiamonds (Gismondi et al., 2015; Passeri et al., 2015; Setyawati et al., 2015; Huynh et al., 2013; Najeeb et al., 2016d), and nanofibers (Zafar et al., 2015a,b, 2016; Kim et al., 2014a,b; Zafar and Al-Samadani, 2014) for potential biodental applications. The word “nano” is a Greek word that means dwarf. In the metric system, one billionth of 1 m is called a nanometer (nm) (1 nm 5 1029 m) (Sheeparamatti et al., 2007). A nanomaterial has at least one dimension in the range of 1 100 nm and demonstrates size dependent characteristic properties (Gouma, 2009). Nanotechnology deals with the engineering of nanomaterials (scale range of 1 100 nm) and miniaturized devices based on nanomaterials. Nanodentistry is the use of nanomaterials and nanotechnology applications for diagnosis, prevention, and treatment of oral and dental diseases (Hieber and Müller, 2012). The concept of nanotechnology was introduced by Richard Feynman while presenting his conceptual theory “There’s Plenty of Room at the Bottom” at the California Institute of Technology that was later published in “Engineering and Science” (Feynman, 1960). However, the actual term “Nanotechnology or Nanosciences” was described by Prof. Norio Taniguchi (Tokyo Science University) for the first time in 1974 (Taniguchi, 1974). In the early 1980s, physicists achieved major developments in the field of scanning microscopy enabling image analysis at an atomic/molecular level, advancing the science of nanotechnology. Nanotechnology engages the fabrication and characterization of materials at molecular levels and purposeful engineering on a nanoscale ( . 100 nm) (Zhang et al., 2005). Natural and synthetic materials at various scales are compared in Fig. 18.1. The physical, chemical, mechanical, and biological properties of materials at an individual molecular level may differ compared to their bulk properties at the nanoscale (Lieber, 1998). Nanomaterials characterization means the use of different techniques and methods to probe materials structures and properties. Regarding nanomaterials characterization, a variety of techniques are used to assess the structure and properties of existing and developing nanomaterials (Fig. 18.2). Nanobiomaterials have been explored for all types of biomaterials (metals, polymers, ceramics, and composites) for clinical applications. Due to favorable features and promising outcomes, nanomaterials have attracted researchers and biomaterials scientist looking to provide benefits for biomedical and dental applications (Moszner and Klapdohr, 2004). Therefore, the aim of this chapter is to discuss the current and potential dental applications of nanobiomaterials, recent innovations, and future expectations. 18.2 Natural biomaterials and nanoscience Natural biomaterials such as bone, teeth, shells, and wood are composed of nanoscale components (Hussain et al., 2006), having great performance and structural/ functional relationships. Nature has taught us a lot of lessons as the majority of natural biomaterials are synthesized in the biological environment and conditions (Broderick et al., 2007). In addition, a range of hierarchical levels (ranging from
Nanotechnology and nanomaterials in dentistry 479 Figure 18.1 Comparison of natural and synthetic materials at various metric scales. the macro- to microscale) and nanometer-sized inorganic materials are implanted into the soft organic matrix (Mano and Reis, 2005). The formation of nanomaterials in nature takes place in ambient conditions of the surrounding environment
480 Advanced Dental Biomaterials Nanomaterials' characterization Morphology Surface Topography Particle size Porosity Surface area Scanning tunnel microscope (STM) Atomic force microscope (AFM) Transmission electron microscope (TEM) Scanning electron microscope (SEM) Dynamic light scattering (DLS) Gas adsorption Scanning probe microscope Small angle X-ray scattering (SAXS) Crystallography Crystallinity and defects Grain size X-ray diffraction (XRD) Electron crystallography Wide-angle X-ray scattering Scanning electron microscope (SEM) Chemical structure Elements Molecules Chemistry Chemical bond Concentration Energy-dispersive X-ray spectroscopy (EDX) Neutron diffraction Mass Spectroscopy Infra-red spectroscopy (FTIR) UV–Vis spectroscopy Scanning electron microscope (SEM) Thermal properties Thermal conductivity Melting point Thermal stability Glass transition Thermogravimetric analysis (TGA) Differential scanning calorimetry (DSC) Thermomechanical analysis (TMA) Dynamic mechanical analysis (DMA) Mechanical properties Strength Hardness Modulus Wear Viscoelastic behavior AFM Nanoindentation Mechanical tester Dynamic mechanical analysis (DMA) Figure 18.2 Commonly used techniques for the characterization of nanomaterials. (Broderick et al., 2007); for instance, the formation of enamel hydroxyapatite (HA) is a matter of simple dissolution precipitation during amelogenesis (LeGeros, 1991). In contrast, the formation of the same product in the laboratory requires calcium fluoride and high temperature (900 C 1100 C) (Emsley, 1976). Moreover, mineralized body tissues are primarily nanocomposites composed of inorganic/ organic components in order to tailor the tissue properties from one tissue to the other (Low et al., 2008). The inorganic component HA comprises the major part of enamel, dentin, cementum, and bone. HA crystalizes in a hexagonal shape (Jones, 2001) with its chemistry known for more than 85 years (Naray-Szabo, 1930). However, the morphology and structure varies among biological tissues (Summit et al., 2000; Nanci, 2008; Ten Cate, 1994). For instance, HA crystals in tooth enamel are remarkably larger (45 90 nm 3 25 39 nm 3 2 3 nm) than in dentin (3.5 nm 3 20 nm 3 20 nm) (Summit et al., 2000). 18.3 General properties of nanomaterials Nanomaterials are expected to exhibit novel and significantly better physical, chemical, and biological properties (Kelsall et al., 2005). The properties of nanomaterials may vary based on multiple factors such as the materials’ type (metal, ceramic, polymer, or composites) and morphological structure. However, certain properties and materials characteristics are expected from nanomaterials. The detailed discussion of these
Nanotechnology and nanomaterials in dentistry 481 Table 18.1 Classification of nanomaterials. Classification A—Based on morphology B—Based on materials chemistry Nanoclusters Nanofibers Nanowire/tubes Nanoparticles Carbon-based nanomaterials Polymer nanomaterials Metallic nanomaterials Nanoceramics Nanocomposites C—Based on dimensions Zero-dimensional One-dimensional Two-dimensional Three-dimensional Examples Electrospun polymers/silk Carbon-based nanomaterials Silica, HA, gold, silver Fullerenes nanotubes Electrospun silk nanofibers Nanogold and nanosilver Silica, titanium oxide, and HA nanoparticles Ceramic/polymer nanocomposites silica/resin nanocomposites Nanoclusters Nanowires and nanotubes Nanolayer/films coatings Tissue engineering scaffolds polymers and/or composites HA, Hydroxyapatite. properties is beyond the scope of this book; but a few main features of nanoscale materials are enumerated here: G G G G G G At nanoscale, greater surface area the surface functionalities. High surface area facilitates better mechanical interlocking of nanoparticles to the polymer matrix (Arcı́s et al., 2002). Inorganic ceramic nanoparticles are brittle and hard and can reinforce to provide superior mechanical properties (Mota et al., 2006). The areas of stress concentration are reduced, resulting in improved resistance to the crack propagation and higher fatigue strength (Turssi et al., 2005). Optical properties (translucency) and surface finish are improved while using nano-sized fillers (Mota et al., 2006). The biodegradability and biodegradation rates can be better controlled compared to conventional composite materials (Ray and Okamoto, 2003; Mohanty et al., 2003). The various types of nanomaterials can be categorized based on their chemistry, morphology, and dimensional existence (Table 18.1). 18.4 Dental applications of nanobiomaterials Nanobiomaterials have been extensively explored for various dental applications (Najeeb et al., 2016; Zafar et al., 2017; Khurshid et al., 2015), covering almost all dental specialties (Fig. 18.3). The following describes the key nanomaterials being used in various disciplines of dentistry.
482 Advanced Dental Biomaterials Figure 18.3 Applications of nanomaterials in various disciplines of dentistry. 18.4.1 Nanobiomaterials for preventive dentistry “Prevention is better than cure.” Teeth perform in the dynamic environment of the oral cavity where the prevention of tooth decay is very challenging. Due to all the advancements and the better understanding of oral diseases, the role of preventive dentistry is still vital and cannot be denied. Nanomaterials are used for preventive dental care, for the management of biofilm at the tooth surface using nanoapatites, and for the demineralization of early stage submicron-sized enamel lesions (Hannig and Hannig, 2010). 18.4.1.1 Preventive nanocomposites surface coatings The nanocomposites surface coatings are prepared by the incorporation of inorganic nanoparticles in fluoropolymer matrix (Hannig et al., 2007). These materials play beneficial roles by detaching the pellicle and bacterial plaque under the effect of physiological activities in the oral cavity and decreasing the surface energy to inhibit biofilm attachment to the tooth structure. Key benefits include the following (Hannig et al., 2007): G G G G Modification of tooth surface. Easy to clean when used as coating or fissure sealants. Wear resistant. Biocompatible.
Nanotechnology and nanomaterials in dentistry G G 483 Low surface energy that helps in the detachment of bacteria and biofilm. Can be beneficial in compromised patients having xerostomia, poor oral hygiene, and high caries rate. 18.4.1.2 Casein phosphopeptide and amorphous calcium phosphate nanocomplexes Casein phosphopeptide (CPP) and amorphous calcium phosphate (ACP) are thought to accelerate the remineralization of the enamel surface (Madhavan et al., 2012; Rizvi et al., 2016). CPP ACP nanomaterials (B2.12 nm) are used for the prevention of enamel caries, remineralization (Reynolds, 2008) and for the management of oral biofilm (Cross et al., 2007). Nanocomplexes in CPP ACP are significantly smaller in size than microorganisms, therefore establishing an effective interaction with oral microorganisms (Hannig and Hannig, 2010) and binding to the bacterial cell surface to interfere with their adherence with the tooth (Cross et al., 2007). This sequence delays the formation of biofilm (Rahiotis et al., 2008), inhibits demineralization, and promotes remineralization of initial carious lesions (Reynolds, 2008). Furthermore, CPP ACP nanomaterials are used in sugar-free gums due to their proven role in preventing proximal caries (Morgan et al., 2008). 18.4.1.3 Nanohydroxyl apatite toothpaste Carbonate hydroxyl apatite nanoparticles (20 100 nm) were used to repair micronsized carious lesions (Roveri et al., 2009) as a result of hydroxyl apatite deposition in to the demineralized tissues. These toothpastes contain spheroid or needle-like hydroxyl apatite nanoparticles as an active constituent (Lv et al., 2007). Key benefits include ease of application as the medium of delivery is toothpaste and deposition of apatite nanoparticles in the defect. The remineralization of decaying tissues was better in comparison to sodium fluoride-containing formulations (Lv et al., 2007). 18.4.2 Nanomaterials for periodontics Uncontrolled periodontal lesions cause bone resorption and tooth loss (Pihlstrom et al., 2005). Nanomaterials used for the cure of periodontal diseases have shown promising outcomes (Piñón-Segundo et al., 2005). A delivery system based on nanoparticles loaded with triclosan was developed for the cure of periodontal diseases. Poly(D,L-lactide-co-glycolide), poly(D,L-lactide), and cellulose phthalate were used to prepare nanoparticles (,500 nm). Various quantities of triclosan as an active nanomaterial ingredient were proved to reduce the inflammation in animal studies (Piñón-Segundo et al., 2005). A site-specific delivery system such as Arestin is used for clinical applications where microspheres containing tetracycline are introduced directly into the periodontal pocket (Paquette et al., 2004). In recent decades, resorbable guided tissue regeneration (GTR) membranes to stimulate the regeneration of periodontal bone have been explored. The GTR membrane acts as a barrier between the bone and the gingival epithelium
484 Advanced Dental Biomaterials (Bottino et al., 2012). The limitations of using biodegradable GTR materials are poor mechanical properties and the possibility of adverse tissue reactions (Bottino et al., 2012). To reinforce the mechanical properties, poly-L-lactide fibrous and membranous scaffolds loaded with octadecylamine-functionalized diamond nanoparticles showed better strength (Zhang et al., 2011, 2012). Growth factors such as bone morphogenetic protein-2 and fibroblast growth factor carried by nanodiamonds in injectable form were observed to stimulate osteoblast differentiation in vitro (Moore et al., 2013). It is suggested that nanoparticles reinforced membranes can overcome the limitations of existing GTR membranes in addition to delivering bioactive growth factors to the periodontal tissues. Nanodiamond-loaded poly(lactide-co-glycolide) (Parizek et al., 2012; Brady et al., 2015), poly(L-lactide)co-(ε-caprolactone), and poly(LLA-co-CL) scaffolds (Xing et al., 2013) have been electrospun and showed promising results for proliferation and growth of human osteoblast-like MG-63 and mesenchymal cells. Furthermore, the nanodiamondloaded poly(LLA-co-CL) scaffolds were used for periodontal tissues in vivo and showed significantly greater bone growth compared to unmodified scaffolds (Xing et al., 2013; Suliman et al., 2015). Recently, the natural polymeric material chitosan was explored for potential applications in dentistry and tissue regeneration (Qasim et al., 2018; Husain et al., 2017). Chitosan is a natural bioresorbable material used in the form of nanoparticles or films which can be used to deliver medicaments (chlorhexidine, metronidazole, and nystatin) to periodontium (Qasim et al., 2017; Al-Bayaty et al., 2013; Pichayakorn and Boonme, 2013). The drug release is improved, corresponding to the higher surface area and reactivity of nanoparticles (Khurshid et al., 2015; Paul and Robeson, 2008). There is a potential of developing nanodiamond scaffolds in the form of membranes or injectable hydrogels to deliver regenerative growth factors to the dental tissues. It is expected to develop more materials for periodontal applications including periodontal tissue regeneration. 18.4.3 Nanomaterials for dental implants The concept of using dental implants is not new; however, the modern dental implants became popular after Brånemark’s novel work on titanium (Brånemark et al., 1969). Although dental implants have a high success rate (95%), failure may be caused by infection, accelerated bone loss, or lack of osseointegration (Pye et al., 2009). The metallic dental implants have been used successfully for last four decades; however, there are certain shortcomings related to the osseointegration and the mismatch of the mechanical properties of biomaterials and bone (Tomsia et al., 2011). For example, the Young’s modulus is almost five times higher than that of cortical bone (Oh et al., 2002), which may cause areas of stress concentration around the implant biomaterials resulting in bone resorption and failure. The implant’s surface is required to be modified in order to improve the surface biocompatibility, hydrophilicity, and enhanced osseointegration with alveolar bone (Le Guéhennec et al., 2007). Failures of dental implants are mainly associated with peri-implantitis or failure of the bioactive surface coating (Esposito et al., 1998; Esposito, 1998). Plenty of research is being performed on surface modifications of
Nanotechnology and nanomaterials in dentistry 485 dental implants to overcome these shortcomings (Najeeb et al., 2017; Kulkarni Aranya et al., 2017; Matinlinna et al., 2013; Najeeb et al., 2016b). Nanomaterials coatings improved the mechanical as well as biological properties of titanium implants (Valiev et al., 2007). There are high expectations from nanomaterials research to improve the properties and to manufacture new generations of implant materials. Some modifications of the implant surface involving nanomaterials are described here. 18.4.3.1 Nanozirconia and alumina materials These materials were developed using tetragonal zirconia (ZrO2) nanoparticles and alumina matrix and exhibited improved mechanical properties (Pecharroman et al., 2003; Deville et al., 2003). 18.4.3.2 Titanium and silica nanocomposites Silica is well known for its bioactivity for calcified tissue as it promotes the formation of HA (Cao and Hench, 1996). The addition of silica on the implant surface enhanced the bioactivity by forming Si OH groups on the surface. Titanium was added to improve mechanical properties for load-bearing applications (Jurczyk et al., 2013). These nanocomposites have better hardness and facilitate the tissue and bone growth in the micro- and macro-sized pores present in their surface (Jurczyk et al., 2013). 18.4.3.3 Calcium phosphorous nanoparticles Titanium implant surfaces can be modified using calcium phosphorous (Ca P) nanoparticles. Ca P nanoparticles (B20 nm) can be deposited on the implant surface by sol gel technique (Mendonça et al., 2008). There was a rise in bone growth in Ca P-modified surfaces compared to the commercially pure titanium and titanium alloys (Mendes et al., 2007, 2008). 18.4.3.4 Hydroxyapatite nanocrystals HA nanocrystals/particles have been used for the surface modification of dental implants. HA nanocrystals powder can be prepared using the wet chemical process, that can be used to produce either microcrystalline HA (sintering at B1100 C for 60 minutes) or nanocrystalline HA (hydrothermal treatment at 200 C for 24 hours). HA powder deposited on to the implant surface at room temperature resulted in better adhesion of osteoblasts and more calcium deposition in the case of nanocrystalline HA coating compared to traditional HA coating (Sato et al., 2006). 18.4.3.5 Nanodiamond coatings Recently, silanization and immobilization techniques have been used to coat dental implant surface with nanodiamonds (Gonçalves et al., 2014). Nanodiamond coatings have the potential to deliver drugs and growth factors and overcome the
486 Advanced Dental Biomaterials shortcomings of surface modifications (Xu et al., 2014). Nanodiamond coatings are capable of improving the mechanical and biological properties (Khurshid et al., 2015; Najeeb et al., 2015, 2016c; Sheikh et al., 2015); therefore surface nanocoatings can improve the materials tissue interface and provide better stability. 18.4.4 Restorative nanobiomaterials 18.4.4.1 Nanocomposites Resin-based dental composites are the most popular dental restorative materials and are mainly composed of resin polymeric matrix, inorganic fillers particles, and silane coupling agents. The inorganic (quartz, glass, or silica) filler particles (B60 80 vol.%) are added to reinforce the mechanical properties. Considering fillers as a major component, the resin composites are often classified as microfilled, hybrid, and nanocomposites on the basis of filler size and morphology (Sakaguchi and Powers, 2012). In addition, it is well accepted that the properties of resin composites (physical and mechanical) are modified by altering size, proportion, and distribution of filler particles (Terry, 2004). Since the discovery of resin composites, a range of modifications have been introduced to the conventional materials in order to improve the materials’ properties (Buonocore, 1955). The nanocomposites were developed to overcome the shortcomings of conventional resin composites and to gain better performance by incorporating various nanostructures (nanoparticles, nanofibers, and nanoclusters). However, for a true nanocomposites material, the size of all particles must be B1 100 nm. Alternatively, nanohybrid composites contain nanoparticles as well as larger filler particles (0.4 5.0 μm) in the resin matrix (Sakaguchi and Powers, 2012). Incorporation of nanostructures resulted in significantly improved properties of dental composites (Sakaguchi and Powers, 2012) due to the following: G G G G G G G G Larger surface area and availability of more sites for bonding. Nanoparticles size is closer to polymer molecules and results in molecular scale interaction with polymer resin matrix. As particle size is smaller than visible light wavelength (400 800 nm) and responsible for higher translucency and aesthetic. Better smoothness and gloss finish of restorations. Mechanical properties are as good as of microhybrid composites. The values (Mitra et al., 2003) of compressive strength (460 MPa), flexural strength (180 MPa), and diametral tensile strength (81 MPa) are higher than any other types of resin composites and glass-ionomers cement. Excellent wear resistance; as good as that of natural human enamel in 3 and 5 years clinical studies. Nanozirconia particles (5 7 nm) can bring radiopacity without affecting the other properties. The fillers particles size in conventional restorative macroscopic materials is significantly greater (0.04 7 μm) compared to the HA crystal present in tooth enamel (B1 10 nm). This kind of mismatch of properties has an adverse effect on the
Nanotechnology and nanomaterials in dentistry 487 bonding of tooth and synthetic materials (Musselman, 2003). This disparity has remarkably been overcome using nanoparticles in the case of nanocomposites (nanofillers), and therefore having a better capability of establishing a smooth transition and native interface with tooth tissues (Terry, 2004). Conclusively, the incorporation of nanofillers (in the case of nanocomposites and hybrids) improved the physical, mechanical, and optical properties of the resin composites. 18.4.4.2 Nanoglass ionomers (nanoionomers) The conventional glass-ionomer cement (GIC) was introduced by Wilson and Kent (1971) in the 1970s. Since then GIC has emerged as promising dental restorative material. Although there are not many differences, there are a variety of GICs available commercially. Conventionally, GICs are available in powder/liquid form to start the reaction upon mixing. The main component of powder is fluoroaluminosilicate glass with added calcium, strontium, and lanthanum ions. The liquid copolymer (acrylic acid/itaconic acid or maleic acid) is used to start the reaction (Sakaguchi and Powers, 2012). GIC gained popularity due to its unique features such as fluoride release, bonding to the dental tissues, and excellent biocompatibility (Ten Cate and van Duinen, 1995). The shortcomings of GIC included poor mechanical properties (Mount, 1998), aesthetics, longer setting reaction (Mount, 1998; Tyas, 2006), and poor adhesive strength (Mickenautsch et al., 2012). To overcome the shortcomings, research is actively being conducted to modify the properties, such as addition of HA and fluorapatite (Moshaverinia et al., 2008), cellulose fibers (Silva et al., 2013), and nanotechnologies (Jevnikar et al., 2012). Recently, nanotechnology modified the resin-modified glass ionomers (RMGIC) using nanoparticles (nanomers) and nanoclusters. These nanoionomers (such as Ketac Nano) have been used in clinical applications for more than a decade. The addition of nanoparticles improved the aesthetic properties and surface finish without influencing the fluoride releasing properties (Sakaguchi and Powers, 2012). The Knoop hardness of Ketac Nano (B39 KHN) was noticeably lower compared to RMGIC, Vitremer (B70 KHN) (Raggio et al., 2009). In addition, the Knoop hardness of Ketac Nano (B48 KHN) was also lower than the restoration hardness specified by the American Dental Association; therefore it is not recommended for stress bearing areas (Anusavice and Phillips, 2003). It is recommended to use this material for lower stress bearing areas including class I, III, and V restorations or beneath resin composite and in deciduous teeth. Another nanomaterial, Equia system, contains inorganic silica nanofiller particles (B40 nm) homogeneously dispersed into the matrix. The addition of silica nanoparticles (15 wt.%) improveed the wear resistance alongside controlling the initial setting time. The improved wear resistance enables these materials to maintain the surface finish for a prolonged time compared to the original composition. In terms of physical properties, optical properties (translucency and hue) and aesthetics were improved significantly compared to conventional GIC (Basso, 2011). Considering the promising results of adding inorganic nanoparticles (HA, silica,
488 Advanced Dental Biomaterials and fluorapatite) to GIC, it has been a cutting edge research area for scientists for the last few years. Crystalline HA is the main inorganic component of mineralized oral tissues and there are no biocompatibility issues, and synthetic HA is used for biological applications. The nanoscale facilitates additional benefits, for instance, RMGICs’ bond strength to the tooth tissues (B0.75 MPa) was improved (1.02 MPa) by replacing conventional particles with micro-HA (5 10 μm) and further reinforced in the case of nano-HA (1.91 MPa). The higher surface area in the case of HA nanoparticles results in improving the bond strength possibility of better surface finish and bioactivity (Lee et al., 2010). In a similar study, nano-sized HA and fluorapatite synthesized sol gel method were added to modify Fuji II GIC. Such modifications resulted in the improvement of mechanical strengths (compressive, diametral tensile, and biaxial flexural) of GIC (Moshaverinia et al., 2008). In addition, materials such as yttria-stabilized zirconia (Gu et al., 2005), alumina (Jevnikar et al., 2012; Khademolhosseini et al., 2012), and alumina/titania (Khademolhosseini et al., 2012) have been explored to modify existing GICs. On the basis of ongoing research on nano-GIC, further improvement in existing GIC materials may be expected soon. 18.4.5 Endodontic nanobiomaterials Nanomaterials have been explored for various endodontic applications including sealants, irrigators, obturation materials, and endodontic regeneration. This section describes various materials that have been introduced for current and potential endodontic applications. 18.4.5.1 Nanoparticles-based endodontic sealer Biological nanoceramics comprises bioactive components, such as calcium hydroxide, calcium phosphate, calcium silicate, and zirconia, and thickening agent were used to develop EndoSequence BC sealer. The addition of bioactive nanoparticles aimed to improve the materials’ manipulation and physical properties. The EndoSequence BC sealer uses moisture available in the apical area and hydrate to form nano-HA and calcium silicate. Therefore the reaction and setting time may be influenced by canal moisture; for instance, lack of moisture (overly dried canals) results in prolonged setting time (Koch and Brave, 2009). In terms of the material’s properties, EndoSequence BC sealer has a number of benefits, for example, it is available as a premixed injectable paste giving the benefits of a homogeneous mix, being convenient to apply, and saving time (Koch and Brave, 2009; Zoufan et al., 2011). In addition, nano-sized particles facilitate extrusion of materials from capillary needles and adopt to surface irregularities. The setting reaction does not require absolute drying of root canals as the residual moisture ignites the setting reaction to set it hard and formation of bioactive HA (Koch and Brave, 2009; Zoufan et al., 2011). It forms bonding tooth structure and obturation material without any cytotoxicity or biocompatibility issues. It shows excellent dimensional stability and
Nanotechnology and nanomaterials in dentistry 489 antimicrobial properties due to an alkaline pH (12.8). Setting time is B3 4 hours and it provides excellent sealing (Koch and Brave, 2009; Zoufan et al., 2011). 18.4.5.2 GuttaFlow sealer GuttaFlow sealer is a silicon-based material with added silver nanoparticles and gutta percha dust. For convenient application, it is supplied as single-dose capsule to be mixed and injected in to the root canals (Zoufan et al., 2011). Key properties are as follows: G G G G G Good compatibility and limited toxicity to tissue are acceptable. Excellent sealing of the root canal and better resistance to bacterial penetration. No heating required hence no shrinkage and improved sealing capability. Working time is 15 minutes and setting time is 30 minutes. Dimensionally stable and insoluble after setting. 18.4.5.3 Antibacterial nanoparticle modified endodontic sealer Considering the magnificent benefits of antibacterial activity for endodontic sealers, conventional sealers (AH Plus, GuttaFlow, and Epiphany) were modified using quaternary ammonium polyethyleneimine (QPEI) antibacterial nanoparticles (Abramovitz et al., 2012). The addition of QPEI nanoparticles to composites enabled antibacterial activity for a longer period of time without affecting the mechanical properties (Beyth et al., 2008). Itzhak et al. added QPEI nanoparticles (0 2 wt.%) to various endodontic sealers to introduce bactericidal properties (Abramovitz et al., 2012). This study reported the following features while adding antibacterial nanoparticles (Abramovitz et al., 2012): G G G G Excellent antibacterial, that is expected to last longer when nanoparticles are encapsulated in the matrix. QPEI nanoparticles are very stable, prepolymerized, and no by-product or unwanted products are released into the surrounding biological environment. Addition of QPEI nanoparticles (1 wt.%) does not cause any cytotoxicity whereas a higher concentration (2 wt.%) is slightly cytotoxic. No effect on the biocompatibility of modified sealers. 18.4.5.4 Nanohydroxyapatite gutta percha Nano-HA-modified gutta percha was introduced in recent years. In addition to nano-HA, the sealer comprises a powder component, bismuth oxide (opacifier), hexamethylenetetramine (activator), and liquid component (bisphenol-A-diglycidylether). Wet chemistry route with calcium hydroxide and phosphoric acid precursors were used to fabricate nano-HA in the range of B40 60 nm (Farea et al., 2010). These materials showed excellent sealing abilities comparable to AH 26 and Epiphany sealers. There were no unwanted effects during the root canal obturation (Abdo et al., 2012). Recently, diamond nanoparticle reinforced gutta percha has shown better mechanical and antibacterial properties (Lee et al., 2015b).
490 Advanced Dental Biomaterials 18.4.5.5 Bioactive nanoscale glass for root canal disinfection The bioactive glass nanoparticles have demonstrated antibacterial activity. For this purpose, the sol gel route was used to fabricate glass nanoparticles (B20 90 nm) (Mortazavi et al., 2010). Bioactive glasses containing SiO2 CaO P2O5 (such as 58S, 63S, and 72S) have been assessed for the antimicrobial potentials. At the optimal concentration (100 mg/mL) of 58S and 63S for antibacterial affects, it effectively showed control of the bacterial growth on agar plates. The antibacterial activity is further enhanced due to the alkaline nature of nanoglasses. The solution pH rises from B7.3 to 8.8 in a few hours and may become more alkaline (B9.0) in 5 days (Mortazavi et al., 2010). Therefore these nanomaterials release alkaline species in the biological environment and further enhance the antibacterial activity (Waltimo et al., 2007). Bioceramics have proven benefits of bioactivity and excellent biocompatibility; in addition, the release of antibacterial alkaline species highlighted the potential of these materials for root canal disinfection applications (Mortazavi et al., 2010). 18.4.6 Nanomaterials and endodontic regeneration The conventional endodontic treatment involves the extirpation of residual pulp tissues and infected dentin debris using mechanical instrumentation and antimicrobial irrigants (Rizvi et al., 2014; Madarati et al., 2017). Following cleaning and shaping, root canals are filled using gutta purcha (Friedman et al., 1975). However, conventional endodontics has a number of limitations such as loss of vitality, poor mechanical properties, and compromised disinfection (Fouad et al., 2008; Saunders and Saunders, 1994; Siqueira, 2001; AlRahabi, 2017). The endodontic regenerative approaches have the benefit of maintaining pulp vitality and regenerating dental tissues (Horst et al., 2012; Zafar et al., 2015a,b). For the pulp tissue engineering, purified dental pulp stem cells were grown in sheets on scaffolds. The beneficial role of nanomaterials for tissue engineering scaffolds is well accepted because of high surface area and surface energy. For instance, nanofibrous scaffolds made of biodegradable organic matrix (such as collagen and fibronectin) (Venugopal and Ramakrishna, 2005; Fukuda et al., 2006) and hydrogels were explored for regeneration of pulp and showed the formation of nanofibrous mesh supporting cellular growth (Galler et al., 2008). Collagen is the most numerous fibrous protein originating from bone and dentin (Wiesmann et al., 2004). Injectable collagen (type I) scaffolds loaded with exfoliated dental stem cells lead to growth of functional odontoblasts and pulp tissues (Demarco et al., 2011). These findings are suggestive that type I collagen scaffolds encourages odontogenic differentiation and mineralization (Kim et al., 2009; Mizuno et al., 2003). 18.4.7 Nanomaterials and tissue engineering Tissue regeneration is an interdisciplinary approach involving multiple scientific disciplines including biology and engineering (Galler et al., 2010). It was
Nanotechnology and nanomaterials in dentistry 491 considered as a part of biomaterials initially. However, it later developed as an individual subject on its own because of promising and remarkable development. There are a variety of definitions available for tissue engineering in the literature. Simply put, “Tissue engineering applies the principles of biology and engineering to the development of functional substitutes for damaged tissue” (Duncan and Sprehe, 2008; Langer and Vacanti, 1993). For any tissue engineering application, there are three main components: cells (mainly stem cells capable of differentiating into various cell lines), bioactive molecular signals (to assist the cellular differentiation and tissue regeneration), and a material scaffold to expand and carry cellular components (Galler et al., 2010; Hargreaves, 2010). Nanomaterials have been explored for the fabrication of tissue engineering scaffolds. To construct a scaffold of a suitable material and to mimic extracellular matrix to support cell growth, differentiation and functionality are the foremost challenges for researchers. There are different techniques used for constructing nanofibrous polymer scaffolds. Electrospinning is a potential method of fabricating nanomaterials for such applications and is commonly used (Gouma, 2009). 18.4.8 Electrospun nanomaterials Electrospinning technique is used for fabricating nanofibers (Fig. 18.4) (from a few nanometers in diameter to tens of micron); additives such as biomolecules or medicaments can also be incorporated to get the desired properties in the final materials Figure 18.4 SEM images of typical electrospun materials: (A and B) as electrospun mats of natural silk and (C and D) condensation of silica to in order to fabricate silica silk nanocomposites.
492 Advanced Dental Biomaterials (Greiner and Wendorff, 2007). Electrospun nonwoven mats are porous in a threedimensional format, providing support to cells similar to the extracellular matrix (Li et al., 2002; Agarwal et al., 2008). Electrospun nanofibers have useful properties such as availability of high surface area for protein absorption and binding sites (Stevens and George, 2005) and enhanced cellular attachment and interaction (Bosworth and Downes, 2011). There are various properties of electrospun scaffolds, such as material strength, the porosity can be tailored by nanofiber alignment, and orientation, that can facilitate maximum volume fractions (Bosworth and Downes, 2011). In addition, characteristics such as surface morphology and topography can also be controlled adjusting various solution dope and spinning parameters (Liang et al., 2007). Due to such unique properties, electrospun nanofiber scaffolds have emerged as an excellent candidate for various biological applications such as wound dressings (He et al., 2015a; Zhou et al., 2013; Rho et al., 2006), drug delivery (He et al., 2015a; Li et al., 2005; Sill and von Recum, 2008), tissue engineering scaffolds (Ardeshirzadeh et al., 2015; Dinis et al., 2015; He et al., 2015b; Tang et al., 2015), and dental applications (Zafar et al., 2016; Qasim et al., 2017; Samprasit et al., 2015; Ohkawa et al., 2009). Various types of nanofibers that have been electrospun for dentistry are shown in Table 18.2. Nanomaterial scaffolds, either electrospun or fabricated by any other technique, are being used for a variety of dental tissue regenerations, such as dentin (VallésLluch et al., 2010), pulp (Cavalcanti et al., 2013), ligament (Peh et al., 2007; Park et al., 2010), bone (Saiz et al., 2013), neural tissues (Lee and Livingston Arinzeh, 2011; Yang, 2004), and tooth (Honda et al., 2010). Although electrospun materials exhibited beneficial properties such as biocompatibility, surface smoothness, and controlled release of the incorporated drugs (Zafar et al., 2016; Li et al., 2005; Norowski et al., 2015), further research is needed to translate these materials to clinical applications. 18.5 Potential of nanomaterials Although a lot of research has been carried out to improve the properties of existing materials, there is no dental material that has ideal properties for any specific application (Anusavice and Phillips, 2003). For example, dental amalgam has been used successfully for more than a century however, there are major concerns about mercury toxicity (Eley, 1997a,b; Jones, 1998; Warfvinge, 1995; Smart et al., 1995) and poor aesthetics (Eley, 1997c; Mclean, 1984; Yardley, 1984). Alternatively, the adhesive composites overcome the aesthetic issues; however, poor mechanical properties and technique sensitivity remain the major issues (Moszner and Salz, 2007). In addition, the synthetic materials lack intelligence to respond to environmental stimuli and activate self-repair and regeneration process (Mano and Reis, 2005). Biological components (such as biomolecules, materials scaffolds, and cells) interact with biomaterials at the nanoscale. Materials scientists may benefit from
Nanotechnology and nanomaterials in dentistry 493 Table 18.2 Applications of electrospun nanofibers in dentistry. Applications in dentistry Electrospun nanofibers References Nanocomposites Polyvinyl alcohol/gelatin Silk/silica PVA, PDS Linh and Lee (2012) Foo et al. (2006) Kim et al. (2008a, 2014), Bottino et al. (2015) Kim et al. (2008b, 2014), Ranjbar-Mohammadi et al. (2016), Zhang et al. (2003, 2007), Jin et al. (2002) Samprasit et al. (2015) Tooth tissue engineering Periodontal regeneration Collagen, PLGA, PLLA, PCL, PEO, silk Prevention of caries Reinforcement of resin composites Chitosan Implant surface modification Cartilage tissue engineering PLGA, collagen Drug delivery Poly(ethylene-co-vinyl acetate), PLLA, PLGA Mucosal/wound and repair Poly-L-(lactic acid), PLGA, chitin, chitosan, silk fibroin, collagen PVA, polyacrylonitrile, polystyrene, nylon PCL, PEO, chitosan Wang et al. (2008), Behler et al. (2009), Demir et al. (2014), Borges et al. (2015), Lee et al. (2015a) Ravichandran et al. (2012) Ainola et al. (2016), Subramanian et al. (2004), Levorson et al. (2013), Thorvaldsson et al. (2008), Li et al. (2003) Vacanti et al. (2012), Song et al. (2013), Zeng et al. (2013), Xue et al. (2014), Kim et al. (2004), Verreck et al. (2003), Kenawy et al. (2002) He et al. (2015a), Tang et al. (2015), Thakur et al. (2008), Noh et al. (2006), Blackwood et al. (2008) PCL, Poly carpolactone; PDS, polydiaxonone; PEO, polyethylene oxide; PLGA, poly(lactide-co-glycolide); PLLA, poly-L-lactic acid; PVA, polyvinyl alcohol. using nanotechnology approaches to improve the material tissues interaction as pointed below: 1. To replace lost dental tissues following nature’s principles (biomimetic approaches) and producing biomaterials resembling the properties very closely of the replaced materials. 2. To produce synthetic materials matching morphology and properties similar to natural dental tissues. 3. Use of regenerative medicine and tissue engineering to replace lost dental tissue by regeneration. There are plenty of hopes for nanomaterials in terms of either developing new materials or significant improvements in the properties of existing materials. The scope of nanomaterials in dentistry is bright and will be helpful to enhance the quality of life in patients.
494 18.6 Advanced Dental Biomaterials Conclusive remarks The application of nanomaterials is the most challenging in the field of medicine and biomedical purposes (Bauer et al., 2004). Nanomaterials exhibited promising results for a range of current and potential applications in various disciplines of dentistry. In addition to new materials, nanotechnology helped researchers to improve the physical and mechanical properties of existing materials. For instance, adding nanoparticles can reinforce strength of polymer composites and facilitate a greater surface area for cellular attachment in tissue engineering scaffolds. This is an area of very active research all around the globe involving a lot of research funding. It can be expected in the future that the science of dental materials may change significantly with the better understanding and the introduction of new nanobiomaterials. It can be hoped that promising nanomaterials will provide a variety of materials for dental applications in the next decade. References Abdo, S.B., Al Darrat, A., Luddin, N., Husien, A., 2012. Sealing ability of gutta-percha/nano HA versus resilon/epiphany after 20 months using an electrochemical model—an in vitro study. Braz. J. Oral Sci. 11, 387 391. Abramovitz, I., Beyth, N., Weinberg, G., Borenstein, A., Polak, D., Kesler-Shvero, D., et al., 2012. In vitro biocompatibility of endodontic sealers incorporating antibacterial nanoparticles. J. Nanomater. 2012, 1 9. Agarwal, S., Wendorff, J.H., Greiner, A., 2008. Use of electrospinning technique for biomedical applications. Polymer (Guildf.) 49, 5603 5621. Ainola, M., Tomaszewski, W., Ostrowska, B., Wesolowska, E., Wagner, H.D., Swieszkowski, W., et al., 2016. A bioactive hybrid three-dimensional tissue-engineering construct for cartilage repair. J. Biomater. Appl. 30, 873 885. Available from: https:// doi.org/10.1177/0885328215604069. Al-Bayaty, F.H., Kamaruddin, A.A., Ismail, M.A., Abdulla, M.A., 2013. Formulation and evaluation of a new biodegradable periodontal chip containing thymoquinone in a chitosan base for the management of chronic periodontitis. J. Nanomater. 2013, 16. AlRahabi, M.K., 2017. Evaluation of complications of root canal treatment performed by undergraduate dental students. Libyan J. Med. 12, 1345582. Anusavice, K., Phillips, R., 2003. Phillips’ Science of Dental Materials. Saunders, St. Louis, MO; Great Britain. Arcı́s, R.W., López-Macipe, A., Toledano, M., Osorio, E., Rodrı́guez-Clemente, R., Murtra, J., et al., 2002. Mechanical properties of visible light-cured resins reinforced with hydroxyapatite for dental restoration. Dent. Mater. 18, 49 57. Ardeshirzadeh, B., Anaraki, N.A., Irani, M., Rad, L.R., Shamshiri, S., 2015. Controlled release of doxorubicin from electrospun PEO/chitosan/graphene oxide nanocomposite nanofibrous scaffolds. Mater. Sci. Eng. C 48, 384 390. Basso, M., 2011. Teeth restoration using a high-viscosity glass ionomer cement: the Equias system. J. Minimum Intervention Dent. 4, 74 76.
Nanotechnology and nanomaterials in dentistry 495 Bauer, L.A., Birenbaum, N.S., Meyer, G.J., 2004. Biological applications of high aspect ratio nanoparticles. J. Mater. Chem. 14, 517 526. Behler, K.D., Stravato, A., Mochalin, V., Korneva, G., Yushin, G., Gogotsi, Y., 2009. Nanodiamond-polymer composite fibers and coatings. ACS Nano 3, 363 369. Beyth, N., Houri-Haddad, Y., Baraness-Hadar, L., Yudovin-Farber, I., Domb, A.J., Weiss, E.I., 2008. Surface antimicrobial activity and biocompatibility of incorporated polyethylenimine nanoparticles. Biomaterials 29, 4157 4163. Blackwood, K.A., McKean, R., Canton, I., Freeman, C.O., Franklin, K.L., Cole, D., et al., 2008. Development of biodegradable electrospun scaffolds for dermal replacement. Biomaterials 29, 3091 3104. Borges, A.L., Münchow, E.A., de Oliveira Souza, A.C., Yoshida, T., Vallittu, P.K., et al., 2015. Effect of random/aligned nylon-6/MWCNT fibers on dental resin composite reinforcement. J. Mech. Behav. Biomed. Mater. 48, 134 144. Bosworth, L.A., Downes, S., 2011. Electrospinning for Tissue Regeneration. Woodhead Publishing, Oxford. Bottino, M.C., Thomas, V., Schmidt, G., Vohra, Y.K., Chu, T.G., Kowolik, M.J., et al., 2012. Recent advances in the development of GTR/GBR membranes for periodontal regeneration—a materials perspective. Dent. Mater. 28, 703 721. Bottino, M.C., Yassen, G.H., Platt, J.A., Labban, N., Windsor, L.J., Spolnik, K.J., et al., 2015. A novel three-dimensional scaffold for regenerative endodontics: materials and biological characterizations. J. Tissue Eng. Regen. Med. 9, E116 E123. Brady, M.A., Renzing, A., Douglas, T.E.L., Liu, Q., Wille, S., Parizek, M., et al., 2015. Development of composite poly(lactide-co-glycolide)-nanodiamond scaffolds for bone cell growth. J. Nanosci. Nanotechnol. 15, 1060 1069. Brånemark, P., Breine, U., Adell, R., Hansson, B., Lindström, J., Ohlsson, Å., 1969. Intraosseous anchorage of dental prostheses: I. Experimental studies. Scand. J. Plast. Reconstr. Surg. Hand. Surg. 3, 81 100. Broderick, M., Neville, F., Gibson, T., Millner, P., 2007. Practical nanobiotechnology: functional nanoparticle production using silaffin R5 peptide. J. Biotechnol. 131, S11. Buonocore, M.G., 1955. A simple method of increasing the adhesion of acrylic filling materials to enamel surfaces. J. Dent. Res. 34, 849 853. Cao, W., Hench, L.L., 1996. Bioactive materials. Ceram. Int. 22, 493 507. Cavalcanti, B.N., Zeitlin, B.D., Nör, J.E., 2013. A hydrogel scaffold that maintains viability and supports differentiation of dental pulp stem cells. Dent. Mater. 29, 97 102. Cross, K., Huq, N., Reynolds, E., 2007. Casein phosphopeptides in oral health-chemistry and clinical applications. Curr. Pharm. Des. 13, 793 800. Demarco, F.F., Conde, M., Cavalcanti, B.N., Casagrande, L., Sakai, V.T., Nör, J.E., 2011. Dental pulp tissue engineering. Braz. Dent. J. 22, 3 14. Demir, M.M., Horzum, N., Ta demirci, A., Turan, K., Güden, M., 2014. Mechanical interlocking between porous electrospun polystyrene fibers and an epoxy matrix. ACS Appl. Mater. Interfaces 6, 21901 21905. Deville, S., Chevalier, J., Fantozzi, G., Bartolomé, J.F., Requena, J., Moya, J.S., et al., 2003. Low-temperature ageing of zirconia-toughened alumina ceramics and its implication in biomedical implants. J. Eur. Ceram. Soc. 23, 2975 2982. Dinis, T.M., Elia, R., Vidal, G., Dermigny, Q., Denoeud, C., Kaplan, D.L., et al., 2015. 3D multi-channel bi-functionalized silk electrospun conduits for peripheral nerve regeneration. J. Mech. Behav. Biomed. Mater. 41, 43 55. Duncan, L., Sprehe, C., 2008. Mosby’s Dental Dictionary. Mosby, St. Louis, MO.
496 Advanced Dental Biomaterials Eley, B.M., 1997a. The future of dental amalgam: a review of the literature. 2. Mercury exposure in dental practice. Br. Dent. J. 182, 293 297. Eley, B.M., 1997b. The future of dental amalgam: a review of the literature. 4. Mercury exposure hazards and risk assessment. Br. Dent. J. 182, 373 381. Eley, B.M., 1997c. The future of dental amalgam: a review of the literature. 7. Possible alternative materials to amalgam for the restoration of posterior teeth. Br. Dent. J. 183, 11 14. Elgendy, E.A., Abo Shady, T.E., 2015. Clinical and radiographic evaluation of nanocrystalline hydroxyapatite with or without platelet-rich fibrin membrane in the treatment of periodontal intrabony defects. J. Indian Soc. Periodontol. 19, 61 65. Emsley, J., 1976. The Chemistry of Phosphorus: Environmental, Organic, Inorganic, Biochemical and Spectroscopic Aspects. Harper & Row, London. Esposito, M., 1998. Biological factors contributing to failures of osseointegrated oral implants. (l). Success criteria and epidemiology. Eur. J. Oral Sci. 106, 527 551. Esposito, M., Hirsch, J., Lekholm, U., Thomsen, P., 1998. Biological factors contributing to failures of osseointegrated oral implants, (II). Etiopathogenesis. Eur. J. Oral Sci. 106, 721 764. Farea, M., Masudi, S., Bakar, W., Zaripah, W., 2010. Apical microleakage evaluation of system B compared with cold lateral technique: in vitro study. Aust. Endod. J. 36, 48 53. Feynman, R.P., 1960. There’s plenty of room at the bottom. Eng. Sci. 23, 22 36. Foo, C.W.P., Patwardhan, S.V., Belton, D.J., Brandon, K., Anastasiades, D., Huang, J., et al., 2006. Novel nanocomposites from spider silk-silica fusion (chimeric) proteins. Proc. Natl. Acad. Sci. U.S.A. 103, 9428 9433. Fouad, A., Torabinejad, M., Walton, R.E., 2008. Endodontics: Principles and Practice. Elsevier Health Sciences. Friedman, C.M., Sandrik, J.L., Heuer, M.A., Rapp, G.W., 1975. Composition and mechanical properties of gutta-percha endodontic points. J. Dent. Res. 54, 921 925. Fukuda, J., Khademhosseini, A., Yeh, J., Eng, G., Cheng, J., Farokhzad, O.C., et al., 2006. Micropatterned cell co-cultures using layer-by-layer deposition of extracellular matrix components. Biomaterials 27, 1479 1486. Gahlot, S., Sharma, P.P., Gupta, H., Kulshrestha, V., Jha, P.K., 2014. Preparation of graphene oxide nano-composite ion-exchange membranes for desalination application. RSC Adv. 4, 24662 24670. Galler, K.M., Cavender, A., Yuwono, V., Dong, H., Shi, S., Schmalz, G., et al., 2008. Selfassembling peptide amphiphile nanofibers as a scaffold for dental stem cells. Tissue Eng. Part A 14, 2051 2058. Galler, K.M., D’Souza, R.N., Hartgerink, J.D., 2010. Biomaterials and their potential applications for dental tissue engineering. J. Mater. Chem. 20, 8730 8746. Gismondi, A., Reina, G., Orlanducci, S., Mizzoni, F., Gay, S., Terranova, M.L., et al., 2015. Nanodiamonds coupled with plant bioactive metabolites: a nanotech approach for cancer therapy. Biomaterials 38, 22 35. Gonçalves, J.P.L., Shaikh, A.Q., Reitzig, M., Kovalenko, D.A., Michael, J., Beutner, R., et al., 2014. Detonation nanodiamonds biofunctionalization and immobilization to titanium alloy surfaces as first steps towards medical application. Beilstein J. Org. Chem. 10, 2765 2773. Gouma, P., 2009. Nanomaterials for Chemical Sensors and Biotechnology. World Scientific, Singapore; London. Greiner, A., Wendorff, J.H., 2007. Electrospinning: a fascinating method for the preparation of ultrathin fibres. Angew. Chem. Int. Ed. Engl. 46, 5670 5703.
Nanotechnology and nanomaterials in dentistry 497 Gu, Y., Yap, A., Cheang, P., Koh, Y., Khor, K., 2005. Development of zirconia-glass ionomer cement composites. J. Non-Cryst. Solids 351, 508 514. Hannig, M., Hannig, C., 2010. Nanomaterials in preventive dentistry. Nat. Nanotechnol. 5, 565 569. Hannig, M., Kriener, L., Hoth-Hannig, W., Becker-Willinger, C., Schmidt, H., 2007. Influence of nanocomposite surface coating on biofilm formation in situ. J. Nanosci. Nanotechnol. 7, 4642 4648. Hargreaves, K.M., 2010. Cohen’s Pathways of the Pulp. Mosby, Edinburgh. He, T., Wang, J., Huang, P., Zeng, B., Li, H., Cao, Q., et al., 2015a. Electrospinning polyvinylidene fluoride fibrous membranes containing anti-bacterial drugs used as wound dressing. Colloids Surf., B: Biointerfaces 130, 278 286. He, X., Cheng, L., Zhang, X., Xiao, Q., Zhang, W., Lu, C., 2015b. Tissue engineering scaffolds electrospun from cotton cellulose. Carbohydr. Polym. 115, 485 493. Hieber, S., Müller, B., 2012. In: Logothetidis, S. (Ed.), Nanodentistry. Springer, Berlin Heidelberg. Honda, M.J., Tsuchiya, S., Shinohara, Y., Shinmura, Y., Sumita, Y., 2010. Recent advances in engineering of tooth and tooth structures using postnatal dental cells. Jpn. Dent. Sci. Rev. 46, 54 66. Horst, O.V., Chavez, M.G., Jheon, A.H., Desai, T., Klein, O.D., 2012. Stem cell and biomaterials research in dental tissue engineering and regeneration. Dent. Clin. North Am. 56, 495 520. Husain, S., Al-Samadani, K.H., Najeeb, S., Zafar, M.S., Khurshid, Z., Zohaib, S., et al., 2017. Chitosan biomaterials for current and potential dental applications. Materials 10, 602. Hussain, F., Hojjati, M., Okamoto, M., Gorga, R.E., 2006. Review article: polymer-matrix nanocomposites, processing, manufacturing, and application: an overview. J. Compos. Mater. 40, 1511. Huynh, V.T., Pearson, S., Noy, J., Abboud, A., Utama, R.H., Lu, H., et al., 2013. Nanodiamonds with surface grafted polymer chains as vehicles for cell imaging and cisplatin delivery: enhancement of cell toxicity by POEGMEMA coating. ACS Macro Lett. 2, 246 250. Jevnikar, P., Golobič, M., Kocjan, A., Kosmač, T., 2012. The effect of nano-structured alumina coating on the bond strength of resin-modified glass ionomer cements to zirconia ceramics. J. Eur. Ceram. Soc. 32, 2641 2645. Jin, H.J., Fridrikh, S., Rutledge, G.C., Kaplan, D., 2002. Electrospinning Bombyx mori silk with poly(ethylene oxide). Biomacromolecules 3, 1233 1239. Jones, D.W., 1998. A Canadian perspective on the dental amalgam issue. Br. Dent. J. 184, 581 586. Jones, F.H., 2001. Teeth and bones: applications of surface science to dental materials and related biomaterials. Surf. Sci. Rep. 42, 75 205. Jurczyk, M.U., Jurczyk, K., Niespodziana, K., Miklaszewski, A., Jurczyk, M., 2013. Titanium SiO2 nanocomposites and their scaffolds for dental applications. Mater Charact. 77, 99 108. Kang, M., Chen, P., Jin, H., 2009. Preparation of multiwalled carbon nanotubes incorporated silk fibroin nanofibers by electrospinning. Curr. Appl. Phys. 9, S95 S97. Kelsall, R., Hamley, I.W., Geoghegan, M., 2005. Nanoscale Science and Technology. Wiley, John Wiley & Sons, Ltd. Kenawy, E., Bowlin, G.L., Mansfield, K., Layman, J., Simpson, D.G., Sanders, E.H., et al., 2002. Release of tetracycline hydrochloride from electrospun poly(ethylene-co-vinylacetate), poly(lactic acid), and a blend. J. Controlled Release 81, 57 64.
498 Advanced Dental Biomaterials Khademolhosseini, M., Barounian, M., Eskandari, A., Aminzare, M., Zahedi, A., Ghahremani, D., 2012. Development of new Al2O3/TiO2 reinforced glass-ionomer cements (GICs) nano-composites. J. Basic Appl. Sci. Res. 2, 7526 7529. Khurshid, Z., Zafar, M., Qasim, S., Shahab, S., Naseem, M., AbuReqaiba, A., 2015. Advances in nanotechnology for restorative dentistry. Materials 8, 717 731. Kim, K., Luu, Y.K., Chang, C., Fang, D., Hsiao, B.S., Chu, B., et al., 2004. Incorporation and controlled release of a hydrophilic antibiotic using poly(lactide-co-glycolide)-based electrospun nanofibrous scaffolds. J. Controlled Release 98, 47 56. Kim, G., Asran, A.S., Michler, G.H., Simon, P., Kim, J., 2008a. Electrospun PVA/HAp nanocomposite nanofibers: biomimetics of mineralized hard tissues at a lower level of complexity. Bioinspir. Biomim. 3, 046003. Kim, I., Seo, S., Moon, H., Yoo, M., Park, I., Kim, B., et al., 2008b. Chitosan and its derivatives for tissue engineering applications. Biotechnol. Adv. 26, 1 21. Kim, N.R., Lee, D.H., Chung, P., Yang, H., 2009. Distinct differentiation properties of human dental pulp cells on collagen, gelatin, and chitosan scaffolds. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 108, e94 e100. Kim, B., Nguyen, T.B.L., Min, Y., Lee, B., 2014a. In vitro and in vivo studies of BMP-2loaded PCL gelatin BCP electrospun scaffolds. Tissue Eng. Part A 20, 3279 3289. Kim, J.J., Bae, W.J., Kim, J.M., Kim, J.J., Lee, E.J., Kim, H.W., et al., 2014b. Mineralized polycaprolactone nanofibrous matrix for odontogenesis of human dental pulp cells. J. Biomater. Appl. 28, 1069 1078. Koch, K., Brave, D., 2009. A new day has dawned: the increased use of bioceramics in endodontics. Dent. Town 10, 39 43. Kostarelos, K., Bianco, A., Prato, M., 2009. Promises, facts and challenges for carbon nanotubes in imaging and therapeutics. Nat. Nanotechnol. 4, 627 633. Kulkarni Aranya, A., Pushalkar, S., Zhao, M., LeGeros, R.Z., Zhang, Y., Saxena, D., 2017. Antibacterial and bioactive coatings on titanium implant surfaces. J. Biomed. Mater. Res. A 105, 2218 2227. Langer, R., Vacanti, J., 1993. Tissue engineering. Science 260, 920 926. Le Guéhennec, L., Soueidan, A., Layrolle, P., Amouriq, Y., 2007. Surface treatments of titanium dental implants for rapid osseointegration. Dent. Mater. 23, 844 854. Lee, Y., Livingston Arinzeh, T., 2011. Electrospun nanofibrous materials for neural tissue engineering. Polymers 3, 413 426. Lee, J., Lee, Y., Choi, B., Lee, J., Choi, H., Son, H., et al., 2010. Physical properties of resin-reinforced glass ionomer cement modified with micro and nano-hydroxyapatite. J. Nanosci. Nanotechnol. 10, 5270 5276. Lee, M.W., An, S., Jo, H.S., Yoon, S.S., Yarin, A.L., 2015a. Self-healing nanofiber-reinforced polymer composites: tensile testing and recovery of mechanical properties. ACS Appl. Mater. Interfaces 7, 19546 19554. Lee, D., Kim, S.V., Limansubroto, A.N., Yen, A., Soundia, A., Wang, C., et al., 2015b. Nanodiamond-gutta percha composite biomaterials for root canal therapy. ACS Nano 9, 11490 11501. LeGeros, R.Z., 1991. Calcium Phosphates in Oral Biology and Medicine. Karger, Basel. Levorson, E.J., Sreerekha, P.R., Chennazhi, K.P., Kasper, F.K., Nair, S.V., Mikos, A.G., 2013. Fabrication and characterization of multiscale electrospun scaffolds for cartilage regeneration. Biomed. Mater. 8, 014103. Li, W., Laurencin, C.T., Caterson, E.J., Tuan, R.S., Ko, F.K., 2002. Electrospun nanofibrous structure: a novel scaffold for tissue engineering. J. Biomed. Mater. Res. 60, 613 621.
Nanotechnology and nanomaterials in dentistry 499 Li, W., Danielson, K.G., Alexander, P.G., Tuan, R.S., 2003. Biological response of chondrocytes cultured in three-dimensional nanofibrous poly(E-caprolactone) scaffolds. J. Biomed. Mater. Res. A 67, 1105 1114. Li, W., Mauck, R.L., Tuan, R.S., 2005. Electrospun nanofibrous scaffolds: production, characterization, and applications for tissue engineering and drug delivery. J. Biomed. Nanotechnol. 1, 259 275. Liang, D., Hsiao, B.S., Chu, B., 2007. Functional electrospun nanofibrous scaffolds for biomedical applications. Adv. Drug Deliv. Rev. 59, 1392 1412. Lieber, C.M., 1998. One-dimensional nanostructures: chemistry, physics & applications. Solid State Commun. 107, 607 616. Linh, N.T.B., Lee, B., 2012. Electrospinning of polyvinyl alcohol/gelatin nanofiber composites and cross-linking for bone tissue engineering application. J. Biomater. Appl. 27, 255 266. Low, I.M., Duraman, N., Mahmood, U., 2008. Mapping the structure, composition and mechanical properties of human teeth. Mater. Sci. Eng. C 28, 243 247. Lv, K.L., Zhang, J.X., Meng, X.C., Li, X.Y., 2007. Remineralization effect of the nano-HA toothpaste on artificial caries. Key Eng. Mat. 330, 267 270. Madarati, A.A., Zafar, M.S., Sammani, A.M., Mandorah, A.O., Bani-Younes, H.A., 2017. Preference and usage of intracanal medications during endodontic treatment. Saudi Med. J. 38, 755 763. Madhavan, S., Nayak, M., Shenoy, A., Shetty, R., Prasad, K., 2012. Dentinal hypersensitivity: a comparative clinical evaluation of CPP-ACP F, sodium fluoride, propolis, and placebo. J. Conserv. Dent. 15, 315. Mano, J.F., Reis, R.L., 2005. Some trends on how one can learn from and mimic nature in order to design better biomaterials. Mater. Sci. Eng. C 25, 93 95. Matinlinna, J.P., Tsoi, J.K., de Vries, J., Busscher, H.J., 2013. Characterization of novel silane coatings on titanium implant surfaces. Clin. Oral Implants Res. 24, 688 697. Mclean, J.W., 1984. Alternatives to amalgam alloys. 1. Br. Dent. J. 157, 432 433. Mendes, V.C., Moineddin, R., Davies, J.E., 2007. The effect of discrete calcium phosphate nanocrystals on bone-bonding to titanium surfaces. Biomaterials 28, 4748 4755. Mendes, V.C., Moineddin, R., Davies, J.E., 2008. Discrete calcium phosphate nanocrystalline deposition enhances osteoconduction on titanium-based implant surfaces. J. Biomed. Mater. Res. A 90, 577 585. Mendonça, G., Mendonça, D.B.S., Aragão, F.J.L., Cooper, L.F., 2008. Advancing dental implant surface technology—from micron- to nanotopography. Biomaterials 29, 3822 3835. Mickenautsch, S., Yengopal, V., Banerjee, A., 2012. Retention of orthodontic brackets bonded with resin-modified GIC versus composite resin adhesives—a quantitative systematic review of clinical trials. Clin. Oral Investig. 16, 1 14. Mitra, S.B., Wu, D., Holmes, B.N., 2003. An application of nanotechnology in advanced dental materials. J. Am. Dent. Assoc. 134, 1382 1390. Mizuno, M., Miyamoto, T., Wada, K., Watatani, S., Zhang, G.X., 2003. Type I collagen regulated dentin matrix protein-1 (Dmp-1) and osteocalcin (OCN) gene expression of rat dental pulp cells. J. Cell. Biochem. 88, 1112 1119. Mohanty, A., Drzal, L., Misra, M., 2003. Nano reinforcements of bio-based polymers—the hope and the reality. J. Am. Chem. Soc. 225, 33. Moore, L., Gatica, M., Kim, H., Osawa, E., Ho, D., 2013. Multi-protein delivery by nanodiamonds promotes bone formation. J. Dent. Res. 92, 976 981.
500 Advanced Dental Biomaterials Morgan, M.V., Adams, G.G., Bailey, D.L., Tsao, C.E., Fischman, S.L., Reynolds, E.C., 2008. The anticariogenic effect of sugar-free gum containing CPP-ACP nanocomplexes on approximal caries determined using digital bitewing radiography. Caries Res. 42, 171 184. Mortazavi, V., Nahrkhalaji, M.M., Fathi, M.H., Mousavi, S.B., Esfahani, B.N., 2010. Antibacterial effects of sol-gel-derived bioactive glass nanoparticle on aerobic bacteria. J. Biomed. Mater. Res. A 94A, 160 168. Moshaverinia, A., Ansari, S., Moshaverinia, M., Roohpour, N., Darr, J.A., Rehman, I., 2008. Effects of incorporation of hydroxyapatite and fluoroapatite nanobioceramics into conventional glass ionomer cements (GIC). Acta Biomater. 4, 432 440. Moszner, N., Klapdohr, S., 2004. Nanotechnology for dental composites. Int. J. Nanotechnol. 1, 130 156. Moszner, N., Salz, U., 2007. Recent developments of new components for dental adhesives and composites. Macromol. Mater. Eng. 292, 245 271. Mota, E.G., Oshima, H., Burnett Jr., L.H., Pires, L., Rosa, R.S., 2006. Evaluation of diametral tensile strength and Knoop microhardness of five nanofilled composites in dentin and enamel shades. Stomatologija 8, 67 69. Mount, G.J., 1998. Clinical performance of glass-ionomers. Biomaterials 19, 573 579. Musselman, M., 2003. Composites make large difference in “small” medical, dental applications. Compos. Technol. 9, 24 27. Najeeb, S., Khurshid, Z., Matinlinna, J.P., Siddiqui, F., Nassani, M.Z., Baroudi, K., 2015. Nanomodified peek dental implants: bioactive composites and surface modification—a review. Int. J. Dent. 2015, (Article ID 381759, in press). Najeeb, S., Khurshid, Z., Zafar, M.S., Khan, A.S., Zohaib, S., Martı́, J.M.N., et al., 2016a. Modifications in glass ionomer cements: nano-sized fillers and bioactive nanoceramics. Int. J. Mol. Sci. 17, 1134. Najeeb, S., Khurshid, Z., Zohaib, S., Zafar, M.S., 2016b. Bioactivity and osseointegration of PEEK are inferior to those of titanium—a systematic review. J. Oral Implantol. 42, 512 516. Najeeb, S., Zafar, M.S., Khurshid, Z., Siddiqui, F., 2016c. Applications of polyetheretherketone (PEEK) in oral implantology and prosthodontics. J. Prosthodont. Res. 60, 12 19. Najeeb, S., Khurshid, Z., Agwan, A.S., Zafar, M.S., Alrahabi, M., Qasim, S.B., et al., 2016d. Dental applications of nanodiamonds. Sci. Adv. Mater. 8, 2064 2070. Najeeb, S., Zafar, M.S., Khurshid, Z., Zohaib, S., Hasan, S.M., Khan, R.S., 2017. Bisphosphonate releasing dental implant surface coatings and osseointegration: a systematic review. J. Taibah Univ. Med. Sci. 12, 369 375. Nanci, A., 2008. Ten Cate’s Oral Histology: Development, Structure, and Function. Mosby, St. Louis, MO; London. Naray-Szabo, S., 1930. Structure of apatite (CaF)Ca4(PO4)3. Z. Kristallogr. 75, 387 398. Noh, H.K., Lee, S.W., Kim, J., Oh, J., Kim, K., Chung, C., et al., 2006. Electrospinning of chitin nanofibers: degradation behavior and cellular response to normal human keratinocytes and fibroblasts. Biomaterials 27, 3934 3944. Norowski, P.A., Fujiwara, J., Clem, T., Adatrow, W.C., Eckstein, P.C., Haggard, E.C., et al., 2015. Novel naturally crosslinked electrospun nanofibrous chitosan mats for guided bone regeneration membranes: material characterization and cytocompatibility. J. Tissue Eng. Regen. Med. 9, 577 583. Oh, T., Yoon, J., Misch, C.E., Wang, H., 2002. The causes of early implant bone loss: myth or science? J. Periodontol. 73, 322 333.
Nanotechnology and nanomaterials in dentistry 501 Ohkawa, K., Hayashi, S., Kameyama, N., Yamamoto, H., Yamaguchi, M., Kimoto, S., et al., 2009. Synthesis of collagen-like sequential polypeptides containing O-phospho-Lhydroxyproline and preparation of electrospun composite fibers for possible dental application. Macromol. Biosci. 9, 79 92. Paquette, D., Hanlon, A., Lessem, J., Williams, R., 2004. Clinical relevance of adjunctive minocycline microspheres in patients with chronic periodontitis: secondary analysis of a phase 3 trial. J. Periodontol. 75, 531 536. Parizek, M., Douglas, T.E.L., Novotna, K., Kromka, A., Brady, M.A., Renzing, A., et al., 2012. Nanofibrous poly(lactide-co-glycolide) membranes loaded with diamond nanoparticles as promising substrates for bone tissue engineering. Int. J. Nanomed. 7, 1931. Park, C.H., Rios, H.F., Jin, Q., Bland, M.E., Flanagan, C.L., Hollister, S.J., et al., 2010. Biomimetic hybrid scaffolds for engineering human tooth-ligament interfaces. Biomaterials 31, 5945 5952. Passeri, D., Rinaldi, F., Ingallina, C., Carafa, M., Rossi, M., Terranova, M., et al., 2015. Biomedical applications of nanodiamonds: an overview. J. Nanosci. Nanotechnol. 15, 972 988. Paul, D.R., Robeson, L.M., 2008. Polymer nanotechnology: nanocomposites. Polymer (Guildf.) 49, 3187 3204. Pecharroman, C., Bartolomé, J.F., Requena, J., Moya, J.S., Deville, S., Chevalier, J., et al., 2003. Percolative mechanism of aging in zirconia containing ceramics for medical applications. Adv. Mater. 15, 507 511. Peh, R.F., Suthikum, V., Goh, C.H., Toh, S.L., 2007. Novel electrospun-knitted silk scaffolds for ligament tissue engineering. World Congr. Med. Phys. Biomed. Eng. 2006 (4), 3287 3290. Pts 1-6. 14. Pichayakorn, W., Boonme, P., 2013. Evaluation of cross-linked chitosan microparticles containing metronidazole for periodontitis treatment. Mater. Sci. Eng. C 33, 1197 1202. Pihlstrom, B.L., Michalowicz, B.S., Johnson, N.W., 2005. Periodontal diseases. Lancet 366, 1809 1820. Piñón-Segundo, E., Ganem-Quintanar, A., Alonso-Pérez, V., Quintanar-Guerrero, D., 2005. Preparation and characterization of triclosan nanoparticles for periodontal treatment. Int. J. Pharm. 294, 217 232. Pye, A., Lockhart, D., Dawson, M., Murray, C., Smith, A., 2009. A review of dental implants and infection. J. Hosp. Infect. 72, 104 110. Qasim, S.B., Najeeb, S., Delaine-Smith, R.M., Rawlinson, A., Rehman, I.U., 2017. Potential of electrospun chitosan fibers as a surface layer in functionally graded GTR membrane for periodontal regeneration. Dent. Mater. 33, 71 83. Qasim, S.B., Zafar, M.S., Najeeb, S., Khurshid, Z., Shah, A.H., Husain, S., et al., 2018. Electrospinning of chitosan-based solutions for tissue engineering and regenerative medicine. Int. J. Mol. Sci. 19, 407. Raggio, D.P., Calvo, A.F.B., Branco, E.C., Bonifácio, C.C., Imparato, J.C.P., Camargo, L.B., 2009. Knoop hardness of resin-modified glass-ionomer cements. Int. J. Dent. 8, 119 123. Rahiotis, C., Vougiouklakis, G., Eliades, G., 2008. Characterization of oral films formed in the presence of a CPP ACP agent: an in situ study. J. Dent. 36, 272 280. Ranjbar-Mohammadi, M., Zamani, M., Prabhakaran, M., Bahrami, S.H., Ramakrishna, S., 2016. Electrospinning of PLGA/gum tragacanth nanofibers containing tetracycline hydrochloride for periodontal regeneration. Mater. Sci. Eng. C 58, 521 531.
502 Advanced Dental Biomaterials Ravichandran, R., Ng, C.C., Liao, S., Pliszka, D., Raghunath, M., Ramakrishna, S., et al., 2012. Biomimetic surface modification of titanium surfaces for early cell capture by advanced electrospinning. Biomed. Mater. 7, 015001. Ray, S.S., Okamoto, M., 2003. Biodegradable polylactide and its nanocomposites: opening a new dimension for plastics and composites. Macromol. Rapid Commun. 24, 815 840. Reynolds, E., 2008. Calcium phosphate-based remineralization systems: scientific evidence? Aust. Dent. J. 53, 268 273. Rho, K.S., Jeong, L., Lee, G., Seo, B.M., Park, Y.J., Hong, S.D., et al., 2006. Electrospinning of collagen nanofibers: effects on the behavior of normal human keratinocytes and early-stage wound healing. Biomaterials 27, 1452 1461. Rizvi, A., Zafar, M.S., Farid, W.M., Gazal, G., 2014. Assessment of antimicrobial efficacy of MTAD, sodium hypochlorite, EDTA and chlorhexidine for endodontic applications: an in vitro study. Middle East J. Sci. Res. 21, 353 357. Rizvi, A., Zafar, M., Al-Wasifi, Y., Fareed, W., Khurshid, Z., 2016. Role of enamel deminerlization and remineralization on microtensile bond strength of resin composite. Eur. J. Dent. 10, 376 380. Rodrı́guez-Lozano, F.J., GarcÃa-Bernal, D., Aznar-Cervantes, S., Ros-Roca, M., AlguerÃ, M.C., Atucha, N.M., et al., 2014. Effects of composite films of silk fibroin and graphene oxide on the proliferation, cell viability and mesenchymal phenotype of periodontal ligament stem cells. J. Mater. Sci. Mater. Med. 25, 2731 2741. Roveri, N., Battistella, E., Bianchi, C.L., Foltran, I., Foresti, E., Iafisco, M., et al., 2009. Surface enamel remineralization: biomimetic apatite nanocrystals and fluoride ions different effects. J. Nanomater. 2009, 8. Saiz, E., Zimmermann, E.A., Lee, J.S., Wegst, U.G.K., Tomsia, A.P., 2013. Perspectives on the role of nanotechnology in bone tissue engineering. Dent. Mater. 29, 103 115. Sakaguchi, R.L., Powers, J.M., 2012. Craig’s Restorative Dental Materials. Elsevier/Mosby, Philadelphia, PA. Samprasit, W., Kaomongkolgit, R., Sukma, M., Rojanarata, T., Ngawhirunpat, T., Opanasopit, P., 2015. Mucoadhesive electrospun chitosan-based nanofibre mats for dental caries prevention. Carbohydr. Polym. 117, 933 940. Sato, M., Sambito, M.A., Aslani, A., Kalkhoran, N.M., Slamovich, E.B., Webster, T.J., 2006. Increased osteoblast functions on undoped and yttrium-doped nanocrystalline hydroxyapatite coatings on titanium. Biomaterials 27, 2358 2369. Saunders, W.P., Saunders, E.M., 1994. Coronal leakage as a cause of failure in root-canal therapy: a review. Dent. Traumatol. 10, 105 108. Setyawati, M.I., Mochalin, V.N., Leong, D.T., 2015. Tuning endothelial permeability with functionalized nanodiamonds. ACS Nano 10, 1170 1181. Sheeparamatti, B., Sheeparamatti, R., Kadadevaramath, J., 2007. Nanotechnology: inspiration from nature. IETE Tech. Rev. 24, 5 8. Sheikh, Z., Najeeb, S., Khurshid, Z., Verma, V., Rashid, H., Glogauer, M., 2015. Biodegradable materials for bone repair and tissue engineering applications. Materials 8, 5744 5794. Sill, T.J., von Recum, H.A., 2008. Electrospinning: applications in drug delivery and tissue engineering. Biomaterials 29, 1989 2006. Silva, R.M., Santos, P.H.N., Souza, L.B., Dumont, V.C., Soares, J.A., Santos, M.H., 2013. Effects of cellulose fibers on the physical and chemical properties of glass ionomer dental restorative materials. Mater. Res. Bull. 48, 118 126. Silva, L.H., Belli, R., Lazar, D.R.R., Tango, R.N., Lohbauer, U., Cesar, P.F., 2014. Y-TZP reinforced with carbon nanotubes: synthesis, microstructure and flexural strength. Dent. Mater. 30, e9.
Nanotechnology and nanomaterials in dentistry 503 Siqueira, J.F., 2001. Aetiology of root canal treatment failure: why well-treated teeth can fail. Int. Endod. J. 34, 1 10. Smart, E.R., Macleod, R.I., Lawrence, C.M., 1995. Resolution of lichen-planus following removal of amalgam restorations in patients with proven allergy to mercury salts—a pilot-study. Br. Dent. J. 178, 108 112. Song, W., Yu, X., Markel, D.C., Shi, T., Ren, W., 2013. Coaxial PCL/PVA electrospun nanofibers: osseointegration enhancer and controlled drug release device. Biofabrication 5, 035006. Stevens, M.M., George, J.H., 2005. Exploring and engineering the cell surface. Interface Sci. 310, 1135 1138. Subramanian, A., Lin, H.Y., Vu, D., Larsen, G., 2004. Synthesis and evaluation of scaffolds prepared from chitosan fibers for potential use in cartilage tissue engineering. Biomed. Sci. Instrum. 40, 117 122. Suliman, S., Xing, Z., Wu, X., Xue, Y., Pedersen, T.O., Sun, Y., et al., 2015. Release and bioactivity of bone morphogenetic protein-2 are affected by scaffold binding techniques in vitro and in vivo. J. Controlled Release 197, 148 157. Summit, J., Robins, W., Hilton, T., Schwartz, R., 2000. Fundamentals of Operative Dentistry: A Contemporary Approach. Quintessence Publishing, Chicago, IL; London. Sun, X., Liu, Z., Welsher, K., Robinson, J.T., Goodwin, A., Zaric, S., et al., 2008. Nanographene oxide for cellular imaging and drug delivery. Nano Res. 1, 203 212. Tang, J., Han, Y., Zhang, F., Ge, Z., Liu, X., Lu, Q., 2015. Buccal mucosa repair with electrospun silk fibroin matrix in a rat model. Int. J. Artif. Organs 38, 105 112. Taniguchi, N., 1974. On the Basic Concept of ‘Nano-Technology’ Proceedings of the International Conference of Production Engineering. Society of Precision Engineering, Tokyo, Japan. Ten Cate, A.R., 1994. Oral Histology: Development, Structure, and Function. Mosby, St. Louis, MO; London. Ten Cate, J.M., van Duinen, R.N.B., 1995. Hypermineralization of dentinal lesions adjacent to glass-ionomer cement restorations. J. Dent. Res. 74, 1266 1271. Terry, D.A., 2004. Direct applications of a nanocomposite resin system: Part 1—The evolution of contemporary composite materials. Pract. Proced. Aesthet. Dent. 16, 417 432. Thakur, R., Florek, C., Kohn, J., Michniak, B., 2008. Electrospun nanofibrous polymeric scaffold with targeted drug release profiles for potential application as wound dressing. Int. J. Pharm. 364, 87 93. Thorvaldsson, A., Stenhamre, H., Gatenholm, P., Walkenström, P., 2008. Electrospinning of highly porous scaffolds for cartilage regeneration. Biomacromolecules 9, 1044 1049. Tomsia, A.P., Launey, M.E., Lee, J.S., Mankani, M.H., Wegst, U.G., Saiz, E., 2011. Nanotechnology approaches for better dental implants. Int. J. Oral Maxillofac. Implants 26, 25. Turssi, C.P., Ferracane, J.L., Vogel, K., 2005. Filler features and their effects on wear and degree of conversion of particulate dental resin composites. Biomaterials 26, 4932 4937. Tyas, M.J., 2006. Clinical evaluation of glass-ionomer cement restorations. J. Appl. Oral Sci. 14, 10. Vacanti, N.M., Cheng, H., Hill, P.S., Guerreiro, J.D., Dang, T.T., Ma, M., et al., 2012. Localized delivery of dexamethasone from electrospun fibers reduces the foreign body response. Biomacromolecules 13, 3031 3038. Valiev, R.Z., Zehetbauer, M.J., Estrin, Y., Höppel, H.W., Ivanisenko, Y., Hahn, H., et al., 2007. The innovation potential of bulk nanostructured materials. Adv. Eng. Mater. 9, 527 533.
504 Advanced Dental Biomaterials Vallés-Lluch, A., Novella-Maestre, E., Sancho-Tello, M., Pradas, M.M., Ferrer, G.G., Batalla, C.C., 2010. Mimicking natural dentin using bioactive nanohybrid scaffolds for dentinal tissue engineering. Tissue Eng. Part A 16, 2783 2793. Vallet-Regı́, M., Arcos, D., 2008. Biomimetic Nanoceramics in Clinical Use: From Materials to Applications. Royal Society of Chemistry. Venugopal, J., Ramakrishna, S., 2005. Applications of polymer nanofibers in biomedicine and biotechnology. Appl. Biochem. Biotechnol. 125, 147 157. Verreck, G., Chun, I., Peeters, J., Rosenblatt, J., Brewster, M.E., 2003. Preparation and characterization of nanofibers containing amorphous drug dispersions generated by electrostatic spinning. Pharm. Res. 20, 810 817. Waltimo, T., Brunner, T.J., Vollenweider, M., Stark, W.J., Zehnder, M., 2007. Antimicrobial effect of nanometric bioactive glass 45S5. J. Dent. Res. 86, 754 757. Wang, W., Ciselli, P., Kuznetsov, E., Peijs, T., Barber, A.H., 2008. Effective reinforcement in carbon nanotube-polymer composites. Philos. Trans. A Math. Phys. Eng. Sci. 366, 1613 1626. Warfvinge, K., 1995. Mercury exposure of a female dentist before pregnancy. Br. Dent. J. 178, 149 152. Wiesmann, H., Meyer, U., Plate, U., Höhling, H., 2004. Aspects of collagen mineralization in hard tissue formation. Int. Rev. Cytol. 242, 121 156. Wilson, A.D., Kent, B.E., 1971. The glass-ionomer cement, a new translucent dental filling material. J. Appl. Chem. Biotechnol. 21, 313. Xing, Z., Pedersen, T.O., Wu, X., Xue, Y., Sun, Y., Finne-Wistrand, A., et al., 2013. Biological effects of functionalizing copolymer scaffolds with nanodiamond particles. Tissue Eng. Part A 19, 1783 1791. Xu, T., Wu, L., Yu, Y., Li, W., Zhi, J., 2014. Synthesis and characterization of diamond silver composite with anti-bacterial property. Mater. Lett. 114, 92 95. Xue, J., He, M., Niu, Y., Liu, H., Crawford, A., Coates, P., et al., 2014. Preparation and in vivo efficient anti-infection property of GTR/GBR implant made by metronidazole loaded electrospun polycaprolactone nanofiber membrane. Int. J. Pharm. 475, 566 577. Yang, F., 2004. Fabrication of nano-structured porous PLLA scaffold intended for nerve tissue engineering. Biomaterials 25, 1891. Yardley, R.M., 1984. Alternatives to amalgam alloys. 2. Br. Dent. J. 157, 434 435. Zafar, M.S., Al-Samadani, K.H., 2014. Potential use of natural silk for bio-dental applications. J. Taibah Univ. Med. Sci. 9, 171 177. Zafar, M., Khurshid, Z., Almas, K., 2015a. Oral tissue engineering progress and challenges. Tissue Eng. Regener. Med. 12, 387 397. Zafar, M.S., Belton, D.J., Hanby, B., Kaplan, D.L., Perry, C.C., 2015b. Functional material features of Bombyx mori silk light vs. heavy chain proteins. Biomacromolecules 16, 606 614. Zafar, M., Najeeb, S., Khurshid, Z., Vazirzadeh, M., Zohaib, S., Najeeb, B., et al., 2016. Potential of electrospun nanofibers for biomedical and dental applications. Materials 9, 73. Zafar, M.S., Khurshid, Z., Najeeb, S., Zohaib, S., Rehman, I.U., 2017. Chapter 26— Therapeutic applications of nanotechnology in dentistry. In: Andronescu, E., Grumezescu, A.M. (Eds.), Nanostructures for Oral Medicine. Elsevier. Zeng, H., Li, F., Wei, H., Shi, J.F., Rao, G.Z., Li, A., et al., 2013. Preliminary study of the dual release baicalin and rhBMP-2 system to improve periodontal tissue regeneration in minipigs. Shanghai Kou Qiang Yi Xue 22, 126 131. Zhang, Y., Ni, M., Zhang, M., Ratner, B., 2003. Calcium phosphate-chitosan composite scaffolds for bone tissue engineering. Tissue Eng. 9, 337 345.
Nanotechnology and nanomaterials in dentistry 505 Zhang, Y., Lim, C.T., Ramakrishna, S., Huang, Z.M., 2005. Recent development of polymer nanofibers for biomedical and biotechnological applications. J. Mater. Sci. Mater. Med. 16, 933 946. Zhang, Y., Su, B., Ramakrishna, S., Lim, C., 2007. Chitosan nanofibers from an easily electrospinnable UHMWPEO-doped chitosan solution system. Biomacromolecules 9, 136 141. Zhang, Q., Mochalin, V.N., Neitzel, I., Knoke, I.Y., Han, J., Klug, C.A., et al., 2011. Fluorescent PLLA-nanodiamond composites for bone tissue engineering. Biomaterials 32, 87 94. Zhang, Q., Mochalin, V.N., Neitzel, I., Hazeli, K., Niu, J., Kontsos, A., et al., 2012. Mechanical properties and biomineralization of multifunctional nanodiamond-PLLA composites for bone tissue engineering. Biomaterials 33, 5067 5075. Zhou, Y., Yang, H., Liu, X., Mao, J., Gu, S., Xu, W., 2013. Electrospinning of carboxyethyl chitosan/poly(vinyl alcohol)/silk fibroin nanoparticles for wound dressings. Int. J. Biol. Macromol. 53, 88 92. Zoufan, K., Jiang, J., Komabayashi, T., Wang, Y., Safavi, K.E., Zhu, Q., 2011. Cytotoxicity evaluation of Gutta Flow and Endo Sequence BC sealers. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 112, 657 661.
19 Digital dentistry Touraj Nejatian1,2, Sanam Almassi3, Azita Farhadi Shamsabadi4, Gaurav Vasudeva5,6, Zoe Hancox7, Amritpaul Singh Dhillon9 and Farshid Sefat7,8 1 Eastman Dental Institute, University College of London, London, United Kingdom, 2 Notthingham Dental Clinic, Nottingham, United Kingdom, 3Almassi Specialist Clinic, Tehran, Iran, 4Centre for English Language Education, Nottingham University, Nottingham, United Kingdom, 5School of Dentistry, James Cook University, Townsville, QLD, Australia, 6Oral Health Services, Hobart, TAS, Australia, 7Biomedical and Electrical Engineering Department, School of Engineering, University of Bradford, Bradford, United Kingdom, 8Interdisciplinary Research Center in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom, 9Dental Institute, King’s College London, London, United Kingdom Chapter Outline 19.1 Introduction 508 19.2 Digital radiography and magnetic resonance imaging 508 19.2.1 Intraoral, extraoral, including cone beam computed tomography 508 19.2.2 Clinical applications 509 19.2.3 Limitations 512 19.3 Caries detection 516 19.4 Photography and shade selection 516 19.5 Computer-aided design computer-aided manufacturing systems in dentistry 517 19.5.1 19.5.2 19.5.3 19.5.4 Chairside milling 518 Laboratory and industrial milling 518 Machining of the restorations 519 Three-dimensional printing 519 19.6 Computer-supported implant dentistry 19.6.1 19.6.2 19.6.3 19.6.4 19.6.5 19.6.6 19.6.7 19.6.8 19.6.9 521 Three-dimensional printing in implant dentistry 521 Recent advances in implant technologies 522 Computer-guided implant surgery 524 Computer-navigated implant surgery 524 Computer-aided design computer-aided manufacturing systems in implant restorative dentistry 527 Prosthetic abutments 528 Computer-aided design computer-aided manufacturing abutments in implant dentistry 528 Materials used 529 Computer-aided design computer-aided manufacturing custom implant abutments 529 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00019-0 Copyright © 2019 Elsevier Ltd. All rights reserved.
508 Advanced Dental Biomaterials 19.7 Lasers and dental applications 530 19.7.1 History of lasers in dentistry 531 19.7.2 Types of lasers 531 19.7.3 Mechanism of laser action 532 19.8 Technology and dental education References 535 Further reading 540 19.1 533 Introduction Dentistry like other branches of health science has benefited enormously from digital technologies. Computers are faster and more accurate than their human counterparts. Nowadays almost all aspects of contemporary clinical dentistry from admission to treatment are assisted by digital technology. Computers have also helped to balance the cost of dental services despite the increasing price of dental products and materials. This chapter provides an insight into the various aspects of dental care which have benefited from digital technology. 19.2 Digital radiography and magnetic resonance imaging 19.2.1 Intraoral, extraoral, including cone beam computed tomography Without computer imaging of the teeth and underlying bone, dentists and orthodontists cannot adequately assess the condition of the mouth. This could mean that a subgingival tissue could be missed and treatment may be planned without addressing this hidden problem. Noninvasive radiography can provide more detailed images of the structures inside of the mouth, rather than a dentist merely probing and examining the teeth and soft tissue by eye. In this section, a variety of dental imaging techniques will be discussed, later followed by clinical applications and limitations of these techniques. Digital radiography is a way of obtaining radiographed scans without the use of conventional films. To capture an image of a patient’s teeth using conventional films for a long time requires recording of the image, which increases the dosage of radiation compared to digital imaging (Masri and Driscoll, 2015). The time is needed so that the X-rays can combine with the film electrons, and then chemicals need to be used so that the image can be distinguished (Parks and Williamson, 2002; Ti-Sun et al., 2002). By using digital radiography instead of the conventional films not only is less waste created as the developing chemicals are not required, but also time and money are saved for the dental practices (Masri and Driscoll, 2015). For a digital radiograph, the following equipment is needed: a computer
Digital dentistry 509 with an analog-to-digital converter, a digital interface card, an electronic sensor to react to X-rays, and, finally, an X-ray emitting source. Digital radiography uses electronic pixels sensors instead of conventional films. Digital images require the change from analog photos to digital images to show the picture on a computer screen. This is done by the computer assigning a number to every single shade between white and black; this is based on the voltage present in the analog form, with 0 being black and 255 being white and this then creates a binary image which can be used by dentists (Parks and Williamson, 2002). Intraoral and extraoral determines whether an X-ray sensor is internally or externally placed in the mouth. Intraoral radiography is one of the most frequently used assessment techniques by dentists, due to the high-quality image provided and the ease of adjusting equipment (Vandenberghe et al., 2010). There are two forms of digital radiography, direct and indirect. The direct imaging means that the X-ray picture has been stored digitally straight away, whereas indirect requires analog information to be converted to digital information. Either method can be used for both intraoral and extraoral scanning (Parks and Williamson, 2002). Cone beam computed tomography (CBCT) creates images by having an X-ray source and a detector mounted to moving platforms, these two devices rotate about a pivot, placed appropriately to circle the area that needs to be imaged. The cone beam part of the name CBCT comes from the cone-shaped X-ray beam, which is pointed through a specific location on a patient, depending on the specialists need. This X-ray is then received by the detector opposite the X-ray source (Scarfe and Farman, 2008). MRI is a form of digital imaging that does not involve irradiating a patient, which enables analysis of soft and to some extent hard tissues. MRI works by picking up the energy emitted from protons located in the atoms of water within the body. This energy occurs after a magnetic field is introduced to the biological tissue; the protons line up with the magnetic field, then a radio frequency pulse is applied by the tissue causing the protons to become unaligned from the magnetic field. When this radio frequency is removed, the protons realign once more and release energy. This energy is picked up by the MRI machine, and differentiation between tissues can be shown based on the speed of energy release. Teeth have a higher density than soft tissues; this means water content is less and so proton energy is very quickly emitted, resulting in an unclear image as the MRI cannot pick up the energy fast enough (Idiyatullin et al., 2011). Extraoral MRI is the most regularly used technique because of its ease and practicality. However, intraoral MRI can also be performed by placing wireless coils inside the mouth; this provides a more precise image, which is more focused than common extraoral MRI techniques (Flügge et al., 2016). 19.2.2 Clinical applications Radiography and MRI can produce three-dimensional (3D) images that can aid a dentist in decision-making, especially to assess bone quality and structure before
510 Advanced Dental Biomaterials planning for surgery or implants that improve the outcome for patients. Tables 19.1 and 19.2 give examples of types of digital dental imaging techniques for both intraoral and extraoral (adapted from Vandenberghe et al., 2010; Cleveland Clinic, 2018; WebMD, Boots, 2018). CBCT takes only a short amount of time to scan the maxillofacial area, with the ability to produce a clear image of the bone (Scarfe et al., 2006), allowing dentists to quickly analyze the situation of the patient’s mouth and act there and then if required. The resolution of the images can be improved by increasing the thickness of the image layer/slice taken of the person. CBCT allows cross-sectional images to be acquired; this is useful for surgical needs, as a dentist can gain knowledge of the bone quality and structure locations prior to surgery to prepare appropriately (Vandenberghe et al., 2010), and also allowing the development of implants that suit the need of the shape of the patients’ mouth. CBCT can assess the hard tissue within the mouth. This type of radiograph can be used to show the outcomes associated with periodontal diseases, such as bony attachment loss, compared to computed tomography (CT) radiography which lacks the inability to show the attachment of teeth to the periodontal tissue (Sedentexct.eu, 2012). Table 19.1 Intraoral digital imaging. Type of imaging technique Image location Applications Occlusal Can provide an image of the entire maxillary arch or mandibular arch Bitewing Allows imaging of upper and lower teeth, this type of X-ray shows the tooth from the crown down to the top of the bone holding the tooth Periapical Gives an image of the whole length of the teeth on the top or bottom of the mouth, displaying the crown all the way to the root attachment to the bone Aids orthodontists with the progression of teeth movement and development to correct alignment via the use of braces As the bottom and top teeth can both be seen, areas between the teeth can be viewed with ease. Shows bone and teeth relation, aiding detection of caries, periodontal disease, and quality of the restorations. Also, it can support dentists with correctly fitting replacement restorations such as crowns Commonly used to show any abnormalities of root structures or nearby bone changes, good for producing images with reasonable resolution
Table 19.2 Extraoral digital imaging. Type of imaging technique Image location Applications Tomography X-ray tube and plate move in opposite directions, resulting in a clear focused point. One layer of an image with other structures blurred Panoramic Gives an image of the whole mouth including upper and lower jaws CT Shows a 3D image of the mouth, made up of layers of 2D scans, from a triangle-shaped ray CBCT Rotates around the patients’ head, creating a 3D picture of both hard and soft tissues Digital imaging Creates a 2D image of a chosen area on a screen rather than using film MRI imaging Provides a 3D image of the desired location of the oral and maxillofacial area Sialogram A dye called radiopaque contrast agent is injected into the salivary gland, this allows the X-rays to pick up soft tissue structures that would not normally be picked up Gives an image of the left or right side of the head Surrounding structures of an area that needs to be scanned can be blocked out so that the practitioner is able to just see a focused image of the structures they what to analyze and ignore the others Panoramic X-rays can be used to help in tumor diagnosis. However, it is more frequently used to show teeth positioning, whether they are visible, still submerged, or whether they are impacted Can be used to find problems in the bone or face. Enables discovery of fractures of the bone, and cysts or tumor findings CBCT for dentistry has many applications, of which will be discussed further in Section 19.2.2 Creates very quick images so can allow quick diagnosis, also can be zoomed in on giving a dentist better clarity There are many applications for MRI imaging in dentistry, Section 19.2.2 will go through this in detail Allows dentists to uncover any underlying salivary problems, such as duct blockages Cephalometric projections Used in orthodontic applications to examine teeth locations concerning the patient’s specific features, allowing them to make the best judgment for treatment to correct teeth positioning 2D, Two-dimensional; 3D, three-dimensional; CBCT, cone beam computed tomography; MRI, magnetic resonance imaging.
512 Advanced Dental Biomaterials A study took place involving the use of dry skulls, which had periodontal defects created artificially imposed on a variety of teeth, using wax to act as soft tissue (Bagis et al., 2015). All of the skulls had a bite block between the teeth and were scanned by CBCT equipment and an intraoral X-ray. Experienced dentists were then asked to identify the defects present, using the images created by both the intraoral and CBCT techniques. Bagis et al. concluded that CBCT allowed the most accurate diagnosis to be made, this was due to the clarity of the image and the sensitivity of the equipment compared to the intraoral X-ray techniques used. Commonly, endodontic imaging uses intraoral radiography techniques at different stages throughout, from a patient diagnosis, during the correction procedure, and after the procedure to check a patients’ situation. CBCT can be used extraorally, and operating microscopes could be used to reduce radiation dose (Sedentexct.eu, 2012). Temporomandibular joint disorders can be seen clearly using MRI, producing a more transparent image than the other radiography methods. The MRI can also show the alteration of disk placement within the joint (Niraj, 2016). Due to MRI enabling both hard and soft tissue to be viewed via imaging, a more in-depth diagnosis can be made as a wider variety of structures can be evaluated. Although standard MRI has been shown to reduce the resolution of hard tissue structures, Flügge et al. showed that using an intraoral coil within the mouth allowed hard tissues to appear clearer and would enable dentists to examine the structures accurately to make a diagnosis. In abnormality investigations, MRI can be used to carefully allow a tumor, within the salivary gland, to be identified as either benign or malignant so further measurement can be taken (Niraj, 2016). This is useful as the method is completely noninvasive and gives extremely accurate results. Development of MRI techniques is further improved, using methods such as Sweep Imaging Fourier Transformations, such developments aim to improve the quality of bone resolution and increase the availability of these devices to reduce the amount of patient radiation (Idiyatullin et al., 2011). Advancements of MRI are promising to reduce the amount of ionizing radiation that dental patients are exposed to, which is beneficial as dental applications apply more radiation dosage onto people than any other medical application. Despite the possible inaccuracy of hard tissue, MRI sometimes can be the only option for dental diagnosis, for example, in some stages of pregnancy where ionizing radiation should avoided due to the ill effects that it may have on an unborn child. It is also notable that MRI is used in caries detection, which will be discussed further in its own section later. 19.2.3 Limitations The radiation dose on a patient can be a limitation itself, as the risks implied from uncontrolled radiation may be worse than the dental condition itself. Digital technology has reduced the amount of patient radiation. Table 19.3 shows the approximate range of radiation dose that can be present from the different types of digital imaging techniques used in dentistry. This data is only a rough estimate of the ranges, these ranges can depend on the equipment/system brands and can be
Digital dentistry 513 Table 19.3 The radiation dose received for different types of dental imaging techniques. Type of digital imaging technique used in dental application Approximate radiation dose received by patient (Sieverts, Sv) Intraoral digital imaging Bitewing imaging Entire mouth imaging ,1.5 µSv (Sedentexct.eu, 2012) 5 µSv, with a skin dose of 26 µSv (Ritenour and Gibbs, 2010) 34.9 170.7 µSv, with skin dose of B107 µSv and thyroid dose of B330 µSv (Ritenour and Gibbs, 2010) 29 477 µSv (Scarfe and Farman, 2008) and can reach 1073 µSv if the field of view is increased (Sedentexct.eu, 2012) 2.7 24 µSv (Sedentexct.eu, 2012) with a thyroid dose of B42 µSv and skin dose of B5 µSv (Ritenour and Gibbs, 2010) 0 Sv CBCT Panoramic radiography MRI CBCT, Cone beam computed tomography; MRI, magnetic resonance imaging. different based on the area in which the X-ray is applied to, as tissues of varying densities can undergo less or more radiation than other areas (Sedentexct.eu, 2012). The two main units for radiation dose that are commonly used are Gray (Gy) and Sieverts (Sv), Gray is the unit given to state the skin or organ absorption of radiation which is measured in joules per kg, whereas Sieverts is the unit used to measure equivalent dosage of various radiation sources (e.g., gamma or beta) within the body (Masri and Driscoll, 2015). The 2007 recommendations of the International Commission on Radiological Protection (from here on abbreviated to 2007RICRP) states that there is a division between the types of the effects that radiation can have on a person; this can be either a deterministic effect or a stochastic effect. A deterministic effect is only when a radiation level has been exceeded will cell death or fault occur, whereas stochastic effects are from the damage to genetic material commonly resulting in impacts on cell mutation. In both types of these effects, insignificant results are shown for radiation below 100 mSv; under this value 2007RICRP does not believe that health risks can occur to adults. Above 100 mSv there has been evidence to show that the risk of carcinogenesis increases linearly to that of radiation increase. Although to be on the cautious side, 2007RICRP recommended that over a 5-year period a maximum of 20 mSv is the most a patient should be exposed to each year and that one single dose should be no more than 50 mSv at a time (2007). There are a variety of calculations that can be done to find the maximum dose of radiation a person should undergo, this takes into consideration variables such as age, gender, weight, health condition, and many more factors such as genetic conditions, working out an individual’s risk when exposed to radiation (2007RICRP). It should be noted that younger people (under 10 years old) have more radiosensitive tissues and can be three times as susceptible to radiation consequences than people over 80 years old (Sedentexct.eu, 2012).
514 Advanced Dental Biomaterials Due to the possible outcomes of too much radiation, it is sensible for dental professionals to limit and reduce the number of exposures patients received from digital radiography. Some methods are used to reduce the amount of radiation exposure, the system itself and radiographers can perform these measures. X-ray beams can turn on and off to work in time to the capturing from the detector, this reduces radiation to a patient rather than an unnecessary constant beam, while maintaining image quality (Scarfe and Farman, 2008). FDI World Dental Federation also suggests several other methods for exposure reduction. These include ensuring the beam is of the smallest possible diameter and accurately setting the alignment up to get the correct area, taking care on the first exposure to reduce the likelihood of necessary retakes, and wearing leaded aprons and thyroid collars to minimize tissue exposure in other areas. On top of that, the radiographer should take their safety into consideration too, they should either leave the room and remotely switch on the X-ray, or stand behind a protective screen (FDI World Dental Federation, 2018). Radiograph operators can also wear a badge that shows how much radiation they have been exposed to, so they can stay within the safe limit when working with X-rays. Comparisons of radiographic measurement techniques to the intrasurgical assessment of patients with periodontal disease have taken place to analyze the difference in the interproximal bone loss. One study compared linear measurement (LMSRT) and FRIACOM (a type of computer-assisted analysis device to compare radiographic images) as their computer-assisted analysis devices against intrasurgical assessment (Kim et al., 2002). Using statistical methods to find the difference between these assessments, they showed that the computer-assisted analysis methods overestimated bone loss when calculating the distance between the cementoenamel junction to the alveolar crest, by approximately 1.5 2 mm. On the other hand, LMSRT and FRIACOM appeared to be reliable and produce the same results as each other. Another study compared LMSRT to a loupe, in this case, these two techniques underestimated interproximal bone loss in comparison with intrasurgical methods (Eickholz et al., 1998). Digital X-ray systems utilize different types of sensors including charged-coupled device (CCD), photostimulable phosphor plates (PSP), or complementary metal oxide semiconductor active pixel sensor (CMOSAPS). These sensors used for direct digital imaging have various disadvantages and limitations. CMOS sensors can only create an image over a small area and have noise that hinders the quality, therefore reducing diagnosis ease. Hygiene is vital for CCD usage as otherwise infections could spread between patients. This requires the sensors to be covered with a clean protective material, and for every patient, this cover needs to be replaced. CCDs can be uncomfortable when used intraorally and because of this more X-rays may need to be taken due to patient movement (Parks and Williamson, 2002). PSP can be made to a variety of sizes meaning the patient’s comfort can be improved, and fewer retakes are needed (Vandenberghe et al., 2010), PSP is an indirect system however and so can result in information lost when converting from analog-to-digital images. Intraoral digital radiography sometimes can, unfortunately, result in an unclear picture; this is from the appearance of overlapping structures within the mouth in
Digital dentistry 515 the form of a two-dimensional (2D) image. Similar to intraoral digital radiography, extraoral 2D photos are commonly overlapping, blurry, and can be misleading due to the magnification differences present (Vandenberghe et al., 2010). CT and CBCT both cannot display soft tissue. Clearly, this means the investigation of issues with soft tissues such as dental pulp, mucosa, and the neurovascular bundle cannot be explored with clarity (Flügge et al., 2016). Image intensifier tubes (IIT) and CCDs can be used in unison for CBCT imaging, or flat-panel imagers can be used for detectors, these are indirect imaging methods. IITs compared to flat-panel detectors can produce falsifications, in turn causing an unreliable result. Flat-panel detectors can sometimes produce incorrect pixelations; however, this can be easily corrected by these pixels being replaced digitally to appear as a mean pixel based on the surrounding pixels (Scarfe and Farman, 2008). Despite CBCT having an improved image quality compared to intraoral X-rays, CBCT should not replace the use of intraoral X-rays, this is because CBCT inflicts a higher dose of radiation onto patients (Bagis et al., 2015). Therefore CBCT should only be used in cases where intraoral X-rays are inadequate or where the risk of radiation dose outweighs the risk of dental problems. Also, if a dentist wishes to analyze soft tissue, CBCT is an unwise choice as soft tissue image quality is low due to noise so that the only bone can be picked up clearly (Scarfe and Farman, 2008). The areas that sustain the highest dose of radiation, when using CBCT for a dental application, are the salivary glands and thyroid (Sedentexct.eu, 2012), this definitely indicates that it would be sensible to distribute a protective thyroid collar to patients undergoing CBCTs. MRI is less commonly used in dentistry; this is due to the cost and size of MRI machines, also scanning the patient can take a long time. Image quality for hard tissues is not as good as that of soft tissues, hindering the usability of MRI, as often bone and enamel examinations need to be determined more importantly (Flügge et al., 2016). It should also be noted that patients may be claustrophobic and placing them in a confined MRI machine may cause them to move; this movement may distort the imaging and require longer imaging time (Niraj, 2016). Although MRI does not emit damaging radiation, some very powerfull types of MRI can create such strong magnetic filed that moves magnetizable objectes uncontroablly towards the machine within the same room. Unfortunately, this means that patients with metal implants cannot have this type of MRI, otherwise these implants may be dislodged (Magnetic Resonance Imaging (MRI) National Institute of Biomedical Imaging and Bioengineering, 2018). Commonly, dental fixtures are made of magnetic materials and need to be assessed before an MRI can take place; examples of these restorations that might cause health concerns if used in MRI include orthodontic braces, crowns, and dentures (Mathew et al., 2013). In particular, it is necessary to emphasize the importance of undergraduate dental students knowing what types of implants are magnetic, so that they can avoid health risks, which could be as severe as a ferromagnetic implant being pulled out of the body by the magnetic field.
516 19.3 Advanced Dental Biomaterials Caries detection Caries is the most common chronic disease in the world, and its diagnosis forms an essential skill the dental healthcare professional must master. However, their early reliable and predictable diagnosis remains one of the most challenging conditions in dentistry. Traditional methods of caries detection involve visual and tactile examination of the tooth surface with the use of air drying and probes, respectively. Radiographs using the bitewing technique are most commonplace in practice to identify and monitor carious lesions of a tooth, as well as help assess any existing restorations and surrounding bone levels, all of which cannot be detected solely by visual inspection. Automated computer diagnosis of dental radiographic imaging has been of limited success; the image analysis is a 2D representation of a 3D tissue; it, therefore, requires an evaluation by the practitioner to then achieve a computer-aided diagnosis (Masri and Driscoll, 2015). Nevertheless, the accuracy with which dental caries is identified remains to be significantly improved, leading to the market demand for a technology which can successfully detect caries. The introduction of the International Caries Assessment and Detection System (ICADS) over 10 years ago highlighted the importance of having a universal framework for grading caries within a tooth and has subsequently allowed for comparisons on the performance of different technologies. Light fluorescence technology has undergone much research and development; it involves measuring the difference in the amount of light reflected off a sound and carious tooth surface. A significant drawback is that any plaque or staining could distort the result and therefore it is necessary to ensure the tooth is cleaned thoroughly before assessment (Gomez, 2015). A further variation of this technology is to direct infrared light at the tooth surface and use a digital camera to measure the amount of light which passes through (healthy enamel) or scatters in all directions because of caries or demineralization present. 19.4 Photography and shade selection Advances in computer, camera, and Internet technology have hugely changed modern society. With an ever-increasing demand for improved esthetics in dentistry, numerous technologies have been developed to improve shade analysis, laboratory communication, dental material selection, as well as reproduction, and, finally, verification in the clinical setting (Chu et al., 2010). Successful shade matching is both art and science combined with any mistakes potentially being costly. It is an important yet difficult task being a matter of color perception and subjective interpretation. Objectively, there are three components to shade: “hue” is the color, “chroma” is the intensity of the hue, and “value” is the brightness of the hue. The most commonly used shade-matching technique is a visual method using a Vitapan Chart system (Chu et al., 2010). This involves the dentist holding a series of colored tabs of distinctive shades next to the patient’s teeth. However, there are strict conditions to
Digital dentistry 517 try and improve the accuracy of color perception; the patient must be sat upright and assessed at eye level at a conversational distance, nonfluorescent color corrected light should be used, and the teeth kept hydrated. Spectrophotometer technologies available on the market offer a greater accuracy with color matching over the human eye (Chu et al., 2010). They work by measuring the quality of light reflected off the tooth’s surface. The main limitation, however, is that the software will average out the color data received for a given area leading to inaccuracies and major errors where the tooth exhibits translucency, especially around the tooth’s edges (Chu et al., 2010; Tam and Lee, 2017). Digital camera and imaging systems are prevalent devices from single-lens reflex cameras to smartphones. They offer an alternative to color matching and measurement instruments with the significant advantage of communicating color distribution and dental morphology (Tam and Lee, 2012). Computer software can analyze images and compare the shade of the teeth with reference shade tabs adjacent to the teeth in the oral cavity. The use of consumer smartphone cameras allows for easier and immediate wireless communication with the laboratory to fabricate dental prostheses which not only match in color but also compliment the natural dentition in shape and design (Tam and Lee, 2012, 2017). Furthermore, the use of camera technology alongside a shade reference is playing an increasing part in monitoring those undergoing tooth whitening (Chu et al., 2010). 19.5 Computer-aided design computer-aided manufacturing systems in dentistry Dr. Duret in 1971 was one of the pioneers to employ an optical impression system linked to a crown designing and milling system (Miyazaki et al., 2009), which was later developed into the current computer-aided design computer-aided manufacturing (CAD CAM) system. A CAD CAM system is generally composed of three stages of data acquisition, data processing, and machining (Qualtrough and Piddock, 1995). A 3D image of the prepared tooth can be either recorded directly using an intraoral scanner or indirectly by scanning either an impression of tooth replicas, cast, or die. A virtual restoration or coping is designed on the virtual replica using dedicated software. Some CAD CAM systems are equipped with a virtual articulator to help to develop the occlusal surface of the restoration accurately. The final design is transferred to a milling machine which carves out a coping or a fully shaped restoration (monolithic restoration) from a high strength ceramic block. Machining the restoration, which is referred to as “subtractive manufacturing,” can take place either chairside or in a laboratory or an industrial milling center. Later the ceramic cores are veneered either by sintering ceramic powders or pressing softened ceramic onto. The CAD CAM technology has been primarily used to fabricate fixed prosthetic restorations, such as onlays, inlays, veneers, and crowns. Further advancements on CAD CAM during the past decade have provided alternative restorative
518 Advanced Dental Biomaterials materials such as composite resin, porcelain, and metallic blocks, all of which have improved the quality of prostheses (Raigrodski, 2004). The CAD CAM system can now be used for fabrication of implant-supported prostheses and implant abutments (Kucey and Fraser, 2000). The CAD CAM systems have three major functional components (Strub et al., 2006): 1. Data gathering or scanning to receive the oral information. To receive this information there are various trading systems: a. Intraoral capture. This approach utilizes 3D optical systems to obtain single components anatomy. Some examples include Interférométrie Moire, laser scan, color coding (e.g., CEREC 79 and Evolution 4D). b. Anatomical dental duplicate capture (plaster cast), often utilizing a laser scan system. Some commercial products include RapidForm (RapidForm), Slim (Slim), poly-Works (polyWorks), and Geometric Studio (Geometric Studio), which are used for the 3D meshes following the process. 2. The CAD technique for geometric design of the restoration. Such CAD systems have their own basic functions to modify the geometry of the restoration. 3. The CAM system to fabricate the restoration. The CAM systems utilize information gathered from the computer to outline a physical object, using subtraction methods which remove the material from a starting block and finalize it to the desired shape, or using additive methods (rapid prototyping) that are increasingly used in the CAD CAM oral technology. Since the introduction of CAD CAM technology in the early 1980s, further advancements have progressed in three directions, depending on the type of production line: (1) chairside production, for example, Cerec System (Sirona Dental GmbH; Salzburg, Austria); (2) laboratory production, for example, inEos X5 scanner and inLab MC XL milling unit (Sirona Dental GmbH; Salzburg, Austria); and (3) centralized fabrication in a production center, for example, Nobel Procera (Nobel Biocare, Zürich, Switzerland). 19.5.1 Chairside milling The Cerec system was the first commercially successful chairside CAD CAM system equipped with an intraoral scanner, a computer for designing virtual restoration, and a chairside milling machine. This system, which was introduced in 1987 by Sirona, initially could produce ceramic inlays. However, further development of this system resulted in the production of a range of ceramic and metal restorations. Joining Dentsply and Sirona provided a new opportunity to develop the system even more. 19.5.2 Laboratory and industrial milling Chairside scan and design data or conventionally made impressions can be sent to a laboratory for machining. Cerec and Planmeca Planmill are two popular CAD CAM systems used by UK dental labs for milling the restorations out of
Digital dentistry 519 various ceramic blocks and disks. Also, the data could be sent to an industrial milling center, which may be in another country, to design and mill the restorations or copings and deliver them within a few days. Straumann, Lava 3M, and Nobel Biocare are examples of industrial machining centers. These systems are classified as “closed systems” because they only serve their own software processed data, whereas there are other machining centers such as Zirkonzahn that use “open systems” which are compatible with the main manufacturer’s scanners, software, and materials. 19.5.3 Machining of the restorations Final restorations are carved out of presintered block or disks of materials using a range of burs set by operating computer to work in four or five different axes. The burs are selected to mill the fitting and external surface of single or multiple restorations in one run. Multiple axes help in cutting occlusal and fitting surface detail with more accuracy. Nowadays, various materials including alloys, ceramics, composites, resins, and waxes are available to be used by CAD CAM machines. Some of these materials are too hard to mill, therefore to facilitate the milling and reduce wear and tear of the milling machines and its burs; they are milled in partially sintered forms. After this milling technique that is referred to as “green machining,” the fully formed restorations are sintered to achieve optimum strength (Kelly and Benetti, 2011). Milled resin and wax copings or partial denture frameworks are converted to the restorations through conventional techniques. The early developed single-shaded ceramic blocks barely simulated the natural tooth color, whereas newly developed multilayered ones offer a better color match. However, to achieve a good result, operator experience in the correct orientation of the virtual design on the block is important. The diameter of the milling bur is a restricting factor in the amount of the detail that can be produced by the milling systems. Also, the length of the milling burs limits the depth of the fitting surface of the crown and the height of the preparation as a result (CEREC, 2016). Generally, it is advised to round the line angles and avoid preparation of long and parallel axial walls. 19.5.4 Three-dimensional printing Several industry subsectors including automotive, aerospace, electronics, machinery, and medical products have utilized additive manufacturing (AM) systems known as 3D printing (3D Printing, 2012). 3D printing, also known as rapid prototyping, is a technology in which objects are fabricated layer by layer at a time. Layers are added to previous ones until a duplicate of an object is formed (Fig. 19.1). The technologies associated with the 3D printing have been used before. However, the “3D printing” concept is quite new and has captured the public’s attention. Mechanically speaking, 3D printers have a simple design. The apparatus could not operate without the presence of the CAD software. The CAD software allows any object to be virtually reassembled. The CAD is commonly
520 Advanced Dental Biomaterials Figure 19.1 3D printed fixed metal restorations, copings, and frameworks on one plate. used in industrial design, manufacturing environments, engineering, and dental laboratories. Some aspects of various dental surgeries can also benefit from this software as well. 3D printing is used to duplicate objects. The CAD software helps us print objects from scratch. In dentistry and surgery, CT and CBCT provide detailed volumetric data as well as intraoral or laboratory optical surface scan data. Recent advancements in CBCT and optical scan technology have significantly improved various aspects of restorative and implant dentistry. Fabrication of anatomical “study models” can be one of the primary applications of 3D printing in the fields of surgery and medical modeling. CBCT access in dental practice has significantly increased over the years and has significantly improved diagnosis and treatment planning in implant dentistry. CT and CBCT provide us with volumetric data that is transferred to a 3D printer before the surgery to duplicate a patient’s jaws precisely. A reproduction of the exact anatomy in detail, particularly in complex, atypical, or unfamiliar cases, can be carefully reviewed beforehand and a suitable surgical approach can be planned accordingly. Fabrication of implant drill guides necessitates the access to precise 3D printers and high-resolution printing materials. Unfortunately, some of the ideal materials that should be incorporated in the fabrication of these drills are not autoclavable. The traditional precious metal casting has been substituted for more modern materials that are utilized in CAD CAM technology (Bammani et al., 2013). Advantages associated with CAD CAM include easier manipulation of materials and elimination of labor-intensive artisanal production techniques (Venkatesh and Nandini, 2013). It also enables mass production of restorations and offers more time to dental technicians to focus on more creative aspects of the fabrication process, including esthetic layering of porcelain. Incremental layers of polymers, ceramics, or metal alloys are laid down and fused together using a high-powered laser that is called selective laser sintering. The term direct metal layer sintering (DMLS) is used when a laser is used to fuse metal alloy powders. Studies show that Co Cr restorations produced by either milling or SLM systems are less porous microstructurally and more homogenous than the conventionally casted ones. Also, internal and marginal accuracy of SLM-made restorations are either equal or better than casted
Digital dentistry 521 restorations, whereas milled restorations were the worst. However, the clinical significance of the differences is not clear (Koutsoukis et al., 2015). Several restorations can be made in one run in the same plate, which saves manufacturing time remarkably (Fig. 19.1). However, the initial cost of 3D laser printing systems are still more than that of milling systems and the production of metal fumes are potentially more hazardous to health than grinding debris (Dawood et al., 2015). Some other challenges include the need for meticulous cleaning and difficult postaccessing. 19.6 Computer-supported implant dentistry New advances in implant technology have allowed the application of CAD CAM to improve implant treatment further. In complex cases where the application of standard abutments is not feasible, CAD CAM technology can be utilized to produce customized implant abutments. Also, customized copings have also been manufactured in such cases to facilitate taking more accurate impressions (Priest, 2005). Recent advancements in 3D printing and CAD CAM make them suitable means for making various components of prostheses and planning different phases of implantation. CBCT can be used in conjunction with CAD CAM technology in fabricating surgical guides for implant placement. Although it is predicted that 3D printing technology will eventually produce customized implants with analogs that closely resemble the roots of the missing teeth, dental implants themselves are still manufactured using conventional techniques. 19.6.1 Three-dimensional printing in implant dentistry Fabrication of dental implants is a multistep process including machining of titanium rods, followed by modification of implant surface design through sandblasting, acidetching, anodization (Choi et al., 2012; Degidi et al., 2012), discrete calcium phosphate crystal deposition (Li et al., 2015), and chemical modifications (Monjo et al., 2012; Choi et al., 2012; Elias et al., 2012). The goal is to improve implant stability and increase osseointegration. Clinical studies have demonstrated excellent longterm success or survival rates for rough surface implants (Mangano et al., 2014a; Degidi et al., 2012; Sesma et al., 2012). The application of conventional methods in implant dentistry does not allow the construction of a precisely controlled exterior and interior porous pattern. To address these shortcomings, AM methods have been proposed (Mullen et al., 2009). DMLS is a laser-based AM technology that employs powdered metals, radiant heaters, and a computer-controlled laser machine to construct an object by adding layers of material together. This technology uses CAD data to construct the object without the need for further tooling costs or inventories. No cutting or milling procedures are involved in this technique, which makes DMLS an efficient technology with less waste and almost zero loss of material (Mangano et al., 2009; Dabrowski et al., 2010). There is insufficient evidence on DMLS
522 Advanced Dental Biomaterials titanium implants within the current literature, even though the AM and DMLS concepts pertaining to implant and biomaterial manufacturing are well accepted. The application of DMLS technology allows fabrication of patient-specific implants. DMLS can be used to make custom-made titanium implants, such as root analogs (Moin et al., 2013; Mangano et al., 2014b) or blade implants (Mangano et al., 2013), causing more adaptation of the implant to the patient’s anatomy compared with preformed standardized fixtures. 3D printing offers the ability to produce batches of dental implants of complex geometries highly adaptive to the patient’s bone. Milling alone is not capable of producing such products; however, the milling/machining procedures can refine the printed object such as the implant platform. It is possible to make implants of complex geometry. However, the ultimate insertion of dental implants with the aid of screws is still the favorable approach. Although the 3D printing technology concepts are straightforward, the postprocessing concepts are not. There are significant health and safety challenges associated with the fine metal powders and nanoparticle wastes created with 3D printing. The 3D printers are readily available in many dental laboratories and the associated postproduction equipment and accessories do not take up much space. Although the use of 3D printers in the fabrication of any material seems logically possible, the maintenance of the machine, particularly switching between an implantable metal and a restorative material, is extremely challenging. 19.6.2 Recent advances in implant technologies The implant dentistry is nowadays computerized due to the introduction of 3D imaging, implant-planning software, CAD CAM technology, computer-guided (CGIS), and computer-navigated implant surgery (CNIS). Implant surgery treatment planning via 3D computer software enables proper preoperative evaluation of anatomic limitations. In addition, CAD CAM technology allows preoperative planning of implant positions along with virtual implant placement and accomplishment of the surgical phase through static (guided) or dynamic (navigated) systems. Presurgical 2D radiograph imaging has traditionally been used to determine implant position, size, number, direction, and placement of the implant. The guiding acrylic stents were prepared over duplicated casts of diagnostic wax-up. Limitations associated with 2D radiography as well as inaccuracies in the fabrication of the stent or guide channels led to improper implant placement, which may cause complications and eventual implant failures, particularly in anatomically complicated cases. 3D CT imaging allows scanning of the patient either with fiducial (artificial) radiographic markers embedded in stents, jaws, etc. or with anatomic (natural) markers that are teeth or bony landmarks. The digital image is taken and imported into one of the available implant-planning software programs. The data are converted into a virtual 3D model that replicates the patient’s bone anatomy, making the execution of the surgery and prosthetic treatments accurate. An undistorted 3D view of the jawbone in axial, sagittal, coronal, panoramic, and cross-sectional views is possible by means of implant-planning software. The software also creates a 3D
Digital dentistry 523 reformatted reconstruction of the jaw. Some additional advantages associated with this software are as follows (Ganeles et al., 2011; Rubio Serrano et al., 2008): G G G G digital planning and fabrication of virtual wax-ups, implant positioning, abutment designing, surgical guidance, fabrication of provisional and final restorations; predetermining size of the implant, the abutment, and the provisional restoration; avoiding possible complications due to inaccuracies in the selection of implant size or positioning during virtual planning by appropriate use of the software; and helping with treatment planning on dental procedures, for example, alveolectomy, alveoplasty, implant positioning in anatomically challenging cases, visualization of bone quantity in each location and assisting in choosing the ideal donor site for osseous grafts, graft location, volume and shape of the graft, sinus lift procedures and placement of implants in a single step, atrophic maxillae treatment, and placement of trans-zygomatic implants. Even though there are many advantages associated with the use of these software programs, there are some drawbacks associated with these programs, including the time required to learn to operate these programs, the high cost associated with investment in these programs, and the need for accurately localizing natural or fiducial markers in image data and on cases. There are many commercial software programs available on the market. Some examples include the following (D’souza and Aras, 2012): G G G G G G G G G G G Procera-Software (Nobel Biocare, Göteborg, Sweden) coDiagnostiX (IVS Solutions AG, Chemnitz, Germany) Easy Guide (Keystone-Dental, Burlington, Massachusetts, United States) SICAT (SICAT GmbH and Co. KG, Brunnenallee, Bonn, Germany) Virtual Implant Planning (BioHorizons, Birmingham, Alabama, United States) ImplantMaster TM (I-Dent Imaging Ltd., Hod Hasharon, Israel) Simplant, SurgiCase (Materialize Inc., Leuven, Belgium) Implant3D Media Lab Software [Media Lab Srl, Follo (SP), Italy] DentalSlice (Bioparts, Brazil) Scan2Guide or S2G (iDent, Ft. Lauderdale, Florida) Tx Studio software (i-CAT, Imaging Sciences International, Hatfield, Pennsylvania) In their paper Martins and Lederman (2013) investigated the efficacy of virtual planning by DentalSlice software and reported the software as being efficient on treatment planning prototype guides for implant positioning and for quantifying and locating bone grafts, hence assuring overall higher success rates. Nkenke et al. (2012) included virtual dental implant-planning software in their undergraduate curriculum after receiving positive feedback from dental students. By means of the CAD CAM technology, turning the virtual treatment plan into actual patient treatment is possible. This possibility has been applied in two guided surgery systems that are either static (template-based system) or dynamic (surgical navigation/computer-aided navigation). Some advantages associated with CAD CAM technology are as follows: G G Minimally invasive surgical procedures with surgical guides (CGIS) for implant placement are better executed with improved predictability. Immediate loading of the implant is possible by fabricating presurgical master cast and accurately fitting custom-designed restorations (Spector, 2008; Fortin et al., 2006).
524 Advanced Dental Biomaterials 19.6.3 Computer-guided implant surgery The static system which takes advantage of a static surgical template/guide to reproduce virtual implant positioning in the surgical field is divided into two types depending on the CAD CAM technology used for fabrication of the surgical guide: 1. static system using rapid prototyping technology and 2. static system using computer-driven drilling technology Some benefits associated with the application of CGIS and surgical template include the following (Spector, 2008; Fortin et al., 2006; Nikzad and Azari, 2008): G G G G G G Directs the osteotomy drills precisely to the location. Guides the surgeon to the exact location and angulation to place the implant based on the virtual treatment plan. Allows flapless surgery, causing less bleeding and swelling, reduced healing time and postoperative pain. Promotes the preservation of both hard and soft tissues and maintains sufficient blood flow to the surgical site. Avoids disturbing vital structures. Reduces surgical time. Even though the CGIS technology is very accurate, still there are some disadvantages associated with it that need further attention. Some drawbacks are as follows (Neugebauer et al., 2010; Valente et al., 2009): G G G G G G G G G G inaccuracies in data acquisition or errors in image processing; possible deviations from previously planned implant positions/angulation particularly in coronal and apical portions of the implants; probable displacement during perforation due to inaccurate fixation of the guide; mechanical faults caused due to inappropriate angulation of the drills during bone perforation; the altered positioning of surgical instruments caused by limited mouth opening; possible surgical guide fracture; complicated cases causing unexpected errors; high costs associated with tools needed such as software programs and surgical templates; the possibility of thermal injury during placement of implant by using a flapless surgical approach and application of surgical guides due to reduced access for external irrigation while preparing the osteotomy site; and inability in the intraoperative modification of implant positioning. Due to these drawbacks, the application of this technique needs further caution and care. However, due to improvements with CNIS many of the mentioned disadvantages associated with CGIS can be ignored. 19.6.4 Computer-navigated implant surgery CNIS employs a surgical navigation dynamic system that uses CT data to reproduce the virtual implant position. This is accomplished by an optical bur tracking system
Digital dentistry 525 without a need for an intraoral surgical guide. There are many navigations or positional tracking systems. In CNIS, the natural (fiducial) markers/reference points are required for registration of the instruments (Neugebauer et al., 2010). Watzinger et al. (1999) discussed in their case report the application of an optical tracking system for intraoperatively transferring preoperative planning into CT scans as the implant drill motor caused distortion of the magnetic field and obstructed direct visualization of the implant socket drilling during application of the electromagnetic tracking system. The 3D positional data are transferred to a camera or detector through sensors that are attached to both the patient and the surgical handpiece. These data make it possible for the computer to instantaneously calculate the virtual position of instruments relative to the image data. They also allow visualization of instrument movements in real time to the surgeon through side viewers or advanced see-through viewers (Winter et al., 2005). Some advantages associated with computer-CNIS over CGIS include the following (Jung et al., 2009): G G G G Permitting intraoperative alterations in implant positioning so that the virtual surgical plan could be modified during surgery and helping the surgeon navigate the system and visualize the patient’s anatomy at the same time. This allows the clinician to address obstacles and defects that were not diagnosed preoperatively. Allowing the drill to be tracked and continuously visualized on the computer screen three dimensionally (x, y, and z). Not suffering the limitations associated with CGIS application, that is, thermal injury, displacement, or fracture of the guide. The image-guided implantology (IGI), a subcategory of the CNIS system, is highly accurate in regard to navigation with an acceptable overall mean spatial navigation error of 0.35 mm (Casap et al., 2004). Elian et al. have demonstrated the high accuracy of implant placement using the CNIS system (IGI) by reporting a mean linear accuracy of less than 1 mm at both the implant neck and apical tip levels. They also showed the mean angular deviation of less than 4 degrees for the implants placed through the CNIS system. In fact, there was a precise match between the planned implant and the final implant (Elian et al., 2008). However, a mean linear accuracy ranging between 1.1 and 1.45 mm at the implant coronal third and between 1.41 and 2.99 mm at the implant apical third along with a mean angular deviation ranging between 2 and 7.25 degrees within implants placed using stereolithographic guides has been illustrated (Ozan et al., 2009; Di Giacomo et al., 2005). Considering all previous researches on this subject, it can be assumed that the CNIS system is more accurate than the CGIS system. One study, however, evaluating the accuracy of optical tracking versus stereolithographic system for implant placement found no significant difference between the two systems (J et al., 2008). Research conducted in 2014 compared the accuracy of a dynamic CNIS system with three commercial CGIS static systems and the application of an acrylic stent for implant osteotomy preparation. The results illustrated that the static and dynamic systems offered higher accuracy compared to laboratory-made acrylic guides and that both static and dynamic systems had an average error of less than 2 mm and 5 degrees (Somogyi-Ganss et al., 2015).
526 Advanced Dental Biomaterials Although the CNIS technology takes advantage of optical tracking systems and is highly accurate, it still has some limitations (Birkfellner et al., 2008; Widmann and Bale, 2006): G G G G The technology is sensitive to reflections and interference with the line of sight between the sensors and the cameras. This may cause some inconvenience, particularly with the common seating arrangement of the dental professional and the assistant, making tracking of the instruments very challenging. An expensive technology with an expensive hardware. Requiring demanding intraoperative referencing. The significantly high learning curve. The significant benefit of the technique is the ability to position dental implants based upon surgical and prosthetic requirements, making the entire surgical procedure safer, faster, and easier. Also, it reduces the postoperative pain and discomfort of the patient (Greenberg, 2015). Considering all the advantages, the accuracy of the guided surgery and success rates associated with the procedure are still controversial. Guided surgery is the digital workflow upon which the entire implant placement is virtually treatment planned after the consideration of the unique surgical and prosthetic needs of the patient. However, this technique has its drawbacks. The deviation of the implant from a correct position is possible following fabrication of an inaccurate surgical guide. Hence, the guided surgery is contraindicated in cases with limited mouth opening within which wrong positioning of the surgical instruments is likely (Gulati et al., 2015). There are three types of surgical guides according to the tissue type used as support: bone, mucosa, or tooth-supported guides. The bone-supported guide is more precise compared to the other two types because it offers better visualization of the surgical field. The mucosa-supported guide, on the other hand, is the most minimally invasive approach, when performing the flapless surgical technique (Raico Gallardo et al., 2016). With all the abovementioned limitations, when correctly performed, the technique is quite safe and precise (Katsoulis et al., 2009). The survival rate is comparable to the conventional technique (Hultin et al., 2012). Schneider et al. (2009) reported different errors and discrepancies that could occur during treatment planning and surgical procedures. It appears that the technique still requires further improvements to avoid any potential prosthetic complications in the future. Arisan et al. (2013) investigated the deviation of 108 implants that were placed using mucosa-supported surgical guides when CT or CBCT was used for virtual planning. Both CT and CBCT showed comparable deviations (Arisan et al., 2013). Petersson et al. assessed the accuracy of the technique by comparing the virtually planned position of implants with the actual implant positions following the surgical procedure. They concluded that a stricter protocol is required to avoid errors and complications (Pettersson et al., 2010). The precision related to guided surgery depends on the surgical guide’s stability inside a patient’s mouth (Gulati et al., 2015). There is still controversy among studies pertaining to fidelity presented by surgical guides to transfer the implant
Digital dentistry 527 positioning from virtual planning to the actual location. Instability of the surgical guide can cause deviations up to 6 mm on implant positioning, particularly in patients with multiple implants (Widmann and Bale, 2006). The conventional method has been compared against computer-aided implant surgery when bone and mucosa-supported guides have been used (Arisan et al., 2010). The authors considered the duration of the surgery and postoperative complications in their comparison. They reported reduced surgical time and postoperative pain and discomfort following application of the mucosa-supported guides for flapless technique. Vasak et al. (2011) discussed flapless surgery as a reliable technique since any implant deviation from the expected position did not exceed the safety distance initially recommended by the planning software. Any minor movements of the surgical guide during the surgery can mainly lead to instability of the template. Di Giacomo et al. (2012) reported a 34% 43% complications from guided surgeries. The potential for implant deviation is higher in the posterior region, particularly in patients with limited mouth opening, which challenges the clinician in positioning the surgical guide. Fortin et al. (2000) concluded the technique was an effective mean for complex rehabilitation and minimally invasive surgeries. 19.6.5 Computer-aided design computer-aided manufacturing systems in implant restorative dentistry The CAD CAM technology has been applied in restorative dentistry for two decades, and now it has been extended into implant dentistry. This technology was introduced to restorative dentistry by Francois Duret in 1971 (Duret et al., 1988) and the first dental CAD CAM restoration was fabricated in 1983 (Priest, 2005). The CAD CAM can be used to fabricate patient-specific abutments which fit very well and are more stringent as they incorporate elements such as titanium, alumina, and zirconium into the abutment. Stereolithography can be utilized to manufacture CAD CAM surgical templates, which can improve accuracy during implant placement. Stereolithography allows flapless, minimally invasive surgery, accomplishing immediate functional loading to the implant and delivery of the prosthesis. The list of products which can be fabricated by the CAD CAM technique depends on the respective production system. Some milling machines can produce ZrO2 frames, while more sophisticated units can accommodate a variety of materials including metals, glass ceramics, resins, and high-performance ceramics. The glass ceramic blocks are grindable silica materials with high stability values. Lithium disilicate ceramic blocks are critical silica-based materials which can be used for the production of full anatomical crowns and copings in both anterior and posterior locations and for three-unit FPD frameworks in the anterior region because of the high mechanical stability, equivalent to 360 MPa, of this material (Sorensen et al., 1998; Taskonak and Sertgöz, 2006). High-performance oxide ceramics are categorized into two groups: aluminum oxide (Al2O3) and yttriumstabilized zirconium oxide (ZrO2, Y-TZP), a high-performance oxide ceramic that
528 Advanced Dental Biomaterials in comparison with other all-ceramic materials, express better mechanical properties, high flexural strength of 900 1200 MPa, and higher fracture toughness of 6 10 MPa (Komine et al., 2010; Kunii et al., 2007). 19.6.6 Prosthetic abutments For various reasons, the application of CAD CAM technology in implant dentistry is increasing. First, the implant frameworks produced by CAD CAM systems are made out of a solid and homogeneous block of material with high mechanical properties. Second, since no waxing, investing, or casting procedure is involved, the chances of any errors or inaccuracies are minimal. Due to this reason, the overall production costs are reduced. Also, by application of CAD CAM technology, any potential for unfavorable implant angulations is minimized, and an appropriate emergence profile is accomplished. Furthermore, the ceramic abutments fabricated by CAD CAM have optimal optical properties similar to natural dentition with expected highly esthetic results for the surrounding soft tissues. Preferably, the abutment head must be similar to a prepared tooth with good stability, morphology, and emergence profile. The appropriate positioning of the implant and proper manipulation of hard and soft tissues are of importance in producing ideal emergence profile, function, esthetics, and periodontal health. 19.6.7 Computer-aided design computer-aided manufacturing abutments in implant dentistry The custom abutments which have been produced via CAD CAM technology have advantages compared with stock and laboratory processed custom abutments (Priest, 2005). Similar to laboratory-made abutments, CAD CAM-fabricated custom abutments are specific for each patient. It is easier for a technician to learn and employ this technique than that of handmade components. By using CAD software, the laboratory technician has control over abutment design. The virtual abutment design is electronically transferred into a CAM milling machine so that the abutment can be fabricated out of a block of selected material. With this approach, any dimensional inaccuracies related to waxing, investing, and casting are omitted. In contrast to stock or cast custom abutments, the surfaces of abutments fabricated from CAD CAM system do not require further modifications following the machining procedure due to their high precision and fitness. The cost associated with implant abutment manufacturing with CAD CAM technology lies somewhere between the cost of fabrication for a stock and cast abutment. As the use of CAD CAM technology grows over time, the cost will most probably drop. However, the labor and laboratory costs are most likely going to increase. The overall effect would be that the cost of prepared stock abutments or handmade cast custom abutments is going to escalate.
Digital dentistry 529 19.6.8 Materials used Metals such as titanium and titanium alloys as well as ceramics such as aluminum oxide and zirconium oxide have been used in CAD CAM technology to fabricate implant abutments (Sherry et al., 2007). The longevity of the restorations fabricated by CAD CAM has increased with the incorporation of such materials of high strength and quality, making this technology favorable among dentists. Some products include CEREC 3D (Sirona Dental Systems) (CEREC 3D), Everest (Everest), and Lava (LAVA). 19.6.9 Computer-aided design computer-aided manufacturing custom implant abutments Many types of implants employ these forms of abutments (Raigrodski, 2004). The process is initiated by using CAD CAM technology and related software to gather information from the patient’s mouth. The laboratory technician uses the data to wax the prosthesis over the corresponding abutment and then scans it. The antagonist arch is used to adapt the structure and the emergence profile. The collected data is then transported to the CAM center so that the designed abutment is milled, and the ceramic is added (Raigrodski, 2004). Recently, with the exception of the external or internal hex, this method has been extensively used to design and fabricate the implant abutment. The available CAD software programs are supplied with databases which can be used to select the abutment. Another choice is to scan and enter the data into the software to achieve the desired shape and design. The designed outline can then be altered depending on the complexity and uniqueness of the case. The digital data is transported to a milling machine that is attached to a computer. This data is used to mill the abutment from a solid block of titanium alloy. The shape and fit of the abutment are then evaluated on the cast (Schneider et al., 2009). A few examples of the commercially available CAD CAM abutment systems include the following: G G G Cerec (Sirona, Patterson Dental Co., Milwaukee, Wisconsin), Atlantis Abutments (Atlantis Components, Inc., Cam-bridge, Massachusetts), and Encode Restorative System (3i Implant Innovations Inc., Palm Beach Gardens, Florida) Computer-aided design and computer-assisted manufacturing Manufacturing of prosthetic implant dentistry Nowadays there are various CAD CAM systems available in the market that can be used in implant dentistry. Some examples include the following: G G G G G G CAD CAM Custom Implant Abutment Systems. The Procera system (Nobel Biocare). The Atlantis abutment (Astra Tech). CARES (Computer-Aided Restoration Service; Straumann). Etkon is another available CAD CAM system which supports the prosthetic portfolio of the Straumann dental implant. CAD CAM Custom Implant Frameworks Systems.
530 G G G G Advanced Dental Biomaterials Procera (Nobel Biocare). CAM StructSURE precision milled bars (Biomet 3i). BioCad milled bars (BioCad Medical). Etkon System. Implant frameworks are divided into four groups depending on the fabrication technique employed: first, the conventional cast frameworks; second, the frameworks produced from carbon/graphite fiber-reinforced polymethylmethacrylate (Björk et al., 1986); third, the laser-welded titanium frameworks (Jemt and Linden, 1992); and finally the most recent, the CAD CAM milled frameworks (Ehrenkranz et al., 2008; Kupeyan et al., 2006). A precise virtual model can be designed by using intraoral optical or laboratory scanners of a prepared tooth, implant position, and the dental arch. The CAD CAM technology can also be used in treatment planning of various aspects of fixed and removable prosthodontics. The scanned data and CAD design are used in milling or printing crown or bridge copings, implant abutments, and dental bridges. The 3D printing can be applied in producing metal structures (Kruth et al., 2005) either directly from metals or metal alloys or indirectly from printing in burnout resins or waxes for a lost-wax process. Printing in resin or wax and then incorporating a traditional casting approach is advantageous because there is not much need for postprocessing procedures compared to the direct 3D printing of metals (Kasparova et al., 2013). In fact, the procedure of casting alloys and facilities is quite routine in dentistry. Some downsides associated with direct printing in metals or metal alloys include higher costs, health and safety concerns, and the need for extra postprocessing care (Ortorp et al., 2011). The 3D printing and the milling/ machining technologies can be used in combination to fabricate a mechanical connection to the implant with high accuracy: reproduction of complex geometry with slight waste due to the application of 3D printing and fabrication of high-precision mechanical connecting surfaces due to the application of milling/machining technology. Although the milling procedure causes some waste of the material, it has advantages because the incorporated material is homogeneous and undisturbed by operating procedures. In fact, the cost is significantly lower, and there is little requirement for postprocessing procedures. 19.7 Lasers and dental applications Lasers in dentistry are not new, but their clinical applications have evolved and diversified in the past decade, making it very popular in a wide range of dental treatments. Lasers give dentists the capability to perform a variety of hard and soft tissue procedures with improved patient outcomes, fewer traumas, and reduced postoperative complications. This new technology dramatically expands the scope of methods a dentist can offer their patients.
Digital dentistry 531 19.7.1 History of lasers in dentistry The term LASER is an acronym for “light amplification by the stimulated emission of radiation” and was first introduced to the public in 1959, in an article by a Columbia University graduate student Gordon Gould (Gross and Hermann, 2007). Theodore Maiman, at the Hughes Research Laboratories in Malibu, California, built the first functioning laser by using a mixture of helium and neon. A laser generated from crystals of yttrium aluminum garnet treated with 1% 3% neodymium (Nd: YAG) was developed in 1961 (Gross and Hermann, 2007) and further led to the development of the argon laser in 1962. The ruby laser became the first medical laser to coagulate retinal lesions when it was used in 1963 (Gross and Hermann, 2007). In 1964 Patel at Bell Laboratories developed the CO2 laser (Gross and Hermann, 2007). The first reported oral surgical application using a CO2 laser occurred in 1977 (Lenz et al., 1977). Because of the thermal nature of these soft tissue lasers, injections were required in most cases. In January 1987 the first FDA clearance for a CO2 laser used in oral surgery paved the way for the acceptance and viability of using lasers in the oral cavity in a clinical environment. Dr. Terry Meyers and his brother William, an ophthalmologist, selected the Nd:YAG laser for experiments on the removal of incipient caries (Meyers and Meyers, 1985). This set the stage for all future lasers to be developed and approved. 19.7.2 Types of lasers 19.7.2.1 Carbon dioxide laser The CO2 laser wavelength has a very high affinity for water, resulting in rapid soft tissue removal and hemostasis with a very shallow depth of penetration. Although it possesses the highest absorbance of any laser, disadvantages of the CO2 laser are its relatively large size and the high cost and hard tissue destructive interactions (Fujiyama et al., 2008). 19.7.2.2 Neodymium:yttrium aluminum garnet laser The Nd:YAG wavelength is highly absorbed by the pigmented tissue, making it a very effective surgical laser for cutting and coagulating dental soft tissues with good hemostasis. In addition to its surgical applications (Fujiyama et al., 2008), there has been research on using the Nd:YAG laser for nonsurgical debridement in periodontal disease control (Aoki et al., 2008) and the laser-assisted new attachment procedure (Slot et al., 2009). 19.7.2.3 Erbium laser The erbium lasers have two distinct wavelength options: Er,Cr:YSGG (yttrium scandium gallium garnet) lasers and Er:YAG (yttrium aluminum garnet) lasers. The erbium wavelengths have a high affinity for hydroxyapatite and the highest absorption of water in any dental laser wavelengths (Harashima et al., 2005). It is
532 Advanced Dental Biomaterials the laser of choice for treatment of dental hard tissues but can also be used for soft tissue ablation (Ishikawa et al., 2008). 19.7.2.4 Diode laser All diode wavelengths are absorbed primarily by tissue pigment (melanin) and hemoglobin. The active medium of the diode laser is a solid-state semiconductor made of aluminum, gallium, arsenide, and, occasionally, indium, which produces laser wavelengths, ranging from approximately 810 to 980 nm. Conversely, they are poorly absorbed by the hydroxyapatite and water present in the enamel. They are useful for specific procedures, such as esthetic gingival recontouring, soft tissue crown lengthening, exposure of soft tissue impacted teeth, removal of inflamed and hypertrophic tissue, frenectomies, and photostimulation of some soft tissue lesions (Hilgers and Tracey, 2004). 19.7.3 Mechanism of laser action Laser light is a monochromatic light and consists of a single wavelength of light. It consists of three principal parts: an energy source, an active lasing medium, and two or more mirrors that form an optical cavity or resonator. Energy is supplied to the laser system by a pumping mechanism, such as a flash-lamp strobe device, an electrical current, or an electrical coil. This energy is pumped into an active medium contained within an optical resonator, producing a spontaneous emission of photons. Subsequently, amplification by stimulated emission takes place as the photons are reflected back and forth through the medium by the highly reflective surfaces of the optical resonator, prior to their exit from the cavity via the output coupler. In dental lasers, the laser light is delivered from the laser to the target tissue via a fiber optic cable, hollow waveguide, or articulated arm. The wavelength and other properties of the laser are determined primarily by the composition of an active medium, which can be a gas, a crystal, or a solid-state semiconductor. The light energy produced by a laser can have four different interactions with a target tissue (Hilgers and Tracey, 2004; Carroll and Humphreys, 2006): reflection, transmission, scattering, and absorption. When a laser is absorbed, it elevates the temperature and produces photochemical effects depending on the water content of the tissues. When a temperature of 100 C is reached, vaporization of the water within the tissue occurs, a process called ablation. At temperatures below 100 C, but above approximately 60 C, proteins begin to denature, without vaporization of the underlying tissue. Conversely, at temperatures above 200 C, the tissue is dehydrated and then burned, resulting in an undesirable effect called carbonization (Sulieman, 2005). The primary chromophores in the intraoral soft tissue are melanin, hemoglobin, and water, and in dental hard tissues are water and hydroxyapatite. Different laser wavelengths have different absorption coefficients with respect to these primary tissue components, making the laser selection procedure dependent (Weiner, 2004).
Digital dentistry 533 Depending on the application of various tissues, use of lasers in dentistry can be categorized as follows: Soft tissue: G G G G G G G G G G G G G G G Incision, excision, vaporization, ablation, and coagulation of oral soft tissues Exposure of unerupted teeth Sulcular debridement Flap preparation—incision of soft tissue to prepare a flap and expose the bone Frenectomy and frenotomy Gingival troughing for crown impressions Gingivectomy or gingivoplasty Hemostasis Implant recovery Incision and drainage of abscesses Operculectomy Oral papillectomies Reduction of gingival hypertrophy Soft tissue crown lengthening Treatment of canker sores, herpetic, and aphthous ulcers of the oral mucosa Hard tissue: G G G G G G G G G G G G G Class I, II, III, IV, and V cavity preparation Caries removal Hard tissue surface roughening and etching Enameloplasty, excavation of pits, and fissures for placement of sealants Tooth preparation to obtain access to a root canal Root canal preparation including enlargement Root canal debridement and cleaning Cutting bone to prepare a window access to the apex (apices) of the root(s) Apicoectomy—amputation of the root end Root end preparation for retrofill amalgam or composite Cutting, shaving, contouring, and resection of oral osseous tissues Osteoplasty and osseous recontouring Ostectomy Laser technology is constantly refining and evolving itself over years of development. Laser-based photochemical reactions hold a great future for additional applications, particularly for targeting specific cells, pathogens, or molecules. Future growth is expected to be a combination of diagnostic and therapeutic laser techniques. It is expected that laser technologies will become an essential component of contemporary dental practice in the future. 19.8 Technology and dental education Digital technology has been increasingly utilized in dental education. Many dental schools in the United Kingdom have adopted computer-assisted learning
534 Advanced Dental Biomaterials tools/platforms such as virtual learning environments, virtual simulators, and computer-based assessments in their undergraduate and postgraduate education. E-learning tools made distance learning possible in postgraduate education in the dental schools. Various learning management systems, such as Blackboard, First Class, and Luvit, along with various shareware or freeware, such as Moodle platforms, have been commonly used worldwide. In addition to day-to-day practice, virtual learning environments assist graduates in their professional development and further education. Advantages such as remote and timeless access make it possible for students to learn anywhere, at any time, and at their own pace. Education providers and practitioners and audience can interact and exchange information while being away from their own place. This is critical when local access to education sources is either limited by time, distance, or resources. Education materials can be customized for individuals’ needs, coauthored, shared, retrieved, or updated easily. Interactive learning, self-assessment, feedback, and scoring systems, which are embedded in e-learning platforms, encourage studentcentered learning. Fast development of hardware technologies such as smartphones has progressively facilitated accessibility to digital learning environments (Mattheos et al., 2008). However, the availability of such devices or Internet is not possible for everybody everywhere. Besides, some dental and dental hygiene students prefer e-learning as an adjunct learning method along with conventional classroom teaching, rather than replacing it entirely (McCann et al., 2010); this method is referred as blended learning. Along with theoretical subjects, digital technology has been vastly used in teaching, learning, and assessing skills in dentistry. Examples of this type of system are SIMODONT dental trainer, PerioSim, and Voxel-Man in which tactile sense (Haptic technology), in combination with 3D graphics, helps in learning the preclinical skills. Clinical scenarios can be fed into the system and students can provide the treatment in a completely digital environment. These systems simulate visual, verbal, and tactile senses and the trainee can examine the dummy patient and experience tooth preparation in a virtual environment. This procedure can be observed and assessed remotely by the tutors and feedback will also be provided. This provides an opportunity to retrieve and review the procedure of the tooth preparation by the tutor and tutee rather than the final product. However, digital simulators need further development to mimic the real clinical conditions such as saliva and water spray, patient movements, and pain, and exercises such as placing restorations and minor surgeries still need to be incorporated into the systems. The complexity of learning management systems makes it necessary for the teachers to be familiar with the function of such systems or receive close technical support to be able to simplify the interface for the users. Also, the content of the virtual learning environments should be evaluated in terms of validity, accuracy, and currency through a peer-review procedure and feedback should be collected from the users to improve the system constantly (McCann et al., 2010). Despite various advantages, some studies have showed that students still prefer conventional teaching methods over digital interactive systems (Arús et al., 2017).
Digital dentistry 535 References ICRP. 2007. Preface, executive summary and glossary. Ann. ICRP, 37, 9 34. 3D Printing. 2012. Report on Additive Manufacturing and 3D Printing State of the Industry. Wohlers Associates, Inc.: Annual Worldwide. Aoki, A., Mizutani, K., Takasaki, A., Sasaki, K., Nagai, S., Schwarz, F., 2008. Current status of clinical laser applications in periodontal therapy. Gen. Den. 56, 674 687. Arisan, V., Karabuda, C., Ozdemir, T., 2010. Implant surgery using bone- and mucosasupported stereolitho-graphic guides in totally edentulous jaws: surgical and postoperative outcomes of computer-aided vs. standard techniques. Clin. Oral Implant Res. 21, 980 988. Arisan, V., Karabuda, Z., Piskin, B., Ozdemir, T., 2013. Conventional multi-slice computed tomography (CT) and cone-beam CT (CBC T) for computer-aided implant placement. Part II: Reliability of mucosa-supported stereolithographic guides. Clin. Implant Dent. Relat. Res. 15, 907 914. Arús, N.A., da Silva, A.M., Duarte, R., Silveira, H., Silveira, H., 2017. Teaching dental students to understand the temporomandibular joint using MRI: comparison of conventional and digital learning methods. J. Dent. Educ. 81, 752 758. Bagis, N., Kolsuz, M.E., Kursun, S., Orhan, K., 2015. Comparison of intraoral radiography and cone-beam computed tomography for the detection of periodontal defects: an in vitro study. BMC Oral Health 15, 64. Bammani, S.S., Birajdar, P., Metan, S.S., 2013. Application of CAD and SLA method in dental prosthesis. AMAE Int. J. Manuf. Mater. Sci. 3, 14 18. Birkfellner, W., Hummel, J., Wilson, E., Cleary, K., 2008. Tracking devices. In: Peters, T., Cleary, K. (Eds.), Text Book of Image-Guided Interventions, first ed. Springer Science. Björk, N., Ekstrand, K., Ruyter, I., 1986. Implant-fixed, dental bridges from carbon/graphite fibre reinforced poly(methyl methacrylate). Biomaterials 7, 73 75. Carroll, L., Humphreys, T., 2006. Laser-tissue interactions. Clin. Dermatol. 24, 2 7. Casap, N., Wexler, A., Persky, N., Schneider, A., Lustmann, J., 2004. Navigation surgery for dental implants: assessment of accuracy of the image guided implantology system. J. Oral Maxillofac. Surg. 62, 116 119. CEREC, 2016. CEREC 3D Preparation Guidelines. Choi, J.Y., Lee, H.J., Jang, J.U., Yeo, I.S., 2012. Comparison between bioactive fluoride modified and bioinert anodically oxidized implant surfaces in early bone response using rabbit tibia model. Implant Dent. 21, 124 128. Chu, S.J., Rd, T., Rd, P., 2010. Dental color matching instruments and systems. Review of clinical and research aspects. J. Dent. 38, e2 e16. Cleveland Clinic 2018. Dental X-rays: Intraoral & Extraoral | Cleveland Clinic. [online] Available at: https://my.clevelandclinic.org/health/articles/11199-types-of-dental-x-rays [Accessed 18 Mar. 2018]. Dabrowski, B., Swieszkowski, W., Godlinski, D., Kurzyd-Lowski, K.J., 2010. Highly porous titanium scaffolds for orthopaedic applications. J. Biomed. Mater. Res. B. Appl. Biomater. 95, 53 61. Dawood, A., Marti Marti, B., Sauret-Jackson, V., Darwood, A., 2015. 3D printing in dentistry. Br. Dent. J. 219, 521 529. Degidi, M., Nardi, D., Piattelli, A., 2012. 10-Year follow-up of immediately loaded implants with TiUnite porous anodized surface. Clin. Implant Dent. Relat. Res. 14, 828 838.
536 Advanced Dental Biomaterials Di Giacomo, G., Cury, P., de Araujo, N., Sendyk, W., Sendyk, C., 2005. Clinical application of stereolithographic surgical guides for implant placement: preliminary results. J. Periodontol. 76, 503 507. Di Giacomo, G., Silva, J., Silva, A., Paschoal, G.H., Cury, P., Szarf, G., 2012. Accuracy and complications of computer-designed selective laser sintering surgical guides for flapless dental implant placement and immediate definitive prosthesis installation. J. Period. Ontol. 83, 410 419. D’souza, K., Aras, M., 2012. Applications of computer-aided design/computer-assisted manufacturing technology in dental implant planning. J. Dent. Implant 2, 37 41. Duret, F., Blouin, J., Duret, B., 1988. CAD-CAM in dentistry. J. Am. Dent. Assoc. 117, 715 720. Ehrenkranz, H., Langer, B., Marotta, L., 2008. Complete-arch maxillary rehabilitation using a custom-designed and manufactured titanium framework: a clinical report. J. Prosthet. Dent. 99, 8 13. Eickholz, P., Kim, T., Benn, D., Staehle, H., 1998. Validity of radiographic measurement of interproximal bone loss. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endodontol. 85, 99 106. Elian, N., Jalbout, Z., Classi, A., Wexler, A., Sarment, D., Tarnow, D., 2008. Precision of flapless implant placement using real-time surgical navigation: a case series. Int. J. Oral Maxillofac. Implants 23, 1123 1127. Elias, C.N., Gravina, P., Silva Filho, C., Nascente, P., 2012. Preparation of bioactive titanium surfaces via fluoride and fibronectin retention. Int. J. Biomater. 2012, 7. FDI World Dental Federation 2018. Radiation Safety in Dentistry. [online] Available at: https://www.fdiworlddental.org/resources/policy-statements-and-resolutions/radiation-safety-in-dentistry [Accessed 23 Mar. 2018]. Flügge, T., Hövener, J., Ludwig, U., Eisenbeiss, A.-K., Spittau, B., Hennig, J., et al., 2016. Magnetic resonance imaging of intraoral hard and soft tissues using an intraoral coil and FLASH sequences. Eur. Radiol. 26, 4616 4623. Fortin, T., Champleboux, G., Lormee, J., Coudert, J., 2000. Precise dental implant placement in bone using surgical guides in conjunction with medical imaging techniques. J. Oral Implants 26, 300 303. Fortin, T., Bosson, J., Isidori, M., Blanchet, E., 2006. Effect of flapless surgery on pain experienced in implant placement using an image-guided system. Int. J. Oral Maxillofac. Implants 21, 298 304. Fujiyama, K., Deguchi, T., Murakam, I.T., Fujii, A., Kushima, K., Takano-Yamamoto, T., 2008. Clinical effect of CO2 laser in reducing pain in orthodontics. Angle Orthod. 78, 299 303. Ganeles, J., Mandelaris, G., Rosenfeld, A., Rose, L., 2011. Image guidance for implants improves accuracy and predictability. Compend. Contin. Educ. Dent. 32, 52 55. Gomez, J., 2015. Detection and diagnosis of the early caries lesion. BMC Oral Health . Greenberg, A., 2015. Digital technologies for dental implant treatment planning and guided surgery. Oral Maxillofac. Surg. Clin. N. Am. 27, 319 340. Gross, A., Hermann, T., 2007. History of lasers. World J. Urol. 25, 217 220. Gulati, M., Anand, V., Salaria, S., Jain, N., Gupta, S., 2015. Computerized implant dentistry: advances toward automation. J. Indian Soc. Periodontol. 19, 5 10. Harashima, T., Kinoshita, J., Kimura, Y., Brugnera, A., Zanin, F., Pecora, J., 2005. Morphological comparative study on ablation of dental hard tissue at cavity preparation by Er:YAG and Er,CR:YSGG lasers. Photomed. Laser. Surg. 23, 52 55. Hilgers, J., Tracey, S., 2004. Clinical uses of diode lasers in orthodontics. J. Clin. Orthod. 38, 266 273.
Digital dentistry 537 Hultin, M., Svensson, K., Trulsson, M., 2012. Clinical advantages of computer guided implant placement: asystematic review. Clin. Oral Implants Res. 23, 124 135. Idiyatullin, D., Corum, C., Moeller, S., Prasad, H.S., Garwood, M., Nixdorf, D.R., 2011. Dental magnetic resonance imaging: making the invisible visible. J. Endod. 37, 745 752. Ishikawa, I., Aoki, A., Takasaki, A., 2008. Clinical application of erbium:YAG Laser in periodontology. J. Int. Acad. Periodontol. 10, 22 30. Jemt, T., Linden, B., 1992. Fixed implant-supported prostheses with welded titanium frameworks. Int. J. Periodontics Restorative Dent. 12, 177 184. Jung, R., Schneider, D., Ganeles, J., Wismeijer, D., Zwahlen, M., Hämmerle, C., 2009. Computer technology applications in surgical implant dentistry: a systematic review. Int. J. Oral Maxillofac. Implants 24, 92 109. Kasparova, M., Grafova, L., Dvorak, P., Dostalova, T., Procházka, A., Kakawand, S., et al., 2013. Possibility of reconstruction of dental plaster cast from 3D digital study models. Biomed. Eng. Online 12, 49. Katsoulis, J., Panzera, P., Mericske-Stern, R., 2009. Prosthetically driven, computer-guided implant planning for the edentulous maxilla: a model study. Clin. Implant Dent. Rel. Res. 11, 238 244. Kelly, J.R., Benetti, P., 2011. Ceramic materials in dentistry: historical evolution and current practice. Aust. Dent. J. 56 (Suppl. 1), 84 96. Kim, T., Benn, D., Eickholz, P., 2002. Accuracy of computer-assisted radiographic measurement of interproximal bone loss in vertical bone defects. Oral Surg. Oral. Med. Oral Pathol. Oral Radiol. Endod. 94, 379 387. Komine, F., Blatz, M., Matsumura, H., 2010. Current status of zirconia-based fixed restorations. J. Oral Sci. 52, 531 539. Koutsoukis, T., Zinelis, S., Eliades, G., Al-Wazzan, K., Rifaiy, M.A., Al Jabbari, Y.S., 2015. Selective laser melting technique of Co-Cr dental alloys: a review of structure and properties and comparative analysis with other available techniques. J. Prosthodont. 24, 303 312. Kruth, J., Vandenbroucke, B., Van Vaerenbergh, J., Naert, I., 2005. Digital manufacturing of biocompatible metal frameworks for complex dental prostheses by means of SLS/SLM. In: Proceedings of the Second International Conference on Advanced Research in Virtual and Rapid Prototyping, Leiria. Kunii, J., Hotta, Y., Tamaki, Y., Ozawa, A., Kobayashi, Y., Fujishima, A., et al., 2007. Effect of sintering on the marginal and internal fit of CAD/CAM-fabricated zirconia frameworks. Dent. Mater. J. 26, 820 826. Kupeyan, H., Shaffner, M., Armstrong, J., 2006. Definitive CAD/CAM-guided prosthesis for immediate loading of bone-grafted maxilla: a case report. Clin. Implant Dent. Relat. Res. 8, 161 167. Li, J., Pow, E., Zheng, L., Ma, L., Kwong, D., Cheung, L., 2015. Effects of calcium phosphate nanocrystals on osseointegration of titanium implant in irradiated bone. Biomed. Res. Int. 2015, 6. Mangano, C., Piattelli, A., Mangano, F., Perrotti, V., Iezzi, G., 2009. Immediate loading of modified acid etched dental implants in postextraction sockets: a histological and histomorphometrical comparative study in nonhuman primate papio ursinus. Implant Dent. 18, 142 150. Mangano, F., Bazzoli, M., Tettamanti, L., Mangano, C., 2013. Custom-made, selective laser sintering (SLS) blade implants as a non-conventional solution for the prosthetic rehabilitation of extremely atrophied posterior mandible. Lasers Med. Sci. 28, 1241 1247.
538 Advanced Dental Biomaterials Mangano, F., Luongo, F., Shibli, J., Anil, S., Mangano, C., 2014a. Maxillary overdentures supported by four splinted direct metal laser sintering implants: a 3-year prospective clinical study. Int. J. Dent. 2014, 9. Mangano, F.G., De Franco, M., Caprioglio, A., Macchi, A., Piattelli, A., Mangano, C., 2014b. Immediate, non-submerged, root-analogue direct laser metal sintering (DLMS) implants: a 1-year prospective study on 15 patients. Lasers Med. Sci. 29, 1321 1328. Martins, R., Lederman, H., 2013. Virtual planning and construction of prototyped surgical guide in implant surgery with maxillary sinus bone graft. Acta Cir. Bras. 28, 683 690. Masri, R., Driscoll, C., 2015. Clinical Applications of Digital Dental Technology. Wiley Blackwell Publishing. Mathew, C., Maller, S., Maheshwaran, 2013. Interactions between magnetic resonance imaging and dental material. J. Pharm. Bioallied Sci. 5, 113 116. Mattheos, N., Stefanovic, N., Apse, P., 2008. Potential of information technology in dental education. Eur. J. Dent. Educ. 12, 85 92. McCann, A., Schneiderman, E., Hinton, R., 2010. E-teaching and learning preferences of dental and dental hygiene students. J. Dent. Educ. 74, 65 78. Miyazaki, T., Hotta, Y., Kunii, J., Kuriyama, S., Tamaki, Y., 2009. A review of dental CAD/ CAM: current status and future perspectives from 20 years of experience. Dent. Mater. J. 28, 44 56. Moin, D.A., Hassan, B., Mercelis, P., Wismeijer, D., 2013. Designing a novel dental root analogue implant using cone beam computed tomography and CAD/CAM technology. Clin. Oral Implants Res. 24, 25 27. Monjo, M., Petzold, C., Ramis, J.M., Lyngstadaas, S.P., Ellingsen, E., 2012. In vitro osteogenic properties of two dental implant surfaces. Int. J. Biomater. 2012, 1 14. Mullen, L., Stamp, R.C., Brooks, W.K., Jones, E., Sutcliffe, C.J., 2009. Selective laser melting: a regular unit cell approach for the manufacture of porous, titanium, bone in-growth constructs, suitable for orthopedic applications. J. Biomed. Mater. Res. B. Appl. Biomater. 89, 325 334. Neugebauer, J., Stachulla, G., Ritter, L., Dreiseidler, T., Mischkowski, R., Keeve, E., et al., 2010. Computer-aided manufacturing technologies for guided implant placement. Expert Rev. Med. Devices 7, 113 129. Nikzad, S., Azari, A., 2008. A novel stereolithographic surgical guide template for planning treatment involving a mandibular dental implant. J. Oral Maxillofac. Surg. 1446 1454. Niraj, L., 2016. MRI in dentistry—a future towards radiation free imaging—systematic review. J. Clin. Diagnostic Res. 10, ZE14 ZE19. Nkenke, E., Vairaktaris, E., Bauersachs, A., Eitner, S., Budach, A., Knipfer, C., 2012. Acceptance of virtual dental implant planning software in an undergraduate curriculum: a pilot study. BMC Med. Educ. 12, 12 90. Ortorp, A., Jonson, D., Mouhsen, A., Vult Von Steyern, P., 2011. The fit of cobalt-chromium three unit fixed dental prostheses fabricated with 4 different techniques: a comparative in vitro study. Dent. Mater. J. 27, 356 363. Ozan, O., Turkyilmaz, I., Ersoy, A., Mc Glumphy, E., Rosenstiel, S.F., 2009. Clinical accuracy of 3 different types of computed tomography-derived stereolithographic surgical guides in implant placement. J. Oral Maxillofac. Surg. 67, 94 401. Parks, E., Williamson, G., 2002. Digital radiography: an overview. J. Contemp. Dent. Pract. 3, 1 10. Pettersson, A., Kero, T., Gillot, L., Cannas, B., Faldt, J., Soderberg, R., et al., 2010. Accuracy of CAD /CAM-guided surgical template implant surgery on human cadavers: Part I. J. Prosthet. Dent. 103, 335 342.
Digital dentistry 539 Priest, G., 2005. Virtual-designed and computer-milled implant abutments. J. Oral Maxillofac. Surg. 63, 22 32. Qualtrough, A.J., Piddock, V., 1995. Dental CAD/CAM: a millstone or a milestone? Dent. Update 22, 200 204. Raico Gallardo, Y., da Silva-Olivio, I.R.T., Mukai, E., Morimoto, S., Sesma, N., Cordaro, L., 2016. Accuracy comparison of guided surgery for dental implants according to the tissue of support: a systematic review and meta-analysis. Clin. Implant Dent. Relat. Res. 28, 602 612. Raigrodski, A., 2004. Contemporary materials and technologies for all-ceramic fixed partial dentures: a review of the literature. J. Prosthet. Dent. 92, 557 562. Ritenour, E., Gibbs, S., 2010. Dental Patient Doses Information. Health Physics Society. Rubio Serrano, M., Albalat Estela, S., Peñarrocha Diago, M., Peñarrocha Diago, M., 2008. Software applied to oral implantology. Med. Oral Patol. Oral Cir. Buca 13, 661 665. Scarfe, W.C., Farman, A.G., 2008. What is cone-beam CT and how does it work?. Dent. Clin. 52, 707 730. Scarfe, W.C., Farman, A.G., Sukovic, P., 2006. Clinical applications of cone-beam computed tomography in dental practice. J. Can. Dent. Assoc. 72, 75 80. Schneider, D., Marquardt, P., Zwahlen, M., Jung, R., 2009. A systematic review on the accuracy and the clinical outcome of computer-guided template-based implant dentistry. Clin. Oral Implant Res. 20, 73 86. Sedentexct.eu 2012. Radiation Protection Cone Beam CT for Dental and Maxillofacial Radiology. [online] Available at: http://www.sedentexct.eu/files/radiation_protection_172.pdf [Accessed 24 Mar. 2018]. Sesma, N., Pannuti, C.M., Cardaropoli, G., 2012. Retrospective clinical study of 988 dual acid-etched implants placed in grafted and native bone for single-tooth replacement. Int. J. Oral Maxillofac. Implants 27, 1243 1248. Sherry, J., Sims, L., Balshi, S., 2007. A simple technique for immediate placement of definitive engaging custom abutments using computerized tomography and flapless guided surgery. Quintessence Int. 38, 755 762. Slot, D., Kranendonk, A., Paraskevas, S., Van der Weijden, F., 2009. The effect of a pulsed Nd:YAG laser in non-surgical periodontal therapy. J. Periodont. 80, 1041 1056. Somogyi-Ganss, E., Holmes, H., Jokstad, A., 2015. Accuracy of a novel prototype dynamic computer-assisted surgery system. Clin. Oral Implants Res. 26, 882 890. Sorensen, J., Choi, C., Fanuscu, M., Mito, W., 1998. IPS empress crown system: three-year clinical trial results. J. Calif. Dent. Assoc. 26, 130 136. Spector, L., 2008. Computer-aided dental implant planning. Dent. Clin. North Am. 52, 761 775. Sulieman, M., 2005. An overview of the use of lasers in general dentist practice, laser wavelengths, soft and hard tissue clinical applications. Dent. Update 32, 286 288. 291 294. Tam, W., Lee, H., 2012. Dental shade matching using a digital camera. J. Dent. 40 (Suppl.), e3 e10. Tam, W., Lee, H., 2017. Accurate shade image matching by using a smartphone camera. J. Prosthodont. Res. 61, 168 176. Taskonak, B., Sertgöz, A., 2006. Two-year clinical evaluation of lithia-disilicate-based allceramic crowns and fixed partial dentures. Dent. Mater. 22, 1008 1013. Ti-Sun, K., Benn, D.K., Eickholz, P., 2002. Accuracy of computer-assisted radiographic measurement of interproximal bone loss in vertical bone defects. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 94, 379 387. Valente, F., Schiroli, G., Sbrenna, A., 2009. Accuracy of computer-aided oral implant surgery: a clinical and radiographic study. Int. J. Oral Maxillofac. Implants 24, 234 242.
540 Advanced Dental Biomaterials Vandenberghe, B., Jacobs, R., Bosmans, H., 2010. Modern dental imaging: a review of the current technology and clinical applications in dental practice. Eur. Radiol. 20, 2637 2655. Vasak, C., Watzak, G., Gahleitner, A., Strbac, G., Schemper, M., Zechner, W., 2011. Computed tomography-based evaluation of template (NobelGuidet)-guided implant positions: a prospective radiological study. Clin. Oral Implants Res. 22, 1157 1163. Venkatesh, K.V., Nandini, V.V., 2013. Direct metal laser sintering: a digitised metal casting technology. J. Indian Prosthodont. Soc. 13, 389 392. Watzinger, F., Birkfellner, W., Wanschitz, F., Millesi, W., Schopper, C., Sinko, K., et al., 1999. Positioning of dental implants using computer-aided navigation and an optical tracking system: case report and presentation of a new method. J. Craniomaxillofac. Surg. 27, 77 81. WebMD Boots 2018. Learn more about dental X-rays. [online] Available at: https://www. webmd.boots.com/oral-health/guide/dental-x-rays [Accessed 18 Mar. 2018]. Weiner, G.P., 2004. Laser dentistry practice management. Dent. Clin. North Am. 48, 1105 1126. Widmann, G., Bale, R., 2006. Accuracy in computer-aided implant surgery—a review. Int. J. Oral Maxillofac. Implants 21, 305 313. Winter, A., Pollack, A., Frommer, H., Koenig, L., 2005. Cone beam volumetric tomography vs. medical CT scanners. N. Y. State Dent. J. 71, 28 33. Further reading Ruppin, J., Popovic, A., Strauss, M., Spüntrup, E., Steiner, A., Stoll, C., 2008. Evaluation of the accuracy of three different computer-aided surgery systems in dental implantology: optical tracking vs. stereolithographic splint systems. Clin. Oral Implants Res. 19, 709 716.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 20 Mohamed-Nur Abdallah1, Tiantong Lou1, Jean-Marc Retrouvey2 and Sunjay Suri1 1 Faculty of Dentistry, University of Toronto, Toronto, ON, Canada, 2Division of Orthodontics, Faculty of Dentistry, McGill University, Montreal, QC, Canada Chapter Outline 20.1 Introduction 541 20.2 Orthodontic brackets 542 20.2.1 Metal brackets 542 20.2.2 Plastic brackets 550 20.2.3 Ceramic brackets 551 20.3 Orthodontic archwires 553 20.3.1 Properties of orthodontic archwires 553 20.3.2 Classification of orthodontic archwires 555 20.4 Clear aligners 20.4.1 20.4.2 20.4.3 20.4.4 20.4.5 20.4.6 20.5 Final remarks References 572 20.1 562 Material composition 562 The thermoforming process 565 Forces of thermoplastic aligners 565 Mechanical properties 567 Attachments 569 Cytotoxicity 571 572 Introduction Orthodontics, similar to other fields in dentistry, is in constant need for research and improvement in the biomaterials used. Indeed, the field of orthodontic materials has vastly expanded in the recent decades with the fabrication of new archwire materials, new brackets materials and systems, as well as the introduction of clear aligners. Box 20.1 summarizes the different classes of orthodontics materials. This chapter will focus on the materials used for manufacturing the orthodontic brackets, archwires, and clear aligners. Most of the other materials are covered under different chapters in this book. Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00020-7 Copyright © 2019 Elsevier Ltd. All rights reserved.
542 Advanced Dental Biomaterials Box 20.1 Classification of orthodontic materials According to their usage: 1. Orthodontic brackets: stainless steel, titanium, plastic or polycarbonates, and ceramic. a. Wire alloys b. Metallic wires (e.g., stainless steel, β-titanium, nickel titanium) 2. Nonmetallic wires (e.g., Optiflex, composite coated wires) 3. Elastomeric ligatures and chains: conventional ligatures, fluoride-releasing elastomerics 4. Adhesive composite resins: (e.g., chemically cured, light cured, dual cured) 5. Cements (e.g., zinc polycarboxylate, glass ionomer) 6. Miscellaneous materials (e.g., alginate impression materials, brass lock pins) According to the structures of orthodontic materials 1. Metallic materials: Used to manufacture brackets, wires, and orthodontics bands. 2. Ceramic materials: Used for fabricating ceramic brackets and as filler particles in cements. 3. Polymeric materials: A broad range of polymeric materials are used in orthodontics including adhesive resins for bonding, impression materials, polycarbonates for brackets, and polyurethane elastomers for tooth movement. 20.2 Orthodontic brackets Orthodontic brackets are attachments that are bonded to the outer surface of the crown of the tooth in order to transmit forces from the appliance to the teeth, with the ultimate aim to induce the desired tooth movement. Orthodontic brackets can be classified according to the following: 1. Morphology: single bracket, Siamese (twin bracket), Lewis bracket, mini-twin, selfligating, etc. 2. Technique used: Begg light wire appliance, edgewise appliance, straight wire appliance, tip-edge appliance, lingual preadjusted edgewise appliance, etc. 3. Bracket slot dimension: 0.018v 3 0.025v, 0.022v 3 0.022v 3 0.028v, etc. 4. Core manufacturing material: metal brackets, plastic brackets, and ceramic brackets (Table 20.1). Materials used for manufacturing brackets have undergone a vast improvements in the past few decades, particularly, with the use of materials other than metals to meet the esthetic demands of patients. Yet, the metal brackets remain the most frequently used for routine orthodontic treatment. This section will focus on the materials used for manufacturing the brackets. 20.2.1 Metal brackets Orthodontic brackets were originally fabricated stainless steel (SS) alloys where the wings and the base were manufactured by machining and/or casting and then the
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 543 Table 20.1 Classification of brackets according to the manufacturing material. Type of bracket Examples Metal Stainless steel brackets Cobalt chromium brackets Titanium brackets Precious metal brackets Polycarbonate brackets Polyurethane-composite brackets Thermoplastic-polyurethane brackets Monocrystalline alumina Polycrystalline alumina Polycrystalline alumina Plastic Ceramic different parts were connected by soldering (Eliades and Brantley, 2016). Recent advances in developing new materials as well as the fabrication techniques (e.g., metal injection molding and laser welding) have supplied the orthodontic field with brackets made of cobalt chromium alloys, gold alloys, and titanium alloys (Eliades and Brantley, 2016). Furthermore, these improvements have led to the fabrication of brackets having a wide array of contradicting properties for each component according to its function. For instance, the wing of the brackets, which is in contact with the archwire, should be manufactured from a material that possesses a relatively high modulus of elasticity to minimize the undesirable plastic deformation and be hard enough to resistant wear induced during the movement of the wire in the slot as well as expression of torque, whereas the base of the bracket should bond sufficiently to the tooth enamel but also allow some degree of deformation in order to facilitate debonding from the tooth after the completion of the orthodontic treatment (Eliades and Brantley, 2016). There are four main types of metal brackets used in contemporary orthodontics: 1. 2. 3. 4. SS brackets; cobalt chromium brackets; titanium brackets; and precious metal brackets. 20.2.1.1 Stainless steel brackets SS alloys were first introduced by Hauptmeyer into dentistry in 1919 to make prostheses, and it was not until the 1930s that Angle introduced them into orthodontics as ligature wires. Following this, SS became the standard material to fabricate brackets because of its superior mechanical properties, corrosion resistance, and relatively acceptable appearance (Brantley and Eliades, 2017). From a material science perspective, SS, also called inox steel, is a steel alloy that contains a minimum of 10.5% chromium content by mass. SS alloy is primarily composed of iron, carbon, and chromium. Iron is the main ingredient of SS alloys, but because of the brittle nature of pure iron, carbon is added to provide strength. Moreover, by means of a process known as sensitization, carbon facilitates at high
544 Advanced Dental Biomaterials temperatures the precipitation of chromium as chromium carbide at the grain boundaries, and thus increases the intergranular corrosion. Chromium protects the alloy from corrosion by forming an impermeable, corrosion-resistant chromium oxide layer. Furthermore, relatively high concentrations of chromium promote the stability of the body-centered cubic (bcc) lattice structure. On the other hand, other elements that might be present in the alloy such as copper, manganese, nickel, and nitrogen favor the stability of the face-centered cubic (fcc) lattice. Table 20.2 summarizes the effect of the different elements on the SS alloy. Classifications of stainless steel alloys SS alloys can be classified into different types of crystalline structures and each type of SS has been designated a specific AISI (American Iron and Steel Institute) or UNS (unified number system) or Society of Automotive Engineers number depending upon the elemental composition of the alloy. Box 20.2 enumerates the different types of SS alloys. In orthodontics, different types of SS alloys are used in manufacturing the wings and the base components of the brackets. The SS alloys vary in composition and demonstrate a specific crystalline structure with the Table 20.2 Effect of different elements on the properties of stainless steel (SS) alloys. Element Effect Carbon Provides hardness and strength Increases chromium carbide formation due to localized corrosion from oral fluids Increases resistance to oxidation by forming a passive surface oxide layer, and thus prevents surface corrosion by blocking oxygen diffusion Stabilizes the austenitic phase Increases resistance to corrosion and oxidation Ion release might cause hypersensitivity reactions Stabilizes the austenitic phase and has been used as a substitute for nickel Similar to nickel and manganese, it is an austenite forming element Increases resistance to pitting corrosion, particularly, by chlorides Stabilizes carbide and increases resistance to corrosion Decreases sintering temperature and increases both strength and corrosion resistance Stabilizes carbon and improves corrosion resistance Chromium Nickel Manganese Nitrogen Molybdenum Titanium Phosphorous Niobium (columbium) and tantalum Copper Selenium Produces precipitation hardening properties Makes SS more machinable and workable but decreases both strength and hardness
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 545 Box 20.2 Types of stainless steel materials 1. 2. 3. 4. 5. 6. Austenitic steels (300 series) Martensitic steels (400 series) Precipitation-hardenable steels Ferrite steel Duplex stainless steel (SAF 2205) Cobalt-containing alloys austenitic SS (303, 304L, 316L) and precipitation-hardenable steels (17-4 PH) being the most commonly used (Eliades et al., 2003; Oh et al., 2005). Table 20.3 lists the composition of some of the commercially available SS brackets. Austenitic stainless steel (300 series) Austenitic SS is one of the SS alloys most commonly used as orthodontic bracket or wire materials due to its superior corrosion resistance, excellent formability, and relative low cost. The SS alloys that fall the 300 series in the AISI system are all austenitic SS alloys. According to this classification, the increase in number indicates an increase in the number of metals added to the iron, whereas the carbon concentration is decreased and the letter “L” indicates a low carbon content. Even though the lower carbon content decreases corrosion susceptibility by minimizing the harmful carbide precipitation, it also decreases the alloy’s strength. 304L steel is one of the popular SS alloys, which is also called 18-8 stainless steel, as it is composed of B18% chromium and B8% nickel, in addition to, B71% iron and B , 0.02% carbon. 18-8 stainless steel brackets are most frequently milled, whereas AISI 316L steel, which is also used to manufacture brackets, is too hard to be milled, and thus manufacturers follow a casting technique to fabricate one-piece brackets with no separate mesh-base. Standard orthodontic brackets are usually manufactured (Oh et al., 2005; Platt et al., 1997; Eliades, 2007; Matasa, 1992). 316 and 316L SS have higher corrosion resistance, particularly to chloride, and have been shown to release less nickel; therefore they are more commonly used for fabricating the base of the bracket (Eliades et al., 2002). SS alloys form as an fcc crystal by heating the alloy above 912 C (Anusavice et al., 2013). Nickel is added to the alloy to stabilize the SS in the austenitic form after cooling. Other austenitizing elements, such as manganese and nitrogen, are added to the alloy to preserve these alloys from turning into the bcc structure (also known as ferrite) at room temperature. Chromium is also important to increase the corrosion resistance, which maximizes the control of the nickel ion release and decreases the possibility of any associated hypersensitivity reactions. Nonetheless, the hostile environment of the oral cavity renders the austenitic SS susceptible to corrosion, particularly in an acidic environment and the presence of chlorine ions that may penetrate the outer protective oxide layer and cause pitting corrosion (Grimsdottir et al., 1992).
Table 20.3 Composition of stainless steel alloys used in manufacturing orthodontic brackets. Designation Example AISI UNS 303 S-30300 304L S-30403 316L Composition (wt.%) Fe Cr Ni Mn Mo C P Si S Ormco diamond Bal. 17 19 8 10 2 0.6 0.15 0.2 1.0 0.15 Advanced orthod Bal. 18 20 8 12 2 0.03 0.04 1.0 0.03 S-31603 “A” company Bal. 16 18 10 14 2 0.03 0.04 1.0 0.03 630/17-4 S-17400 Ormco min diam Bal. 15 17 3 5 1 0.07 0.04 1.0 0.04 4 Cu/ 3 Nb 630/17-7 S-17700 Ormco edge lock Bal. 16 18 6.5 7.5 1 0.09 0.04 1.0 0.04 0.08 1.5 Al SAF 2205 S-31803 CEOSA low nickel Bal. 22 5.5 2 3 0.03 0.03 1.0 0.02 0.16 N 18-8 Plus S-28200 18-8 Plus Bal. 8 0.16 18 1 0.15 0.045 1.0 0.03 0.5 N 431 AI29 S-43100 S-44735 Potential Potential Bal. Bal. 26 29 0.3 0.5 4 4 0.02 0.035 0.35 0.01 0.5 Ti 2.5 Bal. indicates that the rest of the alloy is iron (Fe) by % weight. AISI, American Iron and Steel Institute; Bal., balance iron; UNS, unified number system. Source: Anusavice (1996), K.J. Philips’ Science of Dental Materials, tenth ed. W.B. Saunders Company. Other
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 547 Super austenitic SS (SR-50A or UNS no. S32050) is defined as an SS alloy with a pitting resistance equivalent value of 40 and contains higher molybdenum and nitrogen content than conventional SS (Oh et al., 2002). Super SS exhibits higher resistance to chloride pitting and corrosion as well as good frictional properties (Oh et al., 2005). However, Super SS has been tested only in vitro studies (Oh et al., 2005), and to the best of our knowledge, no information is listed in any company catalogue indicating its use in manufacturing brackets. Precipitation-hardenable steels These alloys are formed through hardening by heat treatment. This type of SS presents a corrosion resistance equivalent to 304SS but with higher strength. The 17-4 PH SS or S174000 PH alloys are widely used to fabricate the wings of normal size or as “mini” brackets due to their relatively higher hardness and strength (Flores et al., 1994). Indeed, an in vitro study tested the microhardness of 12 direct-bonding metal brackets and concluded that the PH 17-4 showed the highest hardness values, while the austenitic 316L presented the lowest hardness values (Matasa, 1998). Despite having lower nickel content, 17-4 PH SS has poorer localized corrosion resistance and more nickel release than 316SS (Oh et al., 2005; Eliades et al., 2002). Ferrite steel These alloys exist in an bcc crystal structure which is similar to the crystal structure of iron, and thus these SS alloys have the designated name “ferrite”. The main alloying element includes chromium, titanium, molybdenum, and a small amount of carbon but no nickel. Generally, ferrite steel exhibits less corrosion resistance than austenitic SS; however, stabilized ferritic SS (AISI 441) has comparable corrosion resistance to that of 316SS. Ferritic SS demonstrates less strength than austenitic hardness because it presents with lower amounts of carbon. In orthodontics, ferrite SS with 19% 30% carbon content, also known as super ferrite SS, is utilized to construct nickel-free brackets. Martensitic steels (400 series) These alloys demonstrate an increased tensile strength and less corrosion resistance than the austenitic counterparts. Due to these properties, their use in dentistry is limited to manufacturing tools that need sharp or wear-resistant cutting edges. Duplex stainless steel (SAF 2205) This SS alloy has a duplex microstructure consisting of austenite and delta-ferrite and is used to manufacture one-piece brackets. It consists of high chromium contents (19% 32%), molybdenum (up to 5%), and relatively low nickel content compared to austenitic SS. This combination of more ductile and softer austenite crystal phases and less ductile harder delta-ferrite crystal phases results in an SS that is harder than the single-phase austenitic 316L SS, but more ductile than the single-phase ferritic 430 SS.
548 Advanced Dental Biomaterials 20.2.1.2 Cobalt chromium brackets Cobalt chromium (CoCr) brackets were introduced in the 1990s as a low nickel alternative to SS. Generally, cobalt-based alloys can be categorized into following: 1. wear-resistant alloys; 2. high-temperature alloys; and 3. corrosion-resistant alloys. From these categories, the cobalt-based wear-resistant alloy is utilized for manufacturing orthodontic brackets (Haddad et al., 2009). According to the ASTM (American Society for Testing and Materials), CoCr brackets are fabricated from ASTM F-75 (Table 20.4), in which the nickel content is kept low (up to 0.05%) (Haddad et al., 2009). CoCr brackets exhibit comparable friction to that of SS brackets when used in combination with SS wires (Moore et al., 2004), while CoCr brackets demonstrate higher frictions than titanium brackets when used with SS wires or β-titanium wires (Titanium brackets and wires are explained in more detail in sections 20.2.1.3 and 20.3) (Nair et al., 2012). Moreover, due to the relatively higher content of chromium, CoCr brackets show a lesser degree of corrosion when compared to titanium-based brackets (Schiff et al., 2005). 20.2.1.3 Titanium brackets Titanium (Ti) and its alloys possess excellent biocompatibility and increased corrosion resistance, as well as adequate mechanical properties; therefore they have been used for several biomedical applications ranging from artificial heart valves to limb prostheses and dental implants (Ratner et al., 2012). Ti brackets were first introduced by Dentaurums in 1995 to overcome nickel hypersensitivity, inadequate retention, and the susceptibility of SS to corrosion in the acidic oral environment (Hamula et al., 1996; Michelberger et al., 2000). In metallurgy, Ti can be classified into the following three types: αTi, βTi, or α β Ti. Table 20.4 Composition of the cobalt-based wear-resistant alloy (ASTM F-75). Element Weight % Cr Mo W C Fe Ni Si Mn Co 25 30 7 max 2 15 0.25 3.3 3 max 0.5 max 2 1 Balanced
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 549 αTi is an unalloyed commercially pure (CP) Ti, whereas the other two types are considered Ti alloys. βTi includes Ti 15V 3Cr 3Sn 3Al alloy, while α β Ti includes Ti 6Al 4V alloy. CP-Ti can be further classified into four grades according to the degree of purity and primary oxygen content. Grade 1CP-Ti exhibits the highest purity, corrosion resistance, and formability but lowest strength when compared to grade 4 CP-Ti, which demonstrates the highest strength and moderate formability. The composition of the different grades of CP-Ti and Ti alloys are shown in Table 20.5. In contemporary orthodontics, αTi (grade 2 and 4) and α β Ti (Ti 6Al 4V) are used to manufacture brackets (Zinelis et al., 2003; Gioka et al., 2004). Grade 2 CP-Ti is usually used to fabricate the base of the brackets because of its relatively lower strength. The stronger and harder Ti 6Al 4V alloy is used to make the wing components, which are then laser welded together. However, the welding process may leave gaps between the different parts, leading to decreased mechanical strength, plaque retentive areas, and increased crevice corrosion (Gioka et al., 2004). In addition to the previous drawbacks, due to reports of vanadium release from Ti 6Al 4V, which could be hazardous (Rodriguez-Mercado et al., 2003), some manufacturers fabricate single-unit brackets from metal injection molded or milled from grade 4 CP-Ti (Gioka et al., 2004). Ti and its alloys exhibit higher corrosion resistance than SS, due to the presence of a thin outer protective layer of passive Ti dioxide (also known as rutile), which is more stable than the chromium layer on the surface of SS (Brantley and Eliades, 2001). Moreover, the Ti dioxide layer remains stable during sliding mechanics process and can be effectively used with SS wires. Ti brackets demonstrate comparable friction resistance to SS brackets when SS archwires were used (Kusy et al., 1998; Kapur et al., 1999; Kusy and O’Grady P, 2000). Ti brackets exhibit lower friction with βTi wires when compared to other metal brackets (Nair et al., 2012). Compared to SS brackets, titanium brackets are more expensive, grayer in color and show a rougher surface texture, and exhibit greater amount of plaque accumulation and staining (Harzer et al., 2004). Ti brackets demonstrate pitting corrosion when used with fluoride containing mouthwashes or toothpaste (Toumelin-Chemla et al., 1996; Reclaru and Meyer, 1998). However, titanium exhibits a low thermal conductivity, which alleviates the patient’s sensitivity to extreme temperature changes. Table 20.5 Composition of different grades of commercially pure titanium. Grade 1 2 3 4 Composition (maximum weight %) Ti N C H Fe O Balanced Balanced Balanced Balanced 0.03 0.03 0.05 0.05 0.08 0.08 0.08 0.08 0.015 0.015 0.015 0.015 0.20 0.30 0.30 0.50 0.18 0.25 0.35 0.40
550 Advanced Dental Biomaterials 20.2.1.4 Precious metal brackets Historically, traditional edgewise brackets were fabricated from gold as they have good mechanical properties and are nonreactive; however, due to their high cost, they were replaced with SS. This type of bracket is usually a steel bracket plated with precious metals including gold, palladium, or platinum. Among these brackets, gold plated brackets (16, 18, 24 carats) are the ones most commonly used, particularly in lingual orthodontics. 20.2.2 Plastic brackets Plastic brackets were first introduced in the early 1970s as an esthetic alternative to the metal brackets. They were initially made of acrylic resins and later from unfilled polycarbonate (Eliades et al., 2004). Plastic brackets are more transparent than SS brackets, and thus more esthetically adaptive to the natural surface color of teeth (Eliades et al., 2004). The bond strength of these brackets to the adhesive bonding agent depends on the mechanical retention of the bracket base and the plastic bracket primer. Although the shear bond strength of plastic brackets was demonstrated to be significantly lower than metal brackets, it was found to be clinically acceptable (Pithon et al., 2007). There were several major inherent disadvantages of plastic brackets that led to its decreased acceptance and use by orthodontists. These problems included slot distortion and creep deformation, tie-wing fractures, water absorption, discoloration, odors, increased friction between the slot and archwire during tooth sliding, as well as inadequate strength and stiffness resulting in debonding problems (Matsui et al., 2015). For instance, polycarbonate brackets undergo permanent deformation or creep with time when subjected to a constant stress in the range of 2000 g/mm (Dobrin et al., 1975). Furthermore, it was shown that polycarbonate brackets exhibit higher torque losses and lower torquing moments when compared with the metal brackets (Harzer et al., 2004). To overcome the lack of strength and rigidity of the polycarbonate (PC) bracket, manufacturers introduced high-grade medical polyurethane brackets (Eliades et al., 2004) and polycarbonate brackets reinforced with ceramic (Arici and Regan, 1997) or fiberglass fillers (Faltermeier et al., 2007) and/or metal slots (Zinelis et al., 2005). Also, the bracket base was modified to strengthen the resin bonding without primers and minimize enamel damage related to debonding procedures (Arici and Regan, 1997). In a study comparing torque deformation of seven commercially available plastic brackets against SS brackets, results demonstrated that the metal reinforced PC brackets undergo the lowest degree of deformation followed by pure PC, pure polycarbonate, fiberglass-reinforced PC brackets, and then ceramicreinforced PC brackets (Sadat-Khonsari et al., 2004). However, it was reported that the addition of ceramic and glass fibers to PC brackets had no beneficial effect regarding frictional forces (Feldner et al., 1994; Sadat-Khonsari et al., 2004; Ali and Mann, 2004; Moller et al., 2009). Even though ceramic reinforcement and the
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 551 addition of metal slot decreased the creep, there was still a 15% loss in torque expression over 24 hours when using the PC brackets (Feldner et al., 1994). Nevertheless, it seems that PC brackets reinforced with metal slot are capable of generating clinically acceptable torque expression and their decreased slot roughness is beneficial when applying sliding mechanics during orthodontic tooth movement (Sadat-Khonsari et al., 2004). Accordingly, it was recommended to use PC brackets only if they had a metal slot in order to decrease the frictional resistance, slot deformation, and to avoid damage from the sliding movements of the archwire (Sadat-Khonsari et al., 2004; Choi et al., 2014). Further recent advancements include plastic self-ligating brackets which have been shown to exhibit lower frictional forces when compared to the conventional counterparts (Russell, 2005). Table 20.6 summarizes the features of the different types of plastic brackets. 20.2.3 Ceramic brackets The first ceramic brackets were introduced into the field of orthodontics in the mid-1980s (Swartz, 1988). Ceramic brackets are composed of tridimensional macromolecules of aluminum oxide or zirconia oxide (Birnie, 1990). Box 20.3 lists the different types of ceramic brackets. These brackets are optically more esthetic than SS brackets, offer the advantage of resisting staining and slot distortion, as well as being chemically inert to oral fluids (Reddy et al., 2013; Ansari et al., 2016). Table 20.6 Characteristics of the different types of plastic brackets. Type of bracket Examples Polycarbonate Polyurethane Reinforced polycarbonate Polyurethane composite Thermoplastic polyurethane Polymer fiber reinforced Ceramic reinforced Fiberglass reinforced Metal slot reinforced Metal slot and ceramic reinforced Characteristics Undergoes creep deformation Water absorption leading to swelling and increased friction Discoloration and odor Improved strength Improved frictional resistance Increased deformation resistance Box 20.3 Types of ceramic brackets 1. Monocrystalline alumina (sapphire) 2. Polycrystalline alumina 3. Polycrystalline zirconia yttrium partially stabilized zirconia
552 Advanced Dental Biomaterials 20.2.3.1 Polycrystalline zirconia Currently, the use of zirconia brackets is limited in the orthodontic field due to their higher friction coefficient (Keith et al., 1994), lower shear bond strength, and inferior esthetic properties (i.e., less transparent and with a yellowish tint) than alumina brackets (Springate and Winchester, 1991; Brantley and Eliades, 2001). Therefore this chapter will only focus on the alumina brackets. 20.2.3.2 Alumina ceramic brackets Contemporary ceramic brackets are mostly made of alumina. According to the manufacturing process, these brackets are categorized into two broad classes with different optical and mechanical properties: monocrystalline (sapphire) and polycrystalline (Birnie, 1990). Polycrystalline alumina brackets are composed of aluminum oxide particles that are fused by a sintering process. This procedure is relatively inexpensive and less complicated than the manufacturing of monocrystalline ceramics, making these brackets more popular (Kusy, 1988; Saunders and Kusy, 1994). During manufacturing, the aluminum oxide particles are first blended using a binder machine and then molded into a certain shape from which the different components of the brackets are cut. The molded mixture is afterward heated at temperatures that fuse the alumina particles and burns out the binder. This firing process is known as “sintering” (Brantley and Eliades, 2001). Of note, polycrystalline zirconia brackets are manufactured following a different process than alumina brackets which entails impression molding followed hot isostatic pressing (Brantley and Eliades, 2001). On the other hand, the process of manufacturing monocrystalline alumina brackets begins with melting pure aluminum oxide particles at temperatures above 2100 C; the resulting mass is slowly cooled to form a large single-crystal alumina rod (Birnie, 1990). The single-crystal alumina rod is then milled into brackets by using ultrasonic cutting or Nd:YAG lasers and heat treated to remove surface imperfections or stresses resulting from the milling process. These steps minimize the imperfections and stress-inducing impurities present in the polycrystalline ceramics. However, this process is relatively more expensive and more difficult than the polycrystaline sintering technique, and introduces sharp corners that may make the tie wings of the brackets more prone to fracture (Brantley and Eliades, 2001). Esthetically, monocrystalline brackets are clearer than the polycrystalline counterparts due to the reduction of grain boundaries and impurities which are responsible for reflecting light, thus making the polycrystalline brackets appear more radiopaque (Swartz, 1988; Birnie, 1990). The main disadvantages of alumina ceramic brackets include the following: G G Brittleness and low fracture toughness that may lead to tie-wing fracture. Increased frictional resistance: In addition to the already increased resistance to sliding between the archwires and ceramic brackets, the harder aluminum oxides can create nicks in the softer SS archwire during orthodontic tooth movement, which, in turn, further increases the binding and frictional resistance (Voudouris et al., 2010; Williams and
Biomaterials used in orthodontics: brackets, archwires, and clear aligners G G G 553 Khalaf, 2013; Kumar et al., 2014). Moreover, polycrystalline ceramics with their rougher surface finish present a higher friction resistance than monocrystalline brackets (Arici et al., 2015; Arash et al., 2015). To overcome this drawback, manufacturers have attempted to improve the frictional characteristics by the addition of metal lined/reinforced slots (Pasha et al., 2015; Arici et al., 2015). High levels of hardness: Ceramic brackets are nine times harder than tooth enamel (Swartz, 1988). Indeed, attrition of the opposing teeth in contact with ceramic brackets is a common clinical finding. Accordingly, these brackets should be avoided in deep bite cases or situations where teeth might be in contact with the bracket even for a short period. Other clinical complications include enamel fractures, flaking, or cracks, especially during debonding procedures, due to the increased bond strength from the silane coupling agent (Reddy et al., 2013; Ansari et al., 2016). Therefore ceramic brackets should be avoided in patients with developmental tooth defects, enamel cracks, several nonvital teeth, or teeth with large restorations (Olsen et al., 1996). To address this problem, ceramic brackets with a polymeric base or with various mechanical base designs have been introduced, such as a mechanical ball, dovetail, silane-coated buttons, or microcrystalline. By utilizing these mechanical retentive modifications, ceramic brackets can be cemented directly to the tooth surface using standard light-cured or chemically cured adhesives without the need for additional saline bonding agents. Moreover, it is argued that these modifications render the bond strengths and debonding characteristics of ceramic brackets comparable to that of metal brackets (Gwinnett, 1988; Bishara et al., 1993). In addition, it was demonstrated that ceramic brackets with polymeric mesh bases exhibited adequate bond strength and resulted in less tooth enamel damage at debonding when compared to conventional ceramic brackets (Ansari et al., 2016; Elekdag-Turk et al., 2009). Ceramic brackets cause less image distortion to the diagnostic quality of magnetic resonance (MRI) images of the head when compared to SS brackets (Zhylich et al., 2017). 20.3 Orthodontic archwires Orthodontic archwires and auxiliaries constitute the basic active components for moving teeth in the fixed appliance system. Orthodontic wires are engaged into the brackets to produce the desired tooth movement or can be utilized as a stable base to which force elements are connected to yield the tooth movement. 20.3.1 Properties of orthodontic archwires The clinician should consider a variety of wire parameters and characteristics according to the stage of treatment and the purpose for which the archwire is being used. A single type of archwire is not appropriate for all treatment stages and there is no archwire that is considered ideal. Some of the main properties of orthodontic wires are discussed below: G Modulus of elasticity or Young’s modulus is defined as the ratio between the stress applied to the wire and the resulting strain. It is used to describe the rigidity or the stiffness of the wire (Fig. 20.1).
554 Advanced Dental Biomaterials Figure 20.1 Diagram illustrating the properties of orthodontic wires according to the stress strain curve. G G G G G Stiffness (or load deflection rate) is defined as the magnitude of force delivered by the appliance and is proportional to the modulus of elasticity. It can also be defined as the resistance of the wire to deformation or to the force required to bend the wire to a certain distance. To illustrate this point with a clinical example, the initial archwires used during the leveling and alignment phase of orthodontic treatment should be flexible, while the working archwires used for space closure or anterior posterior correction need to be stiff enough to sustain the forces applied. Modulus of resilience is the characteristic that reflects the energy storing capacity of the wire and it is represented by the area under the line describing the elastic deformation of the wire (Fig. 20.1). Elastic deformation is a temporary shape change and the material returns to its original dimensions it had before the stress was applied. Range of activation, range of deflection, or working range are defined as the distance that the archwire can be deflected or bent without permanent deformation. This range can also be referred to as springback, maximum elastic deflection, or maximum flexibility. Springback is related to the ratio of yield strength to the modulus of elasticity of the material.Yield strength can be defined as the stress at which a specific amount of permanent plastic deformation is produced, and it is usually determined at a 0.2 percent of the unstressed length. The archwire used for aligning and leveling teeth should have a large working range or high springback in order to engage brackets attached of severely displaced or rotated teeth, while the modulus of elasticity should be low enough for the wire to exert low forces on teeth even at large deflections. Higher springback values refer to the ability of applying large activations while increasing the working time of the appliance. This results in a decrease in the number of archwire changes during treatment. It should be noted that if a wire is stretched beyond its yield strength, it will undergo permanent deformation and will not return to its original shape; however, a certain amount of springback will still occur until the failure point is reached. Strength is defined as the maximum force that an archwire can withstand before undergoing permanent deformation. Plastic or permanent deformation is the permanent distortion that occurs when a material is subjected to stresses beyond its yield strength. Formability is the property that permits bending the wire into any desired configurations (e.g., loops, springs, coils) without fracturing (i.e., without reaching the break point) the
Biomaterials used in orthodontics: brackets, archwires, and clear aligners G G G G 555 wire. Ideally, an archwire should present a combination of formability for ease of preengagement wire bending as well as an appropriate resilience for engagement and subsequent activation (Fig. 20.1). The friction between the archwire and the bracket should be as minimal as possible as excessive friction may slow down tooth movement and lead to loss of anchorage. The archwire material should be biocompatible, highly corrosion resistant to the oral environment, and not favor the growth of microorganisms. It should demonstrate also environmental stability that guarantees the maintenance of the desirable wire properties for an adequate period after manufacturing as well as to ensure a predictable behavior of the wire when in use during treatment. As feasible as possible, the archwires should be manufactured from a relatively inexpensive material. In certain clinical situations, it is preferable that the archwire can be welded or soldered to auxiliaries and attachments. 20.3.2 Classification of orthodontic archwires Orthodontic archwires can be classified according to their composition, sizes, cross sections, configurations, or preformed shapes (Fig. 20.2). Orthodontic archwires are generally manufactured from metal alloys. The manufacturing process begins with cross-sectional reduction and lengthening of the alloy to induce strain hardening and brittleness in the alloy. The wires then undergo various heat treatments to release any residual stresses and accumulated strain energy to restore their ductility and toughness. The fabrication of multistranded wires involves manufacturing of the individual strand and braiding or twisting the wire to produce the final desired configuration (Nikolai, 1997). 20.3.2.1 Gold Gold alloys were initially used to fabricate archwires because they are inert, biocompatible, and stable. Furthermore, before the widespread use of the SS wire, gold alloys were used due to the lack of the availability of other materials that can Figure 20.2 Classification of some of the main orthodontic wires.
556 Advanced Dental Biomaterials tolerate the oral conditions. Gold wires are generally composed by of gold (15 65 wt.%), copper (11 15 wt.%), silver (10 25 wt.%), palladium (5 10 wt.%), platinum (5 10 wt.%), nickel (1 2 wt.%), and traces of zinc. Table 20.7 summarizes the advantage and disadvantages of these wires. To the best of our knowledge, archwires made of gold alloys are not currently used in the clinical orthodontic practice. 20.3.2.2 Stainless steel Ever since it was introduced in 1929 by Wilkinson, SS forms the basis of most orthodontic wires. The composition of the SS archwires is similar to that used in the manufacturing of brackets with the austenitic form, especially the 18-8 SS, being the mostly commonly used. The use of SS wires is dependent on the technique, the stage of treatment, and the required stiffness (aligning, retraction to close space, finishing, etc.). It should be noted that aside from manufacturing brackets and archwires, SS is used to fabricate orthodontic band, auxiliaries (e.g., lingual buttons, lingual cleats), ligature wires, and orthodontic pliers. Table 20.8 describes, in general, the advantages and disadvantages of SS wires. Table 20.7 Advantages and disadvantages of gold wires. Advantages G G G G G G Disadvantages Low modulus of elasticity Strength can be increased by heat treatment or cold working High formability Excellent biocompatibility Good environmental stability Can be soldered or welded G G G Low springback Low yield strength Relatively expensive Table 20.8 Advantages and disadvantages of stainless steel (SS) wires. Advantages G G G G G G G G G High stiffness High resilience Good formability High yield strength Adequate springback Relatively economical Good joinability (i.e., can be welded and soldered) Corrosion resistant (except at weld and solder sites) Biocompatible Disadvantages G G G G G Lower springback than NiTi alloys Higher modulus of elasticity More frequent activations are required to maintain the same force levels Soldering is demanding Delivers high forces, relatively lower springback and high deflection rate, therefore requires more attention to detail to avoid harmful/unwanted effects on teeth (smaller diameter archwires are recommended when using SS wires in the initial stages of aligning crowded teeth)
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 557 Australian stainless steel wires These wires were developed by an Australian metallurgist, Arthur J Wilcock to be used for the Begg light archwire technique. His aim was to produce SS wires that are more tensile and thin enough to exhibit optimal tooth moving forces for an adequate period over long distances, with minimal loss in force intensity. These wires are round in cross section and are available as spools as well as in various straight lengths, grades, and diameters. Over time, the diameter of these wires has been decreasing to accommodate the changing concepts of using lower force levels in orthodontics. The wires that are manufactured using the pulse straightening technique (e.g., newer premium and supreme wires) demonstrate high tensile strength, high load deflection rate (i.e., high working range) good recovery patterns, and a relatively smoother surface finish. Braided or twisted wires (or multistranded stainless steel wires) In these types of wires, SS wires with relatively small diameters (e.g., 0.178 mm) are braided or twisted to fabricate larger wires with round or rectangular cross sections. The resulting larger-diameter wires are capable of sustaining larger elastic deflections with lower load deflection rates when compared with the conventional round SS wires of similar dimensions. On the other hand, the springback of multistranded SS wires is larger than that of conventional SS or β-Ti wires but similar to that of nitinol archwires. It should be noted that the springback is independent of the thickness of twisted archwires. 20.3.2.3 Cobalt chromium-based archwires (also known as Elgiloy) A major disadvantage of SS wires is the inability for the clinician to alter their resilience if required, since it cannot undergo heat hardening treatment. In contrast, Elgiloy wires present somewhat similar properties to those of SS wires but can be supplied in the softer and more formable form which can then hardened by heat treatment to modify their resilience or increase the strength of the wire. Elgiloy or cobalt chromium-based wires were originally developed by the Elgin Watch company for manufacturing watch springs and were introduced to orthodontics in the mid-1950s. Cobalt is the main constituent of these wires, which is present as a solid solution with chromium. Chromium provides the required stiffness and corrosion resistance, whereas nickel strengthens the alloy and improves its ductility. Other elements including iron, copper, molybdenum, tungsten are effective solid solutions hardeners, while manganese and silicon act as oxide scavengers (Table 20.9 shows the composition of Elgiloy wires). There are four main categories of Elgiloy archwires, separated by color coding, which have the same composition but differ in their resilience and other properties by variations in the manufacturing process (Kapila and Sachdeva, 1989): G Blue Elgiloy (soft) is recommended when the wire needs substantial bending, welding, or soldering.
558 Advanced Dental Biomaterials Table 20.9 Composition of cobalt chromium based archwires (Elgiloy). G G Element Weight % Co Cr Ni Fe Mo Mn Be Others 40 20 15 15.4 7 2 0.4 0.5 Yellow Elgiloy (ductile and slightly harder than blue) can be welded or soldered. Yellow Elgiloy is recommended for preadjusted edgewise appliance that requires greater springback properties than those provided by blue Elgiloy. Moreover, Yellow Elgiloy can be used as provided by the manufacturer or can be heat treated, if greater resiliency or springback properties are needed. Green Elgiloy (semiresilient) and Red Elgiloy (resilient) can be plier manipulated before heat treating and recommended for use where adjustments will not be required after heat treating. The advantages of Elgiloy over SS wires include excellent corrosion and tarnish resistance, higher springback efficiency (greater than SS by 20%), and they functionally remain active for a longer duration when used as resilient spring (i.e., greater resistance to fatigue and distortion) (Kapila and Sachdeva, 1989). Elgiloy wires demonstrate a modulus of elasticity that is somewhat comparable or higher to that of SS, and thus, the amount of force delivered to teeth is twice and four times that of β-Ti and NiTi wires, respectively (Goldberg et al., 1983). Even though Elgiloy wires exhibit a 17% higher corrosion resistance than SS wires, the relatively high nickel content could present a biocompatibility problem. Also, the frictional resistance of Elgiloy is comparable or slightly higher than that of SS wires (Frank and Nikolai, 1980). Other disadvantages of Elgiloy wires include the need for heat treatment and the demanding soldering process. 20.3.2.4 Nickel titanium wires NiTi alloys were first developed in the 1960s by William F Buehler for a space program at the Naval Ordnance Laboratory and were marketed as nitinol, which is derived from the elements comprising the alloy and the name of the developing lab —nickel (Ni, 55%), titanium (Ti, 45%), and Naval Ordinance Laboratory (nol). Over the years, several modifications have added additional desired properties such as low stiffness, shape memory, and superelasticity. In addition, the high springback property of nitinol permits the delivery of low forces even at large deflection, which is particularly useful in cases with severe crowding. Moreover, newer NiTi-based alloys are thermally activated (i.e., deliver tooth moving forces only after reaching
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 559 Table 20.10 Advantages and disadvantages of NiTi wires. Advantages G G G G High elasticity High springback High stored energy Biocompatible and environmentally stable (However, some corrosion and failure have been observed in certain occasions) Disadvantages G G G G G G Higher friction than SS but less than TMA wires (friction of NiTi wires can be reduced by coating the surface using metal plating, chemical vapor deposition, ion implantation, etc.) Lower stiffness than SS and is not recommended to be used at the completion stages of orthodontic treatment Fractures easily when bent over a sharp edge Cannot be welded or soldered More expensive than SS wires Poor formability SS, stainless steel; TMA, titanium molybdenum alloy. a certain temperature), which was achieved by the incorporation of a small amount of copper. Table 20.10 summarizes the advantages and disadvantages of NiTi wires. Thermoelasticity of NiTi alloys NiTi alloys are present in two crystalline forms: austenitic and martensitic. The temperature during which the alloy transitions from martensitic to austenitic is called the temperature transition range (TTR). The TTR is dependent on the nickel percentage; the lower the nickel content, the higher will be the TTR. Below the TTR, the crystals form in a martensitic form at low temperatures and high stresses and the crystals can exist in a distorted monoclinic, triclinic, or hexagonal structure. Above the TTR, the crystals form in an austenitic phase under high temperatures and low stresses and demonstrate an ordered bcc system structure. Some variations of the NiTi alloy have been developed in orthodontics. In general, the archwires are manufactured in the desired shape in the martensite phase and then they undergo the TTR to transform into the austenite form. In the austenite form, the NiTi wire is deformed to conform to the irregularities in the arch form. Different generations of NiTi alloys G Stabilized martensitic NiTi or first-generation M-NiTi These wires were marketed as nitinol by Unitek/3M. As the name implies, M-NiTi is manufactured through a work hardening process and stabilized in a martensitic crystallographic form; therefore they do not exhibit shape memory or superelastic properties as the phase transition does not occur. They also possess relatively high springback properties, low elastic modulus, and a wide working range when compared to SS wires, but exhibit poor formability.
560 G G Advanced Dental Biomaterials Superelastic NiTi or austenitic NiTi (A-NiTi) wires A-NiTi wires were introduced in orthodontics in the late 1980s and possess an active austenitic phase. Accordingly, they demonstrate superelasticity and good shape memory. Examples of this group include Chinese NiTi marketed as “NiTi” by Ormco and Japanese NiTi by Miura et al. (“Sentalloy” by GAC International) (Burstone et al., 1985; Miura et al., 1986). Thermoelastic NiTi or copper NiTi or martensitic active NiTi These thermally active wires exhibit a TTR close to the human body temperature and were introduced by Rohit Sachdeva in 1994. Copper is added in order to control the TTR of NiTi alloys as well as to reduce friction and increase strength. Generally, copper NiTi wires are composed of 42.99% Ti, 49.87% Ni, 0.5% Cr, and 5.6% Cu. These wires exploit the thermoelastic effect to achieve shape memory. The wire is cooled down into a martensitic form by means of an ice bath, cold water, or refrigerant spray to make it flexible before engagement into the brackets. At body temperature, the deformed wire transitions back to the austenitic form and its original structure, thus delivering light and continuous forces that are used to align crowded teeth. Moreover, copper NiTi wires possess excellent springback, are more resistant to permanent deformation, and deliver constant continuous forces over a long activation range. Four different types of copper NiTi wires are manufactured with variable transformations temperatures for a variety of applications: G G G G Type I (TTR of 15 C) is not suitable for clinical usage due to its high force levels. Type II (TTR of 27 C) produces heavy forces and it is best used where rapid tooth movement is required in patients with a normal periodontal health and average to high pain threshold. Type III (TTR of 35 C) generates relatively midrange forces and is recommended to be used in patients with a compromised periodontal health and low to normal pain threshold. Type IV (TTR of 40 C) is best used as an initial rectangular wire in patients who are periodontally compromised and are sensitive to pain. Multistranded or braided NiTi wires (Supercable NiTi) Braided NiTi exhibits increased superelastic properties as well as high deflection rate with very light forces, thus allowing these wires to be used as initial archwires for the alignment and leveling of moderate to severely crowded teeth. Plastic-coated NiTi The plastic coating improves the esthetic properties of the wires as they blend with the surrounding dentition and decrease the friction with ceramic, composite, and plastic wires. BioForce wires These newer NiTi wires exhibit superelastic shape memory properties that generate gradually increasing forces from anterior to posterior within the same dental arch. The same BioForce wire is manufactured with variable material composition or structure (i.e., variable transition temperatures) leading to the formation of a graded force delivery system throughout the arch length according to the periodontal surface of the tooth. It is claimed that these wires produce light forces (80 g) anteriorly and heavier forces posteriorly (300 g). G G G
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 561 20.3.2.5 Pure titanium wires and other titanium-based wires As mentioned earlier, pure Ti exists in two forms: α-phase and β-phase. α-Titanium wires are fabricated from grade V Ti that contains 6% aluminum and 4% vanadium. These wires possess good formability and exhibit stiffness comparable to that of SS. However, they demonstrate poor springback and their use is limited to torque control during the finishing stage of orthodontic treatment. β-Titanium or titanium molybdenum alloys Burstone and Goldberg developed β-Ti by alloying pure Ti with molybdenum in order to stabilize the crystals in β-phase at room temperature (Burstone and Goldberg, 1980; Goldberg and Burstone, 1979). β-Ti wires are commercially available as a titanium molybdenum alloy (TMA). The properties and some clinical applications of TMA wires are summarized in Box 20.4 (Burstone and Goldberg, 1980; Goldberg and Burstone, 1979; Cash et al., 2004). Titanium niobium wires Rohit Sachdeva introduced this alloy in orthodontics in 1995. Titanium niobium has a stiffness lower than that of TMA and SS wires but exhibits a springback comparable to that of SS. Furthermore, titanium niobium wires can welded and soldered and can be used to incorporate third-order corrections as well as used as a finishing archwire (Dalstra et al., 2000; Krishnan and Kumar, 2004). Box 20.4 Properties of titanium molybdenum alloy (TMA) wires G G G G G G Springback of TMA is twice that of SS wires with a relatively low deflection rate. This allows full bracket engagement with greater torque control than smaller SS wires. Good formability with adequate strength and springiness are ideal for fabrication of springs. Accordingly, TMA is recommended for fabricating closing loop archwires and auxiliary springs as well as intermediate and finishing archwires during treatment with preadjusted edgewise appliances. Modulus of elasticity of TMA is twice that of NiTi but less than that of SS. The low modulus of TMA combined with its good springback simplify the appliance design by eliminating the need of incorporating complex helices and loops in the archwires as compared to SS wires. TMA wires possess corrosion resistance comparable to that of SS and Elgiloy wires. TMA wires exhibit higher bracket/wire friction than either SS or Elgiloy wires and thus are not recommended for space closure with friction mechanics. However, newer ion-guarded TMA wires are available for decreasing the bracket/wire frictions. TMA wires can be used in all stages of orthodontic treatment from aligning and retraction to the final stages for torque expression.
562 Advanced Dental Biomaterials Timolium wires or titanium vanadium wires These wires are nickel free and can be used in patients with nickel hypersensitivity. Compared to TMA, timolium wires have a smoother surface which significantly reduces friction with the bracket and exhibits a higher yield strength with good formability which permits introducing bends in the wire without breakage. Moreover, the stiffness of timolium is lower than that of SS but higher than that of TMA. This section of the chapter has discussed the major types of metal wires with an overview of some clinical orthodontic applications. Table 20.11 provides a comparison between the most frequently used orthodontic wires. 20.4 Clear aligners Due to increasing demands for esthetic orthodontic treatment, the popularity of clear aligner therapy (CAT) has greatly increased over the years (Boyd, 2008). Surveys have shown that a significant number of patients do not find the commonly used metallic orthodontic appliances attractive and are willing to pay a higher fee for more esthetic options (Rosvall et al., 2009). Contemporary aligners of the 21st century combine the principles pioneered by Remensnyder (1926), Kesling (1945), Nahoum (1964), and others (Ponitz, 1971; McNamara et al., 1985; Sheridan et al., 1993) and integrate them with the modern CAD/CAM technology. The Invisaligns appliance by Align Technology is currently the most popular clear aligner system and it consists of a series of digitally generated thermoplastic aligners created using stereolithographic models (Wong, 2002). The main advantages of CAT include removability for ease of eating, comfort, and oral hygiene maintenance, improved esthetics, and possibly reduced periodontal risk (Boyd, 2008; Kravitz et al., 2009; Miethke and Brauner, 2007; Chhibber et al., 2018). In addition, its smaller size compared to traditional orthodontic appliances has led to improved patient acceptance of orthodontic treatment (Sergl and Zentner, 1998; Miller et al., 2007). The orthodontic force exerted by the thermoplastic appliance in CAT is highly dependent on the mechanical properties of its fabrication material (Zhang et al., 2011). To achieve the ideal treatment outcomes, it is important to understand the numerous clear aligner products available on the market, and how they differ regarding their construction material, thickness, and clinical protocol. 20.4.1 Material composition Appliances used for CAT are mainly based on transparent thermoplastic-based resin polymeric materials. The ideal material properties would be large springback, high stored energy, tolerance to the oral environment, biocompatible, and low surface roughness (Barone et al., 2016). Various types of aligner materials exist, such as polyethylene terephthalate (PET-G), thermoplastic PCs, polypropylene, polycarbonate, and ethylene-vinyl acetate (Table 20.12) (Gao and Wichelhaus, 2017).
Table 20.11 Comparison of the clinical characteristics of metal orthodontic wires. Wire type Stiffness Stored energy Friction with SS brackets Springback Formability Biocompatibility and environmental stability Joinability Stainless steel High Low Low Low Good Good Cobalt chromium High Low Low to moderate Low Good Good Nickel titanium Low High High Poor Goodc β-Titanium Multistranded NiTi Average Low Average High Low to moderate High NA Can be welded and soldered Can be weldeda and solderedb Not joinable Average High Good Poor Good Good SS, Stainless steel. a Only yellow and blue Elgiloy can be welded. b Solder with some difficulty. c Some corrosion and failure have been noted. Can be welded Can be welded and soldered
564 Advanced Dental Biomaterials Table 20.12 Materials used for thermoplastic aligners. Material Product(s) Manufacturer(s) Polyethylene Hardcast Biolon Duran, Clear Aligner Erkodur Zendura Invisalign Copyplast Imprelon “S” Bioplast Essix A 1 Scheu-Dental, GmbH Dreve Dentamid Scheu-Dental, GmbH Erkodent Bay Materials, LLC Align Technology, Inc. Scheu-Dental, GmbH Scheu-Dental, GmbH Scheu-Dental, GmbH Dentsply Raintree Essix, Inc. Ortho Organisers Inc. Keystoen Industries Dentsply Raintree Essix, Inc. PET-G Polyurethane Polypropylene Polycarbonate Ethylene-vinyl acetate Copolyester Polypropylene/polyethylene copolymer Masel Keystone Essix C 1 PET-G, Polyethylene terephthalate glycol. Source: Adapted from Ryokawa, H., Miyazaki, Y., Fujishima, A., Miyazaki, T., Maki, K. 2006. The mechanical properties of dental thermoplastic materials in a simulated intraoral environment. Orthod. Waves 65, 64 72. The current most commonly used thermoplastic polymers for aligners are based on derivatives of PC and polypropylene. PET-G is a noncrystallizing amorphous copolymer of PET, which has good mechanical properties, formability, optical properties, fatigue resistance, and dimensional stability. Polyurethanes have a wide gamut of applications beyond the scope of dentistry, ranging from engineering materials to coatings to adhesives and films (Frick and Rochman, 2004). It combines the properties of vulcanized rubber with the processability of thermoplastic polymers to provide improved physical properties with chemical and abrasion resistance (Lu and Macosko, 2004). When Invisalign was first introduced to the market in 1999, it was made from a single-layered rigid PC material, synthesized from methylene diphenyl diisocyanate and 1,6-hexanediol (Morton et al., 2017). Subsequent generations of aligners were made from Exceed-30, which were selected for improved material transparency, flexibility, and breakage resistance (Brantley and Eliades, 2017). In 2013 SmartTrack was introduced as the latest generation of thermoplastic material, designed to replace the previous Exceed-30 and Exceed-40 aligner materials. SmartTracks is a multilayered thermoplastic-PC-based material combined with an elastomeric component (Align Technology, 2013, 2017). Based on the manufacturer’s claims, this material can maintain more constant and lighter forces over time when compared to previous generations, without losing force expression due to stress relaxation cycles throughout its 2 weeks of usage (Morton et al., 2017). There are currently no scientific studies available on the stress relaxation properties of SmartTrack. Preliminary patient surveys suggest that SmartTrack may be more comfortable than the older materials (Brascher et al., 2016), although further studies
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 565 are needed to verify this (White et al., 2017). A recent clinical study has showed higher mean orthodontic tooth movement achieved by SmartTrack when compared to Exceed-30 over a 25-day period (Wheeler et al., 2017). 20.4.2 The thermoforming process The two basic ways for the creation of clear aligners are the vacuum-formed method initially described by Nahoum (1964) and the positive air pressure method described by Ponitz and McNamara (McNamara et al., 1985; Ponitz, 1971). In the vacuum-form method, a vacuum device is used to apply suck-down pressure as the heated sheet of thermoplastic material is being applied to the dental cast, under atmospheric pressure (Ponitz, 1971). Although a theoretical maximum of 15 PSI is expected, the actual pressure varies between 3 and 14 PSI, based on the efficiency of the vacuum pump (Nahoum, 2014). The positive air pressure method utilizes a pressure chamber to apply positive air pressure, rather than partial vacuum pressure, to conform the softened sheet of thermoplastic material to the shape of the cast (Hahn et al., 2009, 2011). In this system, the air can be compressed up to 100 PSI and the details of the inner surfaces of the aligner improve as the air pressure increases, up to a certain level (Nahoum, 2014). Studies have shown that vacuum-formed aligners deliver lower forces on average than clear aligners created under high-pressure methods (Hahn et al., 2009). It has been hypothesized (Nahoum, 2014) that the improved fit from the high-pressure method may lead to increased friction and; therefore, allows more expression of the programmed forces (Hahn et al., 2009). 20.4.3 Forces of thermoplastic aligners The ideal clear aligner material should be capable of delivering sufficient force to create the desired orthodontic tooth movement in a programmed and predictable fashion, while avoiding damage the surrounding periodontal tissues (Lombardo et al., 2017). In addition, clear aligners should be able to apply light and continuous forces on the teeth, while still providing enough stiffness with a high yield strength to ensure that the force is applied within the elastic range of the material (Lombardo et al., 2017). 20.4.3.1 Differences in force generation between clear aligner therapy and fixed appliances Orthodontic tooth movement is a biochemical process provided by a loading system, represented by a combination of forces and moments applied to the dentition (Ren et al., 2003, 2004). The load delivering modality depends on the specific type of orthodontic appliance used. In traditional fixed appliances, the brackets and bands bonded directly to the crowns of the teeth provide a highly defined contact area. The stained applied is transferred to the center of the bracket
566 Advanced Dental Biomaterials and the subsequent moment-to-force ratio can be defined with respect to the bracket and the center of resistance of the tooth. With CAT, the force system is created through a small geometrical mismatch between the shape of the thermoplastic aligner and the patients’ dentition (dental alignment). Relative to fixed appliances, the load in clear aligner therapy is transferred through a wider and lesser defined contact area. Facilitation of the load transfer can be achieved through usage of auxiliary elements, such as attachment or divots, which modify the contact interface between the aligner and the dentition. This could give rise to localized increases in a mismatch at specific areas of need, as well as improved accuracy in load control in terms of intensity and direction (Barone et al., 2016). Generally, the generation of forces in CAT can be categorized into two systems: (1) a displacement-driven system and (2) a force-driven system. Displacementdriven system is an older concept with limited usefulness today (Schupp and Haubrich, 2016). It involves designing aligners that are formed to the geometry of the dentition in the next stage of the virtual treatment plan (Schupp and Haubrich, 2016). It is considered that the tooth will move until it is lined up to the shape of the aligner. This concept of creating geometric mismatches is effective for simple movements, such as when tipping and intrusion are required, but it is less effective for achieving control of root movements. This concept can still be used today for cases requiring simple and limited movements, via the use of Essix retainers (Sheridan et al., 1993). In the force-driven system, creation of aligners is determined by the fundamental principles of biomechanics. This concept forms aligners in the shapes required to produce the desired forces and movements, not necessarily the shape of the next stage of tooth movement. The aligners are shaped with the intention to deliver the specific forces and movements that can result in the desired tooth movement. This concept is often seen in the comprehensive aligner systems, such as ClearCorrects and Invisaligns. 20.4.3.2 Material factors affecting force delivery Orthodontic forces exerted by clear aligners are governed by numerous factors, including tray thickness, amount of activation as predicted by the software, extension of the gingival edge, type of movement, as well as stiffness and other material properties (Baldwin et al., 2008; Bollen et al., 2003; Clements et al., 2003). Appliances with thicker material have been showed to apply significantly greater orthodontic forces than those of thinner material (Gao and Wichelhaus, 2017; Hahn et al., 2009, 2011; Kohda et al., 2013). The material hardness and elastic modulus were also strongly correlated with the amount of force delivered (Kohda et al., 2013). Furthermore, aligners with larger gingival edge extensions (3 mm) produce more force than aligners with a minimal gingival edge extension (0 1 mm) (Gao and Wichelhaus, 2017). It has been thought that the observed increase in force might be due to enhanced stiffness caused by the changes in material shape.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 567 Aligners with a 2 mm extension beyond the gingival margins have been shown to be more retentive in nature than straight or scalloped edges (Cowley et al., 2012). However, further studies are necessary to investigate the effect of the increased gingival edge extensions on periodontal tissues. Finally, thermoplastic aligners are viscoelastic materials by nature; consequently, the forces generated by its programmed movements will decrease as a function of time (Zhang et al., 2011). The following section will discuss the current progress in seeking aligner materials with the most linear elastic behavior and high yield strength for an efficacious force output (Lombardo et al., 2017; Zhang et al., 2011). 20.4.4 Mechanical properties In general, the two main factors that influence the mechanical properties of thermoplastic aligners are structural and environmental (Landel and Nielsen, 1993). Structural factors refer to properties related to the molecular and crystal structures of the polymer material. For example, PET-G, PCs, and polycarbonates are amorphous plastics, while polyethylenes and polypropylenes are crystalline plastics. The degree of crystallinity of these thermoplastic aligners affects its mechanical properties, and ultimately, the nature of the force that can be delivered by the aligner tray (Ryokawa et al., 2006). Environmental factors refer to the various chemical and mechanical changes that aligners could undergo once inserted into the oral cavity. Variables such as temperature, humidity, pressure, heat, and salivary enzymes can lead to chemical reactions that alter the physical properties of aligners (Hahn et al., 2011; Ryokawa et al., 2006). Furthermore, aligners are also subject to both short- and long-term mechanical forces, while they are worn by the patient (Tuncay, 2006). As the clear aligner is being inserted over the patient’s dentition, it may experience immediate deformation during the fitting process, but once fully seated, it is able to rebound to its original shape due to its elasticity. During the period of prescribed appliance usage, the aligners are further subjected to occlusal forces intermittently exerted throughout the day and night (Zhang et al., 2011). 20.4.4.1 Elastic modulus The elastic modulus (Young’s modulus or modulus of elasticity) is a measurement for the stiffness of a given material and is calculated as a ratio of the elastic stress to elastic strain (Anusavice et al., 2013). It is advantageous for aligners to be made from thermoplastic polymers of high elastic modulus, as it increases its force delivery capability under constant strain (Alexandropoulos et al., 2015). In addition, higher elastic modulus allows the same force delivery from materials of thinner dimensions (Kohda et al., 2013). Polyurethanes of the Invisalign system exhibit a higher elastic index than polyester and PET-G materials (Alexandropoulos et al., 2015).
568 Advanced Dental Biomaterials 20.4.4.2 Creep resistance Creep is defined as the propensity of a solid material to gradually undergo permanent deformation under the influence of a static load or constant stress (Anusavice et al., 2013). With respect to clear aligner therapy, it is desirable for the aligner material to have increased resistance to creep, as it allows for optimal expression of the orthodontic force (Brantley and Eliades, 2017). Having a higher creep value implies that under the constant forces of occlusal stress by the opposing dentition, the material is more likely to deform and exhibit attenuated force levels. Under indentation testing, the proprietary PC material of Invisalign demonstrated inferior creep resistance than other PET-G thermoplastic materials (Alexandropoulos et al., 2015). 20.4.4.3 Stress relaxation Stress relaxation describes the material’s tendency to decrease its load generation when held under a constant strain or deflection (Anusavice et al., 2013). An orthodontic aligner that exhibits rapid stress relaxation may express a decreasing amount of orthodontic force once inserted intraorally, even before the tooth has begun to move (Lombardo et al., 2017). This effect will undoubtedly have a negative impact on the aligner’s performance and could be compounded when used in a sequential fashion, leading to a tracking problem clinically. The ideal aligner should exhibit a stress relaxation curve that is fairly flat, representing its capability to exert constant and continuous forces over time (Fang et al., 2013). Unfortunately, stress relaxation curves for current aligner materials generally follow a pattern of rapid decay within the first 8 hours of application, then diminish to a plateau thereafter (Lombardo et al., 2017). This highlights the importance of measuring forces exerted by aligners not only during the first hours after tray placement inside the oral cavity but also within the first 24- and 48-hours (Lombardo et al., 2017). Both PET-G and PET polymers have a similar stress strain behavior, although PET-G exhibits greater elastic recovery than PET polymers (Dupaix and Boyce, 2005). Furthermore, stress relaxation appears to be temperature dependent, as intraoral temperatures can accelerate this process as compared to ambient room temperatures (Fang et al., 2013). Similar results were found in an in vivo study of von Mises strains experienced by aligners, a parameter used for assessing the amount of load that can be applied before material failure (Jones, 2009). The aligners demonstrated peak von Mises strain within the first 24 hours of wear, followed by a rapid decrease on day 2 and subsequent plateauing throughout the remaining 2 weeks of wear (Vardimon et al., 2010). This result further highlights the importance of wearing the aligners maximally within the first 24-hour period (Brantley and Eliades, 2017). 20.4.4.4 Water absorption The mechanical properties of thermoplastic aligners can be affected by absorption of water through the humidity in the air as well the oral cavity environment
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 569 (Brantley and Eliades, 2017). This could result in dimensional changes to the aligner trays due to hygroscopic expansion, and thus leading to changes in the tray fit to the dentition and force delivery system (Kwon et al., 2008). Moreover, it should be noted that thermoplastic materials with a crystalline structure have a lower water absorption rate than amorphous structures, due to their higher molecular density (Ryokawa et al., 2006). 20.4.4.5 Wear, abrasion, and intraoral aging Assessment of aligner wear is a challenging and complex process, as in vitro experiments may not be able to accurately represent the numerous simultaneous mechanisms that lead to wear in vivo (Lutz et al., 1984; Wassell et al., 1994). Laboratory studies simulating wear on polypropylene materials, including Essix C 1 , and PET-G polymer, demonstrated an increased wear resistance for the latter (Gardner et al., 2003). Furthermore, the increased wear resistance of PET-G polymers has been attributed to its increased hardness . Intraorally used aligners can undergo significant structural deformation, including cracking, wear of contact points, and adsorption of proteinaceous materials (Schuster et al., 2004). These changes can lead to a degradation of its mechanical properties, including elasticity, hardness, and indentation creep (Gerard Bradley et al., 2016), leading to attenuation of its force delivery capacity. These changes were thought to be attributed to the alterations of the polymer crystallinity from cold working produced by the masticatory forces (Gerard Bradley et al., 2016). Similar results were found in PET-G aligners, including increased material stiffness, hardness, and surface roughness (Ahn et al., 2015). However, current studies still differ regarding whether clinically worn Invisalign appliances will undergo changes in their chemical composition (Schuster et al., 2004; Gracco et al., 2009) or remain stable throughout intraoral aging (Gerard Bradley et al., 2016). 20.4.5 Attachments Attachments are part of the armamentarium in clear aligner therapy for achieving the desired tooth movement by improving accuracy and delivering higher levels of forces (Rossini et al., 2015; Simon et al., 2014). Attachments are made out of dental composite materials, including both flowable and packable, which are then polymerized when cured on the crown surface using adhesives (Table 20.13) (Barone et al., 2016). Attachments provide an artificial, flat, and planar surface onto which the aligner can apply additional strains. This portion of the attachment is known as the “active surface” (Schupp and Haubrich, 2016) (Fig. 20.3). The active surface is of critical importance, as it is the one being engaged by the aligner. The remaining portion of the attachment material solely serves to keep the active surface in its desired location. Force systems are applied at specific locations by designing the aligner to only contact at the active surface while avoiding contact at the other areas of the
570 Advanced Dental Biomaterials Table 20.13 Recommendation for composite materials used for attachments in clear aligner therapy. Composite material Compatible adhesive Manufacturer(s) 3M Filtek Supreme Plus & Z250 Tetric EvoCeram 3M Adper prompt L-Pop G-ænial Bond SonicFill Venus Pearl & Diamond GrandioSO Optibond XTR iBond Total Etch Futurabond U 3M ESPE Ivoclar Vivadent GC America Kerr Dental Heraeus-Kulzer VOCO GmbH Packable Flowable Filtek Supreme Ultra Flowable GrandioSO Heavy Flow Tetric EvoFlow Scotchbond Universal Adhesive Futurabond U G-ænial Bond G-ænial Universal Flo G-ænial Bond 3M ESPE VOCO GmbH Ivoclar Vivadent GC America GC America Source: Adapted from Align Technology, 2018. List of recommended materials for bonding attachments. Available from: ,https://learn.invisalign.com/attachments/bonding/recommendedmaterials. (accessed 26.05.18.) (Align Technology, 2018). Figure 20.3 Example of a horizontal rectangular attachment placed on a maxillary central incisor, with bevel toward the gingival aspect, serving as the active surface (arrow). attachment. This design allows for better control of the direction and magnitude of forces applied, leading to more accurate force delivery and enhanced control of tooth movement (Schupp and Haubrich, 2016). Attachments are typically not meant to be fully engaged at the initial time of bonding, as the patient progresses through the aligner trays, their successive aligners gradually engage the active surfaces of the attachments more while
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 571 applying constant strain sresulting in tooth movement. Eventually, the attachment will finally fill in the slot on the aligner. This concept is similar to working through the archwires in preadjusted fixed appliances where the slot of the bracket is filled by larger wires (Hennessy and Al-Awadhi, 2016). 20.4.6 Cytotoxicity Isocyanates are a family of chemicals commonly used for the commercial synthesis of PC-based products; this includes certain types of clear aligners, such as the Invisalign system (Wisnewski et al., 2000). Its popularity in this industry is due to its ability to efficiently react with polyols to form PCs with virtually no by-products (Fig. 20.4). However, the usage of isocyanates does not come without its disadvantages as well, including numerous side effects that can pose significant health issues to the patient (Liljelind et al., 2010). Isocyanates are the leading cause of occupational asthma in industrialized nations. They also cause hypersensitivity and allergic reactions, as well as mucous membrane irritation (Liljelind et al., 2010). It is thought that the cytotoxic properties of thermoplastic materials appear to be influenced by the composition and structure of the polymer, as well as processing and environmental factors, such as temperature, humidity, pressure, and thermal history (Jorge et al., 2003). Polymer materials, especially amorphous plastic including Invisaligns, demonstrate high rates of water absorption, allowing long-term leaching of unreacted chemicals into the oral environment (Ryokawa et al., 2006). This characteristic is of particular concern in orthodontics, as patients are usually prescribed to wear aligners throughout the day and night (Boyd et al., 2000). If a PC-based aligner is used 24 hours a day, they are continuously in contact with the dentition and gingiva. Any residual isocyanates may potentially leak and lead to damage of the epithelial layer or loss of its integrity, leading to various immunologic reactions (Premaraj et al., 2014a,b, Wu, 2014). Figure 20.4 The chemical reaction for the synthesis of polyurethanes from isocyanates.
572 Advanced Dental Biomaterials Awosika et al. (2017) reported the case of a 23-year-old healthy Caucasian female undergoing Invisalign treatment, who experienced intermittent urticarial on her extremities and flanks 2 days after application of the aligners. Her symptoms subsequently worsened, leading to facial and periorbital swelling, with erythema and burning sensation of the lips and oral mucosa. Her treatment was the removal of the aligners with administration of oral prednisone, which resolved her symptoms. Patch testing revealed strong positive reactions to Invisalign and questionable reactions to isocyanates after 96 hours. These findings were consistent with sensitivity to isocyanates and contact sensitization to Invisalign. 20.5 Final remarks Improvements in the orthodontic materials resulted in a great influence on the orthodontic practice, with prominent changes in the types of materials used and biomechanics research. Metals, alloys, polymers, and composites are the materials used for manufacturing orthodontic materials. Each type of material presents its advantages over the other. Accordingly, the clinician should possess a thorough knowledge of the physical and mechanical properties of these materials to determine their clinical behavior and to achieve a satisfactory and predictable outcome that fits the treatment objectives. The rise of the digital age has transformed orthodontics as we know it. The advent of computer-aided design and computer-aided manufacturing has brought clear aligner therapy to the masses and became a mainstay of contemporary orthodontic. Furthermore, digital advances may have a great impact on the materials used for fabricating brackets, wires, and clear aligners. Therefore, it is more important than ever that we use the best available evidence to make sound clinical decisions. References Abolfathi, A., Chen, J.C., Li, C., Tricca, R.E., Wu, B.M., 2009. Systems and Methods for Dental Appliance Compliance Indication. Google Patents. Ahn, H.W., Ha, H.R., Lim, H.N., Choi, S., 2015. Effects of aging procedures on the molecular, biochemical, morphological, and mechanical properties of vacuum-formed retainers. J. Mech. Behav. Biomed. Mater. 51, 356 366. Alexandropoulos, A., Al Jabbari, Y.S., Zinelis, S., Eliades, T., 2015. Chemical and mechanical characteristics of contemporary thermoplastic orthodontic materials. Aust. Orthod. J. 31, 165 170. Ali, S., Mann, D.A., 2004. Signal transduction via the NF-kappa B pathway: a targeted treatment modality for infection, inflammation and repair. Cell Biochem. Funct. 22, 67 79. Align Technology, 2013. Align Technology Announces January 21st Availability of SmartTrack(TM) Invisalign(R) Aligner Material. Align Technology. Align Technology, 2017. Align Technology Receives U.S. Patents for SmartTrack(R) Invisalign(R) Aligner Material. Align Technology.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 573 Align Technology, 2018. List of recommended materials for bonding attachments. Available from: ,https://learn.invisalign.com/attachments/bonding/recommendedmaterials. (accessed 26.05.18). Ansari, M.Y., Agarwal, D.K., Gupta, A., Bhattacharya, P., Ansar, J., Bhandari, R., 2016. Shear bond strength of ceramic brackets with different base designs: comparative invitro study. J. Clin. Diagn. Res. 10, ZC64 ZC68. Anusavice, K.J., Shen, C., Rawls, H.R., 2013. Phillips’ Science of Dental Materials. Elsevier Health Sciences. Arash, V., Rabiee, M., Rakhshan, V., Khorasani, S., Sobouti, F., 2015. In vitro evaluation of frictional forces of two ceramic orthodontic brackets versus a stainless steel bracket in combination with two types of archwires. J. Orthod. Sci. 4, 42 46. Arici, S., Regan, D., 1997. Alternatives to ceramic brackets: the tensile bond strengths of two aesthetic brackets compared ex vivo with stainless steel foil-mesh bracket bases. Br. J. Orthod. 24, 133 137. Arici, N., Akdeniz, B.S., Arici, S., 2015. Comparison of the frictional characteristics of aesthetic orthodontic brackets measured using a modified in vitro technique. Korean J. Orthod. 45, 29 37. Awosika, O., Kao, S., Rengifo-Pardo, M., Ehrlich, A., 2017. Angioedema, stomatitis, and urticaria caused by contact allergy to Invisalign. Dermatitis 28, 323 324. Baldwin, D.K., King, G., Ramsay, D.S., Huang, G., Bollen, A.M., 2008. Activation time and material stiffness of sequential removable orthodontic appliances. Part 3: Premolar extraction patients. Am. J. Orthod. Dentofacial Orthop. 133, 837 845. Barone, S., Paoli, A., Razionale, A.V., Savignano, R., 2016. Computational design and engineering of polymeric orthodontic aligners. Int. J. Numer. Method Biomed. Eng. 33, e2839. Birnie, D., 1990. Ceramic brackets. Br. J. Orthod. 17, 71 74. Bishara, S.E., Fehr, D.E., Jakobsen, J.R., 1993. A comparative study of the debonding strengths of different ceramic brackets, enamel conditioners, and adhesives. Am. J. Orthod. Dentofacial Orthop. 104, 170 179. Bollen, A.M., Huang, G., King, G., Hujoel, P., Ma, T., 2003. Activation time and material stiffness of sequential removable orthodontic appliances. Part 1: Ability to complete treatment. Am. J. Orthod. Dentofacial Orthop. 124, 496 501. Boyd, R.L., 2008. Esthetic orthodontic treatment using the Invisalign appliance for moderate to complex malocclusions. J. Dent. Educ. 72, 948 967. Boyd, R.L., Miller, R.J., Vlaskalic, V., 2000. The Invisalign system in adult orthodontics: mild crowding and space closure cases. J. Clin. Orthod. 34, 203 212. Brantley, W.A., Eliades, T., 2001. Orthodontic Materials: Scientific and Clinical Aspects. Thieme. Brantley, W.A., Eliades, T., 2017. Orthodontic Applications of Biomaterials: A Clinical Guide, 2016, Elsevier Science. Brascher, A.K., Zuran, D., Feldmann J.R., R.E., Benrath, J., 2016. Patient survey on Invisalign(R) treatment comparen the SmartTrack(R) material to the previous aligner material. J. Orofac. Orthop. 77, 432 438. Burstone, C.J., Goldberg, A.J., 1980. Beta titanium: a new orthodontic alloy. Am. J. Orthod. Dentofacial Orthop. 77, 121 132. Burstone, C.J., Qin, B., Morton, J.Y., 1985. Chinese NiTi wire—a new orthodontic alloy. Am. J. Orthod. 87, 445 452. Cash, A., Curtis, R., Garrigia-Majo, D., Mcdonald, F., 2004. A comparative study of the static and kinetic frictional resistance of titanium molybdenum alloy archwires in stainless steel brackets. Eur. J. Orthod. 26, 105 111.
574 Advanced Dental Biomaterials Chhibber, A., Agarwal, S., Yadav, S., Kuo, C.L., Upadhyay, M., 2018. Which orthodontic appliance is best for oral hygiene? A randomized clinical trial. Am. J. Orthod. Dentofacial Orthop. 153, 175 183. Choi, S.H., Kang, D.Y., Hwang, C.J., 2014. Surface roughness of three types of modern plastic bracket slot floors and frictional resistance. Angle Orthod. 84, 177 183. Clements, K.M., Bollen, A.M., Huang, G., King, G., Hujoel, P., Ma, T., 2003. Activation time and material stiffness of sequential removable orthodontic appliances. Part 2: Dental improvements. Am. J. Orthod. Dentofacial Orthop. 124, 502 508. Cowley, D.P., Mah, J., O’toole, B., 2012. The effect of gingival-margin design on the retention of thermoformed aligners. J. Clin. Orthod. 46, 697 702. quiz 705. Dalstra, M., Denes, G., Melsen, B., 2000. Titanium-niobium, a new finishing wire alloy. Clin. Orthod. Res. 3, 6 14. Dobrin, R.J., Kamel, I.L., Musich, D.R., 1975. Load-deformation characteristics of polycarbonate orthodontic brackets. Am. J. Orthod. 67, 24 33. Dupaix, R.B., Boyce, M.C., 2005. Finite strain behavior of poly(ethylene terephthalate) (PET) and poly(ethylene terephthalate)-glycol (PETG). Polymer 46, 4827 4838. Elekdag-Turk, S., Isci, D., Ozkalayci, N., Turk, T., 2009. Debonding characteristics of a polymer mesh base ceramic bracket bonded with two different conditioning methods. Eur. J. Orthod. 31, 84 89. Eliades, T., 2007. Orthodontic materials research and applications: Part 2. Current status and projected future developments in materials and biocompatibility. Am. J. Orthod. Dentofacial Orthop. 131, 253 262. Eliades, T., Brantley, W.A., 2016. Orthodontic Applications of Biomaterials: A Clinical Guide. Elsevier Science. Eliades, T., Zinelis, S., Eliades, G., Athanasiou, A.E., 2002. Nickel content of as-received, retrieved, and recycled stainless steel brackets. Am. J. Orthod. Dentofacial Orthop. 122, 217 220. Eliades, T., Zinelis, S., Eliades, G., Athanasiou, A.E., 2003. Characterization of as-received, retrieved, and recycled stainless steel brackets. J. Orofac. Orthop. 64, 80 87. Eliades, T., Gioka, C., Zinelis, S., Eliades, G., Makou, M., 2004. Plastic brackets: hardness and associated clinical implications. World J. Orthod. 5, 62 66. Faltermeier, A., Rosentritt, M., Faltermeier, R., Muessig, D., 2007. Influence of fibre and filler reinforcement of plastic brackets: an in vitro study. Eur. J. Orthod. 29, 304 309. Fang, D., Zhang, N., Chen, H., Bai, Y., 2013. Dynamic stress relaxation of orthodontic thermoplastic materials in a simulated oral environment. Dent. Mater. J. 32, 946 951. Feldner, J.C., Sarkar, N.K., Sheridan, J.J., Lancaster, D.M., 1994. In vitro torque-deformation characteristics of orthodontic polycarbonate brackets. Am. J. Orthod. Dentofacial Orthop. 106, 265 272. Flores, D.A., Choi, L.K., Caruso, J.M., Tomlinson, J.L., Scott, G.E., Jeiroudi, M.T., 1994. Deformation of metal brackets: a comparative study. Angle Orthod. 64, 283 290. Frank, C.A., Nikolai, R.J., 1980. A comparative study of frictional resistances between orthodontic bracket and arch wire. Am. J. Orthod. 78, 593 609. Frick, A., Rochman, A., 2004. Characterization of TPU-elastomers by thermal analysis (DSC). Polym. Test. 23, 413 417. Gao, L., Wichelhaus, A., 2017. Forces and moments delivered by the PET-G aligner to a maxillary central incisor for palatal tipping and intrusion. Angle Orthod. 87, 534 541. Gardner, G.D., Dunn, W.J., Taloumis, L., 2003. Wear comparison of thermoplastic materials used for orthodontic retainers. Am. J. Orthod. Dentofacial Orthop. 124, 294 297.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 575 Gerard Bradley, T., Teske, L., Eliades, G., Zinelis, S., Eliades, T., 2016. Do the mechanical and chemical properties of InvisalignTM appliances change after use? A retrieval analysis. Eur. J. Orthod. 38, 27 31. Gioka, C., Bourauel, C., Zinelis, S., Eliades, T., Silikas, N., Eliades, G., 2004. Titanium orthodontic brackets: structure, composition, hardness and ionic release. Dent. Mater. 20, 693 700. Goldberg, J., Burstone, C.J., 1979. An evaluation of beta titanium alloys for use in orthodontic appliances. J. Dent. Res. 58, 593 599. Goldberg, A.J., Morton, J., Burstone, C.J., 1983. The flexure modulus of elasticity of orthodontic wires. J. Dent. Res. 62, 856 858. Gracco, A., Mazzoli, A., Favoni, O., Conti, C., Ferraris, P., Tosi, G., et al., 2009. Short-term chemical and physical changes in Invisalign appliances. Aust. Orthod. J. 25, 34 40. Grimsdottir, M.R., Gjerdet, N.R., Hensten-Pettersen, A., 1992. Composition and in vitro corrosion of orthodontic appliances. Am. J. Orthod. Dentofacial Orthop. 101, 525 532. Gwinnett, A.J., 1988. A comparison of shear bond strengths of metal and ceramic brackets. Am. J. Orthod. Dentofacial Orthop. 93, 346 348. Haddad, A.C., Tortamano, A., Souza, A.L., Oliveira, P.V., 2009. An in vitro comparison of nickel and chromium release from brackets. Braz. Oral Res. 23, 399 406. Hahn, W., Dathe, H., Fialka-Fricke, J., Fricke-Zech, S., Zapf, A., Kubein-Meesenburg, D., et al., 2009. Influence of thermoplastic appliance thickness on the magnitude of force delivered to a maxillary central incisor during tipping. Am. J. Orthod. Dentofacial Orthop. 136, 12 13. 12 e1-7; discussion. Hahn, W., Engelke, B., Jung, K., Dathe, H., Kramer, F.J., Rodig, T., et al., 2011. The influence of occlusal forces on force delivery properties of aligners during rotation of an upper central incisor. Angle Orthod. 81, 1057 1063. Hamula, D.W., Hamula, W., Sernetz, F., 1996. Pure titanium orthodontic brackets. J. Clin. Orthod. 30, 140 144. Harzer, W., Bourauel, C., Gmyrek, H., 2004. Torque capacity of metal and polycarbonate brackets with and without a metal slot. Eur. J. Orthod. 26, 435 441. Hennessy, J., Al-Awadhi, E.A., 2016. Clear aligners generations and orthodontic tooth movement. J. Orthod. 43, 68 76. Jones, R.M., 2009. Deformation Theory of Plasticity. Bull Ridge Pub. Jorge, J.H., Giampaolo, E.T., Machado, A.L., Vergani, C.E., 2003. Cytotoxicity of denture base acrylic resins: a literature review. J. Prosthet. Dent. 90, 190 193. Kapila, S., Sachdeva, R., 1989. Mechanical properties and clinical applications of orthodontic wires. Am. J. Orthod. Dentofacial Orthop. 96, 100 109. Kapur, R., Sinha, P.K., Nanda, R.S., 1999. Frictional resistance in orthodontic brackets with repeated use. Am. J. Orthod. Dentofacial Orthop. 116, 400 404. Keith, O., Kusy, R.P., Whitley, J.Q., 1994. Zirconia brackets: an evaluation of morphology and coefficients of friction. Am. J. Orthod. Dentofacial Orthop. 106, 605 614. Kesling, H.D., 1945. The philosophy of the tooth positioning appliance. Am. J. Orthod. Oral Surg. 31, 297 304. Kohda, N., Iijima, M., Muguruma, T., Brantley, W.A., Ahluwalia, K.S., Mizoguchi, I., 2013. Effects of mechanical properties of thermoplastic materials on the initial force of thermoplastic appliances. Angle Orthod. 83, 476 483. Kravitz, N.D., Kusnoto, B., Begole, E., Obrez, A., Agran, B., 2009. How well does Invisalign work? A prospective clinical study evaluating the efficacy of tooth movement with Invisalign. Am. J. Orthod. Dentofacial Orthop. 135, 27 35.
576 Advanced Dental Biomaterials Krishnan, V., Kumar, K.J., 2004. Weld characteristics of orthodontic archwire materials. Angle Orthod. 74, 533 538. Kumar, B.S., Miryala, S., Kumar, K.K., Shameem, K., Regalla, R.R., 2014. Comparative evaluation of friction resistance of titanium, stainless steel, ceramic and ceramic with metal insert brackets with varying dimensions of stainless steel wire: an in vitro multicenter study. J. Int. Oral Health 6, 66 71. Kusy, R.P., 1988. Morphology of polycrystalline alumina brackets and its relationship to fracture toughness and strength. Angle Orthod. 58, 197 203. Kusy, R.P., O’Grady, P.W., 2000. Evaluation of titanium brackets for orthodontic treatment: Part II—The active configuration. Am. J. Orthod. Dentofacial Orthop. 118, 675 684. Kusy, R.P., Whitley, J.Q., Ambrose, W.W., Newman, J.G., 1998. Evaluation of titanium brackets for orthodontic treatment: Part I. The passive configuration. Am. J. Orthod. Dentofacial Orthop. 114, 558 572. Kwon, J.S., Lee, Y.K., Lim, B.S., Lim, Y.K., 2008. Force delivery properties of thermoplastic orthodontic materials. Am. J. Orthod. Dentofacial Orthop. 133, 228 234. quiz 328 e1. Landel, R.F., Nielsen, L.E., 1993. Mechanical Properties of Polymers and Composites, second ed. Taylor & Francis. Liljelind, I., Norberg, C., Egelrud, L., Westberg, H., Eriksson, K., Nylander-French, L.A., 2010. Dermal and inhalation exposure to methylene bisphenyl isocyanate (MDI) in iron foundry workers. Ann. Occup. Hyg. 54, 31 40. Lombardo, L., Martines, E., Mazzanti, V., Arreghini, A., Mollica, F., Siciliani, G., 2017. Stress relaxation properties of four orthodontic aligner materials: a 24-hour in vitro study. Angle Orthod. 87, 11 18. Lu, Q.-W., Macosko, C.W., 2004. Comparing the compatibility of various functionalized polypropylenes with thermoplastic polyurethane (TPU). Polymer 45, 1981 1991. Lutz, F., Phillips, R.W., Roulet, J.F., Setcos, J.C., 1984. In vivo and in vitro wear of potential posterior composites. J. Dent. Res. 63, 914 920. Matasa, C.G., 1992. Direct bonding metallic brackets: where are they heading? Am. J. Orthod. Dentofacial Orthop. 102, 552 560. Matasa, C.G., 1998. Metal strength of direct bonding brackets. Am. J. Orthod. Dentofacial Orthop. 113, 282 286. Matsui, S., Umezaki, E., Komazawa, D., Otsuka, Y., Suda, N., 2015. Evaluation of mechanical properties of esthetic brackets. J. Dent. Biomech. 6, 1758736015574401. Mcnamara, J.A., Kramer, K.L., Juenker, J.P., 1985. Invisible retainers. J. Clin. Orthod. 19, 570 578. Michelberger, D.J., Eadie, R.L., Faulkner, M.G., Glover, K.E., Prasad, N.G., Major, P.W., 2000. The friction and wear patterns of orthodontic brackets and archwires in the dry state. Am. J. Orthod. Dentofacial Orthop. 118, 662 674. Miethke, R.R., Brauner, K., 2007. A comparison of the periodontal health of patients during treatment with the Invisalign system and with fixed lingual appliances. J. Orofac. Orthop. 68, 223 231. Miller, K.B., Mcgorray, S.P., Womack, R., Quintero, J.C., Perelmuter, M., Gibson, J., et al., 2007. A comparison of treatment impacts between Invisalign aligner and fixed appliance therapy during the first week of treatment. Am. J. Orthod. Dentofacial Orthop. 131 (302), e1 e9. Miura, F., Mogi, M., Ohura, Y., Hamanaka, H., 1986. The super-elastic property of the Japanese NiTi alloy wire for use in orthodontics. Am. J. Orthod. Dentofacial Orthop. 90, 1 10.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 577 Moller, M., Klocke, A., Sadat-Khonsari, R., Schlegel, V., Kahl-Nieke, B., 2009. Torque stability of plastic brackets following multiple loading and artificial material aging—an in-vitro comparison. J. Orofac. Orthop. 70, 385 395. Moore, M.M., Harrington, E., Rock, W.P., 2004. Factors affecting friction in the pre-adjusted appliance. Eur. J. Orthod. 26, 579 583. Morton, J., Derakhshan, M., Kaza, S., Li, C.H., 2017. Design of the Invisalign system performance. Semin. Orthod. 23, 3 11. Nahoum, H.I., 1964. The vacuum formed dental contour appliance. N.Y. State Dent. J. 9, 385 390. Nahoum, H.I., 2014. Forces and moments generated by removable thermoplastic aligners. Am. J. Orthod. Dentofacial Orthop. 146, 545 546. Nair, S.V., Padmanabhan, R., Janardhanam, P., 2012. Evaluation of the effect of bracket and archwire composition on frictional forces in the buccal segments. Indian J. Dent. Res. 23, 203 208. Nikolai, R.J., 1997. Orthodontic wire: a continuing evolution. Semin. Orthod. 3, 157 165. Oh, K.-T., Kim, K.-N., Lee, M., Park, Y.-S., 2002. Corrosion wear of high molybdenum and nitrogen stainless steel for biomedical applications. J. Electrochem. Soc. 149, B146 B153. Oh, K.T., Choo, S.U., Kim, K.M., Kim, K.N., 2005. A stainless steel bracket for orthodontic application. Eur. J. Orthod. 27, 237 244. Olsen, M.E., Bishara, S.E., Boyer, D.B., Jakobsen, J.R., 1996. Effect of varying etching times on the bond strength of ceramic brackets. Am. J. Orthod. Dentofacial Orthop. 109, 403 409. Pasha, A., Vishwakarma, S., Narayan, A., Vinay, K., Shetty, S.V., Roy, P.P., 2015. Comparison of frictional forces generated by a new ceramic bracket with the conventional brackets using unconventional and conventional ligation system and the selfligating brackets: an in vitro study. J. Int. Oral Health 7, 108 113. Pithon, M.M., Oliveira, M.V., Ruellas, A.C., Bolognese, A.M., Romano, F.L., 2007. Shear bond strength of orthodontic brackets to enamel under different surface treatment conditions. J. Appl. Oral Sci. 15, 127 130. Platt, J.A., Guzman, A., Zuccari, A., Thornburg, D.W., Rhodes, B.F., Oshida, Y., et al., 1997. Corrosion behavior of 2205 duplex stainless steel. Am. J. Orthod. Dentofacial Orthop. 112, 69 79. Ponitz, R.J., 1971. Invisible retainers. Am. J. Orthod. 59, 266 272. Premaraj, T., Simet, S., Beatty, M., Premaraj, S., 2014a. Oral epithelial cell reaction after exposure to Invisalign plastic material. Am. J. Orthod. Dentofacial Orthop. 145, 64 71. Premaraj, T.S., Simet, S., Beatty, M., Premaraj, S., 2014b. Authors’ response. Am. J. Orthod. Dentofacial Orthop. 145, 551. Ratner, B.D., Hoffman, A.S., Schoen, F.J., Lemons, J.E., 2012. Biomaterials Science: An Introduction to Materials in Medicine. Elsevier Science. Reclaru, L., Meyer, J.M., 1998. Effects of fluorides on titanium and other dental alloys in dentistry. Biomaterials 19, 85 92. Reddy, Y.G., Sharma, R., Singh, A., Agrawal, V., Agrawal, V., Chaturvedi, S., 2013. The shear bond strengths of metal and ceramic brackets: an in-vitro comparative study. J. Clin. Diagn. Res. 7, 1495 1497. Remensnyder, O., 1926. A gum-massaging appliance in the treatment of pyorrhea. Dent. Cosmos 28, 381 384. Ren, Y., Maltha, J.C., Kuijpers-Jagtman, A.M., 2003. Optimum force magnitude for orthodontic tooth movement: a systematic literature review. Angle Orthod. 73, 86 92.
578 Advanced Dental Biomaterials Ren, Y., Maltha, J.C., Van’t Hof, M.A., Kuijpers-Jagtman, A.M., 2004. Optimum force magnitude for orthodontic tooth movement: a mathematic model. Am. J. Orthod. Dentofacial Orthop. 125, 71 77. Rodriguez-Mercado, J.J., Roldan-Reyes, E., Altamirano-Lozano, M., 2003. Genotoxic effects of vanadium(IV) in human peripheral blood cells. Toxicol. Lett. 144, 359 369. Rossini, G., Parrini, S., Castroflorio, T., Deregibus, A., Debernardi, C.L., 2015. Efficacy of clear aligners in controlling orthodontic tooth movement: a systematic review. Angle Orthod. 85, 881 889. Rosvall, M.D., Fields, H.W., Ziuchkovski, J., Rosenstiel, S.F., Johnston, W.M., 2009. Attractiveness, acceptability, and value of orthodontic appliances. Am. J. Orthod. Dentofacial Orthop. 135, 276 277. 276 e1-12; discussion. Russell, J.S., 2005. Current products and practice. J. Orthod. 32, 146 163. Ryokawa, H., Miyazaki, Y., Fujishima, A., Miyazaki, T., Maki, K., 2006. The mechanical properties of dental thermoplastic materials in a simulated intraoral environment. Orthod. Waves 65, 64 72. Sadat-Khonsari, R., Moshtaghy, A., Schlegel, V., Kahl-Nieke, B., Moller, M., Bauss, O., 2004. Torque deformation characteristics of plastic brackets: a comparative study. J. Orofac. Orthop. 65, 26 33. Saunders, C.R., Kusy, R.P., 1994. Surface topography and frictional characteristics of ceramic brackets. Am. J. Orthod. Dentofacial Orthop. 106, 76 87. Schiff, N., Dalard, F., Lissac, M., Morgon, L., Grosgogeat, B., 2005. Corrosion resistance of three orthodontic brackets: a comparative study of three fluoride mouthwashes. Eur. J. Orthod. 27, 541 549. Schott, T.C., Göz, G., 2011. Color fading of the blue compliance indicator encapsulated in removable clear Invisalign Teens aligners. Angle Orthod. 81, 185 191. Schupp, W., Haubrich, J., 2016. Aligner Orthodontics: Diagnostics, Biomechanics, Planning and Treatment. Quintessence Publishing. Schuster, S., Eliades, G., Zinelis, S., Eliades, T., Bradley, T.G., 2004. Structural conformation and leaching from in vitro aged and retrieved Invisalign appliances. Am. J. Orthod. Dentofacial Orthop. 126, 725 728. Sergl, H.G., Zentner, A., 1998. A comparative assessment of acceptance of different types of functional appliances. Eur. J. Orthod. 20, 517 524. Sheridan, J.J., Ledoux, W., Mcminn, R., 1993. Essix retainers: fabrication and supervision for permanent retention. J. Clin. Orthod. 27, 37 45. Simon, M., Keilig, L., Schwarze, J., Jung, B.A., Bourauel, C., 2014. Forces and moments generated by removable thermoplastic aligners: incisor torque, premolar derotation, and molar distalization. Am. J. Orthod. Dentofacial Orthop. 145, 728 736. Springate, S.D., Winchester, L.J., 1991. An evaluation of zirconium oxide brackets: a preliminary laboratory and clinical report. Br. J. Orthod. 18, 203 209. Swartz, M.L., 1988. Ceramic brackets. J. Clin. Orthod. 22, 82 88. Toumelin-Chemla, F., Rouelle, F., Burdairon, G., 1996. Corrosive properties of fluoridecontaining odontologic gels against titanium. J. Dent. 24, 109 115. Tuncay, O.C., 2006. In: Nguyen Cv, C.J. (Ed.), The Invisalign System. Quintessence Publ, London [u.a.]. Vardimon, A.D., Robbins, D., Brosh, T., 2010. In-vivo von Mises strains during Invisalign treatment. Am. J. Orthod. Dentofacial Orthop. 138, 399 409. Voudouris, J.C., Schismenos, C., Lackovic, K., Kuftinec, M.M., 2010. Self-ligation esthetic brackets with low frictional resistance. Angle Orthod. 80, 188 194.
Biomaterials used in orthodontics: brackets, archwires, and clear aligners 579 Wassell, R.W., Mccabe, J.F., Walls, A.W., 1994. A two-body frictional wear test. J. Dent. Res. 73, 1546 1553. Wheeler, T., Patel, N., Mcgorray, S., 2017. On orthodontic tooth movement. J. Aligner Orthod. 1, 21 27. White, D.W., Julien, K.C., Jacob, H., Campbell, P.M., Buschang, P.H., 2017. Discomfort associated with Invisalign and traditional brackets: a randomized, prospective trial. Angle Orthod. 87, 801 808. Williams, C.L., Khalaf, K., 2013. Frictional resistance of three types of ceramic brackets. J. Oral Maxillofac. Res. 4, e3. Wisnewski, A.V., Srivastava, R., Herick, C., Xu, L., Lemus, R., Cain, H., et al., 2000. Identification of human lung and skin proteins conjugated with hexamethylene diisocyanate in vitro and in vivo. Am. J. Respir. Crit. Care Med. 162, 2330 2336. Wong, B.H., 2002. Invisalign A to Z. Am. J. Orthod. Dentofacial Orthop. 121, 540 541. Wu, D., 2014. Oral epithelial cell reaction and Invisalign treatment. Am. J. Orthod. Dentofacial Orthop. 145, 551. Zhang, N., Bai, Y., Ding, X., Zhang, Y., 2011. Preparation and characterization of thermoplastic materials for invisible orthodontics. Dent. Mater. J. 30, 954 959. Zinelis, S., Annousaki, O., Eliades, T., Makou, M., 2003. Metallographic structure and hardness of titanium orthodontic brackets. J. Orthop./Fortschritte der Kieferorthopädie 64, 426 433. Zinelis, S., Eliades, T., Eliades, G., Makou, M., Silikas, N., 2005. Comparative assessment of the roughness, hardness, and wear resistance of aesthetic bracket materials. Dent. Mater. 21, 890 894. Zhylich, D., Krishnan, P., Muthusami, P., Rayner, T., Shroff, M., Doria, A., et al., 2017. Effects of orthodontic appliances on the diagnostic quality of magnetic resonance images of the head. Am J Orthod Dentofacial Orthop 151, 484 499.
21 Dental implants materials and surface treatments Shariq Najeeb1, Maria Mali2, Syed Azeem Ul Yaqin3, Muhammad Sohail Zafar4, Zohaib Khurshid3, Abdullah Alwadaani3 and Jukka P. Matinlinna5 1 National Center for Proteomics, Karachi University, Pakistan, 2Department of Orthodontics, Islamic International Dental College & Hospital, Riphah International University, Islamabad, Pakistan, 3Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Al-Ahsa, Saudi Arabia, 4Department of Restorative Dentistry, College of Dentistry, Taibah University, Almadinah Almunawwarah, Saudi Arabia, 5Dental Materials Science, Applied Oral Sciences, Faculty of Dentistry, The University of Hong Kong, Prince Philip Dental Hospital, Sai Ying Pun, Hong Kong SAR, P.R. China Chapter Outline 21.1 21.2 21.3 21.4 21.5 21.6 Introduction 581 Osseointegration: cellular and biomaterial aspects 583 Biomaterial properties and implant surface characteristics Biomechanical properties of dental implants 584 Surface properties 585 Type of dental implant material 586 21.6.1 Alveolar bone properties 587 21.6.2 Influence of oral health and systemic disease on implant survival 21.7 Modification of the dental implants 584 587 588 21.7.1 Modification of titanium implants 588 21.8 Functionally graded/hierarchical dental implant surfaces 590 21.9 Modification of the polyetheretherketone dental implants 590 21.10 Modification of zirconia implants 592 21.11 Conclusion 592 References 592 21.1 Introduction In general, a dental implant is a synthetic medical device that is surgically placed directly into the alveolar bone and supports a prosthodontic or an orthodontic appliance (Adell, 1981). Ideally, there needs to be a direct physical, chemical, and biological interface between the human tissues and the implant material (Le Guéhennec et al., 2007). There are two main components of a dental implant: the abutment and Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00021-9 Copyright © 2019 Elsevier Ltd. All rights reserved.
582 Advanced Dental Biomaterials Figure 21.1 A titanium screw in the dog mandible. The implant is osseointegrated and the soft tissue integration is present. The abutment connects the implants with the crown, which is made of metal in this experimental setting. Source: Adapted from Gruber, R., Bosshardt, D.D., 2014. Dental implantology and implants —tissue interface. In: Stem Cell Biology and Tissue Engineering in Dental Sciences. Academic Press, pp. 735 747. https://doi.org/10.1016/B978-0-12-397157-9.00078-3, with publisher permission (Gruber and Bosshardt, 2014). the screw (fixture). The abutment is the portion that is visible above the gingiva level and retains or supports a prosthodontic or an orthodontic appliance, while the screw or the root of the implant is submerged in the alveolar bone. As a dental implant is a foreign object, it is imperative for the human hard and soft tissues to “accept” it and form an intimate physical, stable, and functional interface with it (Annunziata and Guida, 2015). The initial goal of implant therapy is to achieve osseointegration, which is the direct stabilized physical and structural interface between the implant and the surrounding bone. A diagram of an osseointegrated implant is shown in Fig. 21.1. Mankind has tried to use various materials as dental implants. Archeological evidence indicates the use of ivory, human teeth, bamboo, and some metal alloys as implants in ancient civilizations. However, it was not until the 1950s that titanium
Dental implants materials and surface treatments 583 was attempted as an implant material when titanium chambers were placed into the soft tissues of rabbit ears at the University of Cambridge (Brånemark and Lindström, 1963). While studying the process of bone healing at Lund University, Per-Ingvar Brånemark noticed that he was unable to remove the titanium chambers from rabbit femurs. Following research on animal and human subjects, Brånemark developed further the contemporary titanium dental implant. In 1965 he placed the implants in the jaws of human volunteers. He termed this stable physical and biofunctional interface between bone and implant as osseointegration (Brånemark and Lindström, 1963). There are two possible scenarios when an implant is placed into the alveolar bone. The first possibility is the formation of a connective tissue between the implant and the bone (Smeets et al., 2016). This scenario is unfavorable as it may cause the implant to be loosened and dislodged from the bone, leading to implant failure. The most favorable outcome of dental implant therapy is osseointegration (also called osteointegration) which is the formation of a direct, stable interface between the implant and surrounding bone. Although the most common material used in implant dentistry has been traditionally titanium and its alloys, more recently several other materials have been used as implants (Lang and Matinlinna, 2014). The aim of this chapter is to provide the reader with a comprehensive background on the concept of osseointegration along with implant biomaterials, prosthodontic and clinical aspects of oral implantology. Moreover, factors governing the failure and success rates of implants will be discussed. 21.2 Osseointegration: cellular and biomaterial aspects Following placement in the alveolar bone, the first tissue that comes into contact with the implant material is blood (Telleman et al., 2010). Just within around 1 minute of implantation, plasma proteins are adsorbed onto the implant surface. Platelets are attracted to, and interact with, the adsorbed proteins. Platelets play a vital role in healing, blood clotting, and thrombus formation. Eventually, other cells migrate to the implant site and interact with the implant surface proteins through membrane receptors. Plasma delivers nutrients (glucose, proteins, amino acids, and cholesterol) and other substances to the surgical site. The interactions of these substances and cells alter the surface properties of the implant. Several cells interact with the dental implant fixture surface following blood clotting. However, mesenchymal stem cells (MSCs) play a determining factor in healing of tissues around the implant. Growth factors released following surgery and during inflammation have been indicated to attract MSCs to the implant surface and adhere to it. In the right environment, MSCs have the potential to differentiate into many different types of cells, including bone-forming cells (osteoblasts), cartilage-forming cells (chondroblasts), and fibroblastic cells. Ideally, MSCs should differentiate into osteoblasts around the portion of the dental implant submerged in
584 Advanced Dental Biomaterials the alveolar bone, leading to osteointegration. However, if a fibrous capsule forms instead, implant failure occurs. There are a number of factors that affect osseointegration. Local factors include biomaterial properties and implant surface characteristics, the quality of the alveolar bone, periodontal disease, and oral hygiene. Systemic factors include general health, immunity smoking status, genetics, and metabolic disease (Marei and El Backly, 2018; Viña et al., 2014; Zafar et al., 2015). 21.3 Biomaterial properties and implant surface characteristics As dental implant therapy primarily involves placement of a foreign object (i.e., the artificial implant) in direct contact with a living tissue, the type of the implant material used, and its surface properties play a pivotal role in the success—or failure—of a dental implant. The abutment (root connection) portion of a dental implant may be divided into two areas: the core and the surface. The core is the main bulk of the dental material used in the construction of the implant (Lang and Matinlinna, 2014; Aboushelib and Matinlinna, 2014), while the surface is the portion of the implant in direct contact with living tissue (Guillaume, 2016). 21.4 Biomechanical properties of dental implants The bone and its related soft tissues are dynamic. They remodel according to the physiomechanical demands throughout life. Most dental implant materials used today are significantly less elastic when compared to human bone. According to Wolff’s Law, the amount of bone remodeling and formation is dependent upon the mechanical load applied to it (Frost, 1994). As evident in Table 21.1, titanium has a significantly higher elastic modulus and tensile strength compared to bone. If a Table 21.1 Comparison of the different hard tissue with dental implants materials with relation to tensile strength and modulus of elasticity. Material Tensile strength Modulus of elasticity References Enamel Dentine Human cortical bone Titanium PEEK CFR-PEEK 47.5 104 104 121 954 976 80 120 40 83 15 14 102 110 3 4 18 Rees and Jacobsen (1993) Rees and Jacobsen (1993) Rho et al. (1993) Niinomi (1998) Sandler et al. (2002) Sandler et al. (2002) CFR-PEEK, carbon-reinforced polyetheretherketone; PEEK, polyetheretherketone.
Dental implants materials and surface treatments 585 relatively inelastic material, such as a titanium implant, is placed in the bone, it may prevent the load from being transferred to the surrounding bone. This may lead to bone resorption. This process is known as stress-shielding (Asgharzadeh Shirazi et al., 2017). That said, one of the ideal properties of a dental implant material would be to have physical properties identical to that of alveolar bone. However, it is very difficult to replicate nature in a cost-effective fashion. Hence, research has been conducted to produce various Ti alloys (Lang and Matinlinna, 2014) and dental implant biocomposites that have similar mechanical behavior to bone (Asgharzadeh Shirazi et al., 2017). Commercially pure titanium is the most commonly used dental implant material. Titanium is graded 1 4 according to the purity, that is, oxygen and iron contents. Commercially pure type 4 titanium (grade 4 cpTi) is used to fabricate dental implants. For example, Ti6Al4V is a grade 5 titanium alloy that has a higher fatigue resistance and strength than other grades. Hence, Ti6Al4V is also used as dental implant material. More recently, materials such as reinforced polymeric composites (e.g., glass fiber reinforced composites), ceramics, and silicon nitrate have been studied for potential dental implant applications (Osman and Swain, 2015; Zhang and Matinlinna, 2012). At the implant surface, two factors affect the initial bone formation around the dental implant: surface roughness and hydrophilicity (Eliasa et al., 2012). Other factors, such as the presence of biomimetic molecules and factors in the implant material, may also impact osteointegration (Khurshid et al., 2018; Najeeb et al., 2017b). Surface roughness dramatically increases the surface area of the dental implant (1) to achieve higher initial implant stability by mechanical interlocking and (2) to maximize cell adhesion at the surface to promote bone ingrowth into the surface porosities. Surface roughness can be introduced at three levels: macro, micro, and nano (Khurshid et al., 2015; Najeeb et al., 2016b). Macrolevel surface roughness is the introduction of certain geometrical features to the implant design which are visible to the naked eye or measuring more than 10 µm. A tapered, root-like design and threading are examples of macrolevel surface modifications that improve initial implant fixation (Cook et al., 1982). However, solely opting out to use macrolevel surface treatments increases the chances of ion-leakage. This said, in order to minimize ion-leakage but maximize bone ingrowth, microlevel (1 10 µm) and nanolevel (,1 µm) implant modifications are more appropriate. They may also be concurrently used with certain macrolevel surface treatments (Ryu et al., 2014). 21.5 Surface properties In general, human tissues comprise up to 60% water. Hence, it is widely understood that it is important for the implant surface to be highly hydrophilic (Ferraris et al., 2015). The hydrophilicity of a material is the measure of its affinity for water. Ideally, the surface of a perfectly hydrophilic material should form a 0 degree angle with a water droplet. However, practically speaking, a material such as a dental
586 Advanced Dental Biomaterials Table 21.2 Some examples of titanium implant surface modifications and their respective properties. Implant surface Surface roughness (Ra) References Pure titanium Titanium-sprayed plasma Plasma-sprayed apatite Biomimetic calcium 0.22 6 0.01 7.01 6 2.09 1.06 6 0.21 1.83 6 0.64 Mabboux et al. (2004) Bagno and Di Bello (2004) Giavaresi et al. (2003) Le Guéhennec et al. (2007) implant is modified to form an angle which is less than 90 degrees with a water droplet (Ferraris et al., 2015; Hong et al., 2013). The more a liquid droplet spreads out over the surface of the implant material, the lower the surface contact angle. Over the last few decades, a significant amount of research has been conducted to improve the hydrophilicity of dental implants (Ferraris et al., 2015; Held et al., 2013; Hong et al., 2013). It has been observed that highly hydrophilic dental implant surfaces not only exhibit closer bone implant contact (Rupp et al., 2014) but they may also activate macrophages to produce antiinflammatory factors (Hotchkiss et al., 2017). Several methods that increase the hydrophilicity of dental implants will be described in this chapter. In addition to having a hydrophilic surface, a dental implant fixture should also possess osseoconductive and/or osseoinductive properties (Le Guéhennec et al., 2007). A material is said to be osseoinductive when it promotes the undifferentiated mesenchymal cells in a tissue to mature into bone-forming osteoblasts, whereas osseoconductive materials promote bone formation on their surface. A significant amount of research has been conducted to improve the osseoconductive/osseoinductive properties of implant materials to enhance the bone implant interface. Methods include production of nanoporous implant surfaces (Carradò et al., 2017), laser treatment only (Han et al., 2017), spraying of osseoconductive materials such as hydroxyapatite and fluoroapatite (Mohseni et al., 2014), using a laser-assisted biomimetic coating of calcium phosphate (Nathanael et al., 2018), and functionalized biomimetic coatings with adhesion peptides for dental implants (Roessler et al., 2001). Moreover, even eicosapentaenoic acid has been attempted in vivo (Mustafa et al., 2016) and some silane coatings in vitro (Villard et al., 2015) as potential, beneficial dental implant coatings. Ultraviolet light treatment and some other disinfection methods may also be worth further study and consideration (Han et al., 2017) (Table 21.2). 21.6 Type of dental implant material As discussed above, grade 4 cpTi and its alloy Ti6Al4V are the most commonly used dental implant material by manufacturers. Apart from titanium, there are several other materials which have been used clinically as dental implants. Carbon fiber reinforced carbon had been studied as a potential dental implant material in
Dental implants materials and surface treatments 587 the 1970 80s (Adams et al., 1978). In addition, E-glass fibers have been attempted and introduced (Zhang and Matinlinna, 2012). However, due to the possible release of fiber debris from the implant into the periimplant tissues, the US Food and Drug Administration consequently did not approve the use of carbon fiber implants in the human body (Petersen, 2016). Nevertheless, ceramics such as zirconia offer a more esthetic alternative to titanium (Cionca et al., 2017). More recently, polymeric reinforced composites, such as polyetheretherketone (PEEK), have been attempted and tested in animals in order to assess their osseointegration with alveolar bone (Najeeb et al., 2016d). In addition, silicon nitride, a ceramic used in spinal reconstruction and maxillofacial rehabilitation, has also been tested as a dental implant material (Webster et al., 2012). Each material does have a different way of interaction with the human tissues. Hence, choosing the right biocompatible material for implant applications is a major factor which plays an absolutely important role in the clinical success of implants. 21.6.1 Alveolar bone properties As proposed by Lekholm and Zarb, human bone can be classified into four types, depending on its density (Al-Ekrish et al., 2018), as follows: 1. 2. 3. 4. Type I—Primarily compact bone Type II—A core of dense spongy (trabecular) bone surrounded by compact bone Type III—A core of dense spongy bone surrounded by a thin layer of compact bone Type IV—A core of low-density spongy bone surrounded by a thin layer compact bone Most studies suggest that implants exhibit the optimal survival rate when placed in Type I or Type II bone. On the other hand, although a 5-year study on 1045 dental implants found no difference in the survival rate of Types I, II, and III implants, a failure rate of 35% was observed in implants placed in Type IV bone (Jaffin and Berman, 1991). Nonetheless, a recent systematic review of the latest studies has suggested that implant surface modifications may improve the survival rate of implants placed in osteoporotic animal bones (Ghanem et al., 2017). However, more clinical studies are necessary before conclusive evidence can be ascertained (Dereka et al., 2018). Conventional radiography and computerized tomography can both be used to assess the bone density before implant placement (Norton and Gamble, 2001). In addition, a meticulous periodontal screening and recording protocol should be employed to document the clinical signs and symptoms of periodontal disease before initiation of implant therapy. The reader is encouraged to refer to textbooks focusing on clinical periodontology and implantology to learn more about periodontal and implant disease and the surgical steps involved in implant therapy. 21.6.2 Influence of oral health and systemic disease on implant survival Several systemic diseases have been implicated in dental implant failures. Smoking, age, uncontrolled diabetes, and a history of radiation therapy in the head and neck
588 Advanced Dental Biomaterials region have been implicated in early implant failures (Dawson and Jasper, 2015). A recently published systematic review by the authors of this chapter has indicated a negative impact of Down syndrome on successful dental implant therapy (Najeeb et al., 2017a). 21.7 Modification of the dental implants Before dental implants are placed in the alveolar bone, their fixture surfaces are modified. They may be modified by several mechanisms such as plasma spraying, ion dispersion, coatings of bioactive materials (Khurshid et al., 2019; Najeeb et al., 2016c; Zafar et al., 2019; Aivazi et al., 2016; Han et al., 2018), and antimicrobial proteins/peptides (Khurshid et al., 2018, 2017a, 2017b, 2016). The mechanism of modification primarily depends on the type of the core implant material. As stated above, implants can be modified at the following levels: macro, micro, and nano. Most contemporary dental implants are endosseous, screw-type devices (Le Guéhennec et al., 2007) which are further modified at the surface and/or by introduction of bioactive substances within the core. 21.7.1 Modification of titanium implants 21.7.1.1 Titanium plasma spraying Titanium plasma spraying (TPS) involves injecting titanium powder through a hot plasma torch, as shown in Fig. 21.2. The molten titanium particles fuse and condense at the surface of the implant to produce a rough surface, enhancing the surface area for improved bone deposition and contact. Studies have indicated that the Figure 21.2 A schematic diagram of the plasma-spraying process. A molten jet of a surfactant is coated onto the implant surface. The molten material solidifies and adheres to the implant surface.
Dental implants materials and surface treatments 589 thickness of the plasma coating is in the range of 40 50 µm and has an average roughness of approximately 7 µm (Buser et al., 1991). However, TPS has a major drawback: residual particles. Not only have titanium particles been found in bone adjacent to TPS implant surfaces (Urban et al., 2000), organs such as liver, spleen, and abdominal lymph nodes have all been observed to contain the metallic particles in patients who have received hip and knee replacements (Urban et al., 2004). Although the long-term effects of metal debris found in organs are unknown, titanium debris may have local and systemic inflammatory or carcinogenic effects (Le Guéhennec et al., 2007). Hence, more recently, studies have been conducted to produce more stable and bioactive titanium surfaces, which are described in the following subsections. 21.7.1.2 Grit-blasting Grit-blasting (or sometimes called sand-blasting) involves production of a rough implant surface by means of a jet of a particles of an abrasive material under high pressure (Lung and Matinlinna, 2012). Two commonly used abrasives used to gritblast implants are alumina (Al2O3) and titanium dioxide (TiO2). Although gritblasted dental implant surfaces have been observed to exhibit enhanced boneimplant contact in animal subjects, they may, however, exhibit drawbacks. First, it has been generally believed that residual ceramic particles can detach from the implant surface and cause inflammation in the periimplant bone (Esposito et al., 1998; Le Guéhennec et al., 2007). However, an animal study by Piattelli et al. (2003) has found no significant negative effects of residual alumina particles. Moreover, alumina particles may undermine the otherwise excellent corrosion resistance of titanium (Aparicio et al., 2003). Other materials that have been used to grit-blast titanium implants are titanium oxide (Hotchkiss et al., 2017; Ivanoff et al., n.d.) and resorbable materials such as calcium phosphates (Xuereb et al., 2015). More recently, grit-blasting with alumina has also been combined with treatment with NaOH, followed by heat treatment at 600 C to produce rough implant surfaces (Herrero-Climent et al., 2018). 21.7.1.3 Nanostructured titanium implant surfaces To maximize the surface area for cellular adhesion and protein adsorption, modification of titanium surfaces at the nanometer-scale has been suggested. Laser and lithography can be employed to produce nanometer-sized roughness on titanium implants (Anselme et al., 2002; Zhu et al., 2004). Modern 3D-printing technology may also give some answers. Indeed, in vitro studies on 3D-printed implant surfaces possessing surface roughness in the range of 32 6 4 nm on Ti microparticles and 40 6 4 nm on underlying flat Ti have exhibited the potential to promote osteoblast proliferation (Gulati et al., 2017). Nonetheless, although the production of nano-sized implant modifications has the potential to enhance osseointegration, due to the lack of animal and clinical studies, nanostructured implant surfaces are not yet widely used in routine implant dentistry.
590 Advanced Dental Biomaterials 21.7.1.4 Acid-etching of titanium surfaces Another possible way to roughen titanium implants is acid-etching, which is a process like enamel and dentine etching employed prior to bonding to resin composites. However, titanium implants, being alloys, require much stronger acids to be etched. Hydrofluoric acid (HF), hydrochloric acid (HCl), nitric acid (HNO3), and sulfuric acid (H2SO4) have been used to produce rough implant surfaces in the micrometer range (Klokkevold et al., 2001; Zinger et al., 2004). Acid-etched surfaces have been observed to exhibit higher osseointegration when compared to TPS implant surfaces, possibly due to a greater surface area and a higher wettability (Cho and Park, 2003). Moreover, etching by HF incorporates fluoride ions onto the implant, and these make the surface more osseoconductive (Cooper et al., 2006). Recently, fluoride-modified dental implants have been introduced and marketed as OsseoSpeed (Dentsply Sirona). 21.7.1.5 Calcium phosphate coated titanium surfaces Calcium hydroxyapatite (CHAp) is a form of calcium phosphate primarily found in the mineralized component of animal and human hard tissues. In regenerative medicine, CHAp has been used to promote the regeneration and healing of calcified tissue such as bone and teeth (Zhou and Lee, 2011). Similarly, when coated on dental implants, CHAp has been found to enhance bone implant contact and cellular proliferation (Xuereb et al., 2015). 21.8 Functionally graded/hierarchical dental implant surfaces Some recent research in regenerative periodontology has proposed producing “functionally graded” or “hierarchical” biomaterial surfaces (Qasim et al., 2017). Rather than consisting of just a core with a coated or modified surface, functionally graded materials (FGMs) have multiple layers, with each layer possessing different physical, chemical, and biological properties (Hedia and Fouda, 2013). Functionally graded CHAp coatings on titanium implants have shown promising results in vivo (Watari et al., 1997). Moreover, computer simulation studies have suggested that using FGMs may reduce the adverse effects of dental implants on alveolar bone (Lin et al., 2009). Nevertheless, FGMs have yet to see use in routine clinical practice, owing to the need for more clinical and preclinical research. 21.9 Modification of the polyetheretherketone dental implants The ability to modify polymeric composites has enabled them to be used in multiple fields in dentistry. For instance, resin composites have been used as
Dental implants materials and surface treatments 591 tooth-colored materials for decades, poly(methylmethacrylate)-based denture materials have proven to be user-friendly and durable prosthodontic materials. Moreover, polymeric impression materials produce excellent surface detail and elastic properties for producing accurate impressions of soft and hard oral tissues. More recently, PEEK, a polymer produced by the step-growth polymerization of bis-phenolate salts by dialkylation, has been suggested for use as a dental implant material (Najeeb et al., 2015). In addition to being tooth-colored, PEEK may also exert less adverse effects on the periodontal structures due to its physical properties being like that of human bone (as shown in Table 21.1). Studies have suggested that PEEK implants exhibit lesser stress-shielding when compared to titanium implants. However, in its unmodified form, the osseointegration and bioactivity exhibited by PEEK is inferior to those demonstrated by titanium implants (Najeeb et al., 2016a). To address this issue, a number of methods have been proposed to modify the properties of PEEK implants in order to enhance their osseointegration. As shown in Fig. 21.3, PEEK may be combined with various bioactive materials via the process of melt-blending. PEEK may be coated by CHAp through plasma spraying, similarly to titanium implants (Fauchais and Vardelle, 2012). However, plasma spraying at high temperatures on a polymer like PEEK may have deleterious effects on the physical properties of not only the core polymer but also the PEEK coating interface (Molitor et al., 2001). Nonthermal processes such as electron-beam (e-beam) deposition, spin-coating, and plasma gas etching treatment may circumvent that problem (Balmer et al., 2018; Dawson and Jasper, 2015; Han et al., 2017; Lung and Matinlinna, 2012; Najeeb et al., 2017a). E-beam coating involves the formation of a thin, nanorough coating of material on a substrate via electron-induced deposition. Spin-coating is carried out by slowly dropping a solution of apatites in organic solvents onto the implant spinning at high speeds. In plasma gas etching, the implant surface is exposed to plasma gases at high pressure to produce nanorough surfaces. That said, probably the most unique aspect of PEEK implants would be their ability to be combined with bioactive materials such as apatite to produce potentially osteoconductive fiber-reinforced dental implants (Yabutsuka et al., 2018). So far, no large-scale clinical trials have demonstrated Figure 21.3 A schematic diagram of the melt-blending process by which bioactive PEEK composite implants are produced. During the process, the bioactive particles are codispersed with PEEK in molten form in a suitable solvent. When placed in a mold, the bioactive composite can cool down and produce a solid which is then shaped into an implant. PEEK, Polyetheretherketone.
592 Advanced Dental Biomaterials long-term clinical viability of PEEK biocomposites. Nevertheless, positive results from recent animal studies may bring these implants into wider use in the dental clinics. 21.10 Modification of zirconia implants The surface treatments employed for zirconia implants (Aboushelib and Matinlinna, 2014) are like those carried out on titanium implants. Acid-etching, lasermodification, and grit-blasting have all been used to enhance the bone implant interface of zirconia implants (Oliva et al., n.d.). However, clinical data regarding the 5-year survival rate of modified zirconia implants is limited. A systematic review (in 2008) of seven animal studies by Wenz et al. (n.d.) failed to state any recommendation regarding the use of zirconia implants in the clinical setting. Although some reviews of clinical studies show a promising outcome of zirconia implants after 60 months, more comparative studies are required to ascertain their performance when compared to titanium (Balmer et al., 2018; Montero et al., 2015; Wenz et al., n.d.). 21.11 Conclusion The surface of dental implants is a vital factor. It may be improved in various ways but the systemic health of the implant-recipient, surgical procedures, and the oral environmental factors has major impacts on osseointegration. Nonetheless, materials such as PEEK and zirconia may present a viable esthetic alternative to titanium in the near future. References Aboushelib, M.N., Matinlinna, J.P., 2014. Zirconia in implant dentistry. In: Matinlinna, J.P. (Ed.), Handbook of Oral Biomaterials. Pan Stanford Publishing, Singapore, pp. 459 476. Available from: https://doi.org/10.1201/b15644. Adams, D., Williams, D.F., Hill, J., 1978. Carbon fiber-reinforced carbon as a potential implant material. J. Biomed. Mater. Res. 12, 35 42. Available from: https://doi.org/ 10.1002/jbm.820120104. Adell, R., 1981. A 15-year study of osseointegrated implants in the treatment of the edentulous jaw. Int. J. Oral Surg. Available from: https://doi.org/10.1016/S0300-9785(81)80077-4. Aivazi, M., Hossein Fathi, M., Nejatidanesh, F., Mortazavi, V., HashemiBeni, B., Matinlinna, J.P., et al., 2016. The evaluation of prepared microgroove pattern by femtosecond laser on alumina-zirconia nano-composite for endosseous dental implant application. Lasers Med. Sci. 31, 1837 1843. Available from: https://doi.org/10.1007/s10103-016-2059-8. Al-Ekrish, A., Widmann, G., Alfadda, S., 2018. Revised, computed tomography based Lekholm and Zarb Jawbone quality classification. Int. J. Prosthodont. 31, 342 345. Available from: https://doi.org/10.11607/ijp.5714.
Dental implants materials and surface treatments 593 Annunziata, M., Guida, L., 2015. The effect of titanium surface modifications on dental implant osseointegration. Biomaterials for Oral and Craniomaxillofacial Applications. Karger Publishers, pp. 62 77. Available from: https://doi.org/10.1159/000381694. Anselme, K., Bigerelle, M., Noël, B., Iost, A., Hardouin, P., 2002. Effect of grooved titanium substratum on human osteoblastic cell growth. J. Biomed. Mater. Res. 60, 529 540. Available from: https://doi.org/10.1002/jbm.10101. Aparicio, C., Javier Gil, F., Fonseca, C., Barbosa, M., Planell, J.A., 2003. Corrosion behaviour of commercially pure titanium shot blasted with different materials and sizes of shot particles for dental implant applications. Biomaterials 24, 263 273. Available from: https://doi.org/10.1016/S0142-9612(02)00314-9. Asgharzadeh Shirazi, H., Ayatollahi, M.R., Asnafi, A., 2017. To reduce the maximum stress and the stress shielding effect around a dental implant bone interface using radial functionally graded biomaterials. Comput. Methods Biomech. Biomed. Eng. 20, 750 759. Available from: https://doi.org/10.1080/10255842.2017.1299142. Balmer, M., Spies, B.C., Vach, K., Kohal, R.-J., Hämmerle, C.H.F., Jung, R.E., 2018. Threeyear analysis of zirconia implants used for single-tooth replacement and three-unit fixed dental prostheses: a prospective multicenter study. Clin. Oral Implants Res. 29, 290 299. Available from: https://doi.org/10.1111/clr.13115. Brånemark, P.-I., Lindström, J., 1963. A modified rabbit’s ear chamber high-power highresolution studies in regenerated and preformed tissues. Anat. Rec. 145, 533 540. Available from: https://doi.org/10.1002/ar.1091450406. Buser, D., Schenk, R.K., Steinemann, S., Fiorellini, J.P., Fox, C.H., Stich, H., 1991. Influence of surface characteristics on bone integration of titanium implants. A histomorphometric study in miniature pigs. J. Biomed. Mater. Res. 25, 889 902. Available from: https://doi.org/10.1002/jbm.820250708. Carradò, A., Perrin-Schmitt, F., Le, Q.V., Giraudel, M., Fischer, C., Koenig, G., et al., 2017. Nanoporous hydroxyapatite/sodium titanate bilayer on titanium implants for improved osteointegration. Dent. Mater. 33, 321 332. Available from: https://doi.org/10.1016/j. dental.2016.12.013. Cho, S.-A., Park, K.-T., 2003. The removal torque of titanium screw inserted in rabbit tibia treated by dual acid etching. Biomaterials 24, 3611 3617. Available from: https://doi. org/10.1016/S0142-9612(03)00218-7. Cionca, N., Hashim, D., Mombelli, A., 2017. Zirconia dental implants: where are we now, and where are we heading? Periodontology 2000 73, 241 258. Available from: https:// doi.org/10.1111/prd.12180. Cook, S.D., Klawitter, J.J., Weinstein, A.M., 1982. The influence of implant geometry on the stress distribution around dental implants. J. Biomed. Mater. Res. 16, 369 379. Available from: https://doi.org/10.1002/jbm.820160406. Cooper, L.F., Zhou, Y., Takebe, J., Guo, J., Abron, A., Holmén, A., et al., 2006. Fluoride modification effects on osteoblast behavior and bone formation at TiO2 grit-blasted c.p. titanium endosseous implants. Biomaterials 27, 926 936. Available from: https://doi. org/10.1016/J.BIOMATERIALS.2005.07.009. Dawson, D.R., Jasper, S., 2015. Key systemic and environmental risk factors for implant failure. Dent. Clin. North Am. 59, 25 39. Available from: https://doi.org/10.1016/J. CDEN.2014.09.002. Dereka, X., Calciolari, E., Donos, N., Mardas, N., 2018. Osseointegration in osteoporotic-like condition: a systematic review of preclinical studies. J. Periodontal Res. Available from: https://doi.org/10.1111/jre.12566.
594 Advanced Dental Biomaterials Eliasa, C.N., Rocha, F.A., Nascimento, A.L., Coelho, P.G., 2012. Influence of implant shape, surface morphology, surgical technique and bone quality on the primary stability of dental implants. J. Mech. Behav. Biomed. Mater. 16, 169 180. Available from: https://doi. org/10.1016/j.jmbbm.2012.10.010. Esposito, M., Hirsch, J.M., Lekholm, U., Thomsen, P., 1998. Biological factors contributing to failures of osseointegrated oral implants: (II). Etiopathogenesis. Eur. J. Oral Sci. Available from: https://doi.org/10.1046/j.0909-8836.t01-6-.x. Fauchais, P., Vardelle, A., 2012. Solution and suspension plasma spraying of nanostructure coatings. Advanced Plasma Spray Applications. InTech. Available from: https://doi.org/ 10.5772/34449. Ferraris, S., Bobbio, A., Miola, M., Spriano, S., 2015. Micro- and nano-textured, hydrophilic and bioactive titanium dental implants. Surf. Coat. Technol. 276, 374 383. Available from: https://doi.org/10.1016/J.SURFCOAT.2015.06.042. Frost, H.M., 1994. Wolff’s law and bone’s structural adaptations to mechanical usage: an overview for clinicians. Angle Orthod. Available from: https://doi.org/10.1043/00033219(1994)064 , 0175:WLABSA . 2.0.CO;2. Ghanem, A., Kellesarian, S.V., Abduljabbar, T., Al-Hamoudi, N., Vohra, F., Javed, F., 2017. Role of osteogenic coatings on implant surfaces in promoting bone-to-implant contact in experimental osteoporosis. Implant Dent. 26, 770 777. Available from: https://doi.org/ 10.1097/ID.0000000000000634. Gruber, R., Bosshardt, D.D., 2014. Dental implantology and implants—tissue interface. Stem Cell Biology and Tissue Engineering in Dental Sciences. Academic Press, pp. 735 747. Available from: https://doi.org/10.1016/B978-0-12-397157-9.00078-3. Guillaume, B., 2016. Les implants dentaires: revue. Morphologie . Available from: https:// doi.org/10.1016/j.morpho.2016.02.002. Gulati, K., Prideaux, M., Kogawa, M., Lima-Marques, L., Atkins, G.J., Findlay, D.M., et al., 2017. Anodized 3D printed titanium implants with dual micro- and nano-scale topography promote interaction with human osteoblasts and osteocyte-like cells. J. Tissue Eng. Regen. Med. 11, 3313 3325. Available from: https://doi.org/10.1002/ term.2239. Han, A., Tsoi, J., Matinlinna, J., Chen, Z., 2017. Influence of grit-blasting and hydrofluoric acid etching treatment on surface characteristics and biofilm formation on zirconia. Coatings 7, 130. Available from: https://doi.org/10.3390/coatings7080130. Han, A., Tsoi, J.K.H., Matinlinna, J.P., Zhang, Y., Chen, Z., 2018. Effects of different sterilization methods on surface characteristics and biofilm formation on zirconia in vitro. Dent. Mater. 34, 272 281. Available from: https://doi.org/10.1016/j.dental.2017.11.012. Hedia, H.S., Fouda, N., 2013. A new design of dental implant coating using functionally graded material. Materialprufung 55, 765 771. Available from: https://doi.org/10.3139/ 120.110499. Held, U., Rohner, D., Rothamel, D., 2013. Early loading of hydrophilic titanium implants inserted in low-mineralized (D3 and D4) bone: one year results of a prospective clinical trial. Head Face Med. 9, 37. Available from: https://doi.org/10.1186/1746-160X-9-37. Herrero-Climent, M., Romero Ruiz, M.M., Calvo, P.L., Santos, J.V.R., Perez, R.A., Gil Mur, F.J., 2018. Effectiveness of a new dental implant bioactive surface: histological and histomorphometric comparative study in minipigs. Clin. Oral Investig. 22, 1423 1432. Available from: https://doi.org/10.1007/s00784-017-2223-y. Hong, J., Kurt, S., Thor, A., 2013. A hydrophilic dental implant surface exhibit thrombogenic properties in vitro. Clin. Implant Dent. Relat. Res. 15, 105 112. Available from: https:// doi.org/10.1111/j.1708-8208.2011.00362.x.
Dental implants materials and surface treatments 595 Hotchkiss, K.M., Ayad, N.B., Hyzy, S.L., Boyan, B.D., Olivares-Navarrete, R., 2017. Dental implant surface chemistry and energy alter macrophage activation in vitro. Clin. Oral Implants Res. 28, 414 423. Available from: https://doi.org/10.1111/clr.12814. Ivanoff, C.-J., Widmark, G., Johansson, C., Wennerberg, A., n.d. Histologic evaluation of bone response to oxidized and turned titanium micro-implants in human jawbone. Int. J. Oral Maxillofac. Implants 18, 341 348. Jaffin, R.A., Berman, C.L., 1991. The excessive loss of Branemark fixtures in type IV bone: a 5-year analysis. J. Periodontol. 62, 2 4. Available from: https://doi.org/10.1902/ jop.1991.62.1.2. Khurshid, Z., Zafar, M., Qasim, S., Shahab, S., Naseem, M., AbuReqaiba, A., et al., 2015. Advances in nanotechnology for restorative dentistry. Materials (Basel) 8, 717 731. Available from: https://doi.org/10.3390/ma8020717. Khurshid, Z., Naseem, M., Sheikh, Z., Najeeb, S., Shahab, S., Zafar, M.S., 2016. Oral antimicrobial peptides: types and role in the oral cavity. Saudi Pharm. J. Available from: https://doi.org/10.1016/j.jsps.2015.02.015. Khurshid, Z., Najeeb, S., Mali, M., Moin, S.F., Raza, S.Q., Zohaib, S., et al., 2017a. Histatin peptides: pharmacological functions and their applications in dentistry. Saudi Pharm. J. Available from: https://doi.org/10.1016/j.jsps.2016.04.027. Khurshid, Z., Naseem, M., Asiri, F.Y.I., Mali, M., Sannam Khan, R., Sahibzada, H.A., et al., 2017b. Significance and diagnostic role of antimicrobial cathelicidins (LL-37) peptides in oral health. Biomolecules 7. Available from: https://doi.org/10.3390/biom7040080. Khurshid, Z., Zafar, M.S., Naseem, M., Khan, R.S., Najeeb, S., 2018. Human oral defensins antimicrobial peptides: a future promising antimicrobial drug. Curr. Pharm. Des. 24, 1130 1137. Available from: https://doi.org/10.2174/1381612824666180403114615. Khurshid, Z., Husain, S., Alotaibi, H., Rehman, R., Zafar, M.S., Farooq, I., et al., 2019. Novel techniques of scaffold fabrication for bioactive glasses. Biomed. Ther. Clin. Appl. Bioact. Glasses 497 519. Available from: https://doi.org/10.1016/B978-0-08-102196-5.00018-5. Klokkevold, P.R., Johnson, P., Dadgostari, S., Davies, J.E., Caputo, A., Nishimura, R.D., 2001. Early endosseous integration enhanced by dual acid etching of titanium: a torque removal study in the rabbit. Clin. Oral Implants Res. 12, 350 357. Available from: https://doi.org/10.1034/j.1600-0501.2001.012004350.x. Lang, N.P., Matinlinna, J.P., 2014. Titanium in implant dentistry. In: Matinlinna, J.P. (Ed.), Handbook of Oral Biomaterials. Pan Stanford Publishing, Singapore, pp. 339 358. Available from: https://doi.org/10.1201/b15644. Le Guéhennec, L., Soueidan, A., Layrolle, P., Amouriq, Y., 2007. Surface treatments of titanium dental implants for rapid osseointegration. Dent. Mater. Available from: https:// doi.org/10.1016/j.dental.2006.06.025. Lin, D., Li, Q., Li, W., Zhou, S., Swain, M.V., 2009. Design optimization of functionally graded dental implant for bone remodeling. Compos. Part B Eng. 40, 668 675. Available from: https://doi.org/10.1016/J.COMPOSITESB.2009.04.015. Lung, C.Y.K., Matinlinna, J.P., 2012. Aspects of silane coupling agents and surface conditioning in dentistry: an overview. Dent. Mater. 28, 467 477. Available from: https:// doi.org/10.1016/j.dental.2012.02.009. Mabboux, F., Ponsonnet, L., Morrier, J.J., Jaffrezic, N., Barsotti, O., 2004. Surface free energy and bacterial retention to saliva-coated dental implant materials—an in vitro study. Colloids and Surfaces B: Biointerfaces 39 (4), 199 205. Marei, M.K., El Backly, R.M., 2018. Dental mesenchymal stem cell-based translational regenerative dentistry: from artificial to biological replacement. Front. Bioeng. Biotechnol. 6, 49. Available from: https://doi.org/10.3389/fbioe.2018.00049.
596 Advanced Dental Biomaterials Mohseni, E., Zalnezhad, E., Bushroa, A.R., 2014. Comparative investigation on the adhesion of hydroxyapatite coating on Ti-6Al-4V implant: a review paper. Int. J. Adhes. Adhes. 48, 238 257. Available from: https://doi.org/10.1016/j.ijadhadh.2013.09.030. Molitor, P., Barron, V., Young, T., 2001. Surface treatment of titanium for adhesive bonding to polymer composites: a review. Int. J. Adhes. Adhes. 21, 129 136. Available from: https://doi.org/10.1016/S0143-7496(00)00044-0. Montero, J., Bravo, M., Guadilla, Y., Portillo, M., Blanco, L., Rojo, R., et al., 2015. Comparison of clinical and histologic outcomes of zirconia versus titanium implants placed in fresh sockets: a 5-month study in Beagles. Int. J. Oral Maxillofac. Implants 30, 773 780. Available from: https://doi.org/10.11607/jomi.3668. Mustafa, A., Lung, C.Y.K., Mustafa, N.S., Mustafa, B.A., Kashmoola, M.A., Zwahlen, R.A., et al., 2016. EPA-coated titanium implants promote osteoconduction in white New Zealand rabbits. Clin. Oral Implants Res. 27, 303 309. Available from: https://doi.org/10.1111/ clr.12525. Najeeb, S., Khurshid, Z., Matinlinna, J.P., Siddiqui, F., Nassani, M.Z., Baroudi, K., 2015. Nanomodified peek dental implants: bioactive composites and surface modification—a review. Int. J. Dent. Available from: https://doi.org/10.1155/2015/381759. Najeeb, S., Bds, Z.K., Bds, S.Z., Bds, M.S.Z., 2016a. Bioactivity and osseointegration of PEEK are inferior to those of titanium: a systematic review. J. Oral Implantol. 42, 512 516. Available from: https://doi.org/10.1563/aaid-joi-D-16-00072. Najeeb, S., Khurshid, Z., Agwan, A.S., Zafar, M.S., Alrahabi, M., Qasim, S.B., et al., 2016b. Dental applications of nanodiamonds. Sci. Adv. Mater. 8, 2064 2070. Available from: https://doi.org/10.1166/sam.2016.2993. Najeeb, S., Khurshid, Z., Zafar, M.S., Khan, A.S., Zohaib, S., Martı́, J.M.N., et al., 2016c. Modifications in glass ionomer cements: nano-sized fillers and bioactive nanoceramics. Int. J. Mol. Sci. Available from: https://doi.org/10.3390/ijms17071134. Najeeb, S., Zafar, M.S., Khurshid, Z., Siddiqui, F., 2016d. Applications of polyetheretherketone (PEEK) in oral implantology and prosthodontics. J. Prosthodont. Res. Available from: https://doi.org/10.1016/j.jpor.2015.10.001. Najeeb, S., Khurshid, Z., Siddiqui, F., Zohaib, S., Zafar, M.S., 2017a. Outcomes of dental implant therapy in patients with Down syndrome: a systematic review. J. Evid. Based Dent. Pract. Available from: https://doi.org/10.1016/j.jebdp.2017.05.003. Najeeb, S., Zafar, M.S., Khurshid, Z., Zohaib, S., Hasan, S.M., Khan, R.S., 2017b. Bisphosphonate releasing dental implant surface coatings and osseointegration: a systematic review. J. Taibah Univ. Med. Sci. 12, 369 375. Available from: https://doi.org/ 10.1016/J.JTUMED.2017.05.007. Nathanael, A.J., Oyane, A., Nakamura, M., Koga, K., Nishida, E., Tanaka, S., et al., 2018. Calcium phosphate coating on dental composite resins by a laser-assisted biomimetic process. Heliyon 4, e00734. Available from: https://doi.org/10.1016/j.heliyon.2018.e00734. Niinomi, M., 1998. Mechanical properties of biomedical titanium alloys. Mater. Sci. Eng. A 243 (1 2), 231 236. Norton, M.R., Gamble, C., 2001. Bone classification: an objective scale of bone density using the computerized tomography scan. Clin. Oral Implants Res. 12, 79 84. Available from: https://doi.org/10.1034/j.1600-0501.2001.012001079.x. Oliva, J., Oliva, X., Oliva, J.D., n.d. Five-year success rate of 831 consecutively placed zirconia dental implants in humans: a comparison of three different rough surfaces. Int. J. Oral Maxillofac. Implants 25, 336 344. https://doi.org/10.1590/S0103-50532007000100020 Osman, R.B., Swain, M.V., 2015. A critical review of dental implant materials with an emphasis on titanium versus zirconia. Materials (Basel) . Available from: https://doi.org/ 10.3390/ma8030932.
Dental implants materials and surface treatments 597 Petersen, R., 2016. Carbon fiber biocompatibility for implants. Fibers 4, 1. Available from: https://doi.org/10.3390/fib4010001. Piattelli, A., Degidi, M., Paolantonio, M., Mangano, C., Scarano, A., 2003. Residual aluminum oxide on the surface of titanium implants has no effect on osseointegration. Biomaterials 24, 4081 4089. Available from: https://doi.org/10.1016/S0142-9612(03)00300-4. Qasim, S.B., Najeeb, S., Delaine-Smith, R.M., Rawlinson, A., Ur Rehman, I., 2017. Potential of electrospun chitosan fibers as a surface layer in functionally graded GTR membrane for periodontal regeneration. Dent. Mater. 33, 71 83. Available from: https://doi.org/ 10.1016/J.DENTAL.2016.10.003. Rees, J.S., Jacobsen, P.H., 1993. The elastic moduli of enamel and dentine. Clini. Mater. 14 (1), 35 39. Rho, J.Y., Ashman, R.B., Turner, C.H., 1993. Young’s modulus of trabecular and cortical bone material: ultrasonic and microtensile measurements. J. Biomech. 26 (2), 111 119. Roessler, S., Born, R., Scharnweber, D., Worch, H., Sewing, A., Dard, M., 2001. Biomimetic coatings functionalized with adhesion peptides for dental implants. J. Mater. Sci. Mater. Med. 871 877. Available from: https://doi.org/10.1023/A:1012807621414. Rupp, F., Gittens, R.A., Scheideler, L., Marmur, A., Boyan, B.D., Schwartz, Z., et al., 2014. A review on the wettability of dental implant surfaces I: theoretical and experimental aspects. Acta Biomater. 10, 2894 2906. Available from: https://doi.org/10.1016/J. ACTBIO.2014.02.040. Ryu, H.-S., Namgung, C., Lee, J.-H., Lim, Y.-J., 2014. The influence of thread geometry on implant osseointegration under immediate loading: a literature review. J. Adv. Prosthodont. 6, 547 554. Available from: https://doi.org/10.4047/jap.2014.6.6.547. Sandler, J., Werner, P., Shaffer, M.S.P., Demchuk, V., Altstädt, V., Alan, H., 2002. Windle. Carbon-nanofibre-reinforced poly (ether ether ketone) composites. Compos. Part A: Appl. Sci. Manuf. 33 (8), 1033 1039. Smeets, R., Stadlinger, B., Schwarz, F., Beck-Broichsitter, B., Jung, O., Precht, C., et al., 2016. Impact of dental implant surface modifications on osseointegration. Biomed. Res. Int. Available from: https://doi.org/10.1155/2016/6285620. Telleman, G., Albrektsson, T., Hoffman, M., Johansson, C.B., Vissink, A., Meijer, H.J.A., et al., 2010. Peri-implant endosseous healing properties of dual acid-etched miniimplants with a nanometer-sized deposition of CaP: a histological and histomorphometric human study. Clin. Implant Dent. Relat. Res. 12, 153 160. Available from: https://doi.org/10.1111/j.1708-8208.2009.00147.x. Urban, R.M., Jacobs, J.J., Tomlinson, M.J., Gavrilovic, J., Black, J., Peoc’h, M., 2000. Dissemination of wear particles to the liver, spleen, and abdominal lymph nodes of patients with hip or knee replacement. J. Bone Joint Surg. Am. 82, 457 476. Urban, R.M., Tomlinson, M.J., Hall, D.J., Jacobs, J.J., 2004. Accumulation in liver and spleen of metal particles generated at nonbearing surfaces in hip arthroplasty. J. Arthroplasty 19, 94 101. Available from: https://doi.org/10.1016/J.ARTH.2004.09.013. Villard, N., Seneviratne, C., Tsoi, J.K.H., Heinonen, M., Matinlinna, J., 2015. Candida albicans aspects of novel silane system-coated titanium and zirconia implant surfaces. Clin. Oral Implants Res. 26, 332 341. Available from: https://doi.org/10.1111/clr.12338. Viña, J.A., El-Alami, M., Gambini, J., Borras, C., Viña, J., Peñarrocha, M.A., 2014. Application of mesenchymal stem cells in bone regenerative procedures in oral implantology. A literature review. J. Clin. Exp. Dent. 6, e60 e65. Available from: https://doi. org/10.4317/jced.51186. Watari, F., Yokoyama, A., Saso, F., Uo, M., Kawasaki, T., 1997. Fabrication and properties of functionally graded dental implant. Compos. Part B Eng. 28, 5 11. Available from: https://doi.org/10.1016/S1359-8368(96)00021-2.
598 Advanced Dental Biomaterials Webster, T.J., Patel, A.A., Rahaman, M.N., Sonny Bal, B., 2012. Anti-infective and osteointegration properties of silicon nitride, poly(ether ether ketone), and titanium implants. Acta Biomater. 8, 4447 4454. Available from: https://doi.org/10.1016/j.actbio.2012.07.038. Wenz, H.J., Bartsch, J., Wolfart, S., Kern, M., n.d. Osseointegration and clinical success of zirconia dental implants: a systematic review. Int. J. Prosthodont. 21, 27 36. Xuereb, M., Camilleri, J., Attard, N., 2015. Systematic review of current dental implant coating materials and novel coating techniques. Int. J. Prosthodont. 28, 51 59. Available from: https://doi.org/10.11607/ijp.4124. Yabutsuka, T., Fukushima, K., Hiruta, T., Takai, S., Yao, T., 2018. Fabrication of bioactive fiber-reinforced PEEK and MXD6 by incorporation of precursor of apatite. J. Biomed. Mater. Res. Part B Appl. Biomater. 106, 2254 2265. Available from: https://doi.org/ 10.1002/jbm.b.34025. Zafar, M.S., Khurshid, Z., Almas, K., 2015. Oral tissue engineering progress and challenges. Tissue Eng. Regen. Med. 12, 387 397. Available from: https://doi.org/10.1007/s13770015-0030-6. Zafar, M.S., Farooq, I., Awais, M., Najeeb, S., Khurshid, Z., Zohaib, S., 2019. Bioactive surface coatings for enhancing osseointegration of dental implants. Biomed. Ther. Clin. Appl. Bioact. Glasses 313 329. Available from: https://doi.org/10.1016/B978-0-08102196-5.00011-2. Zhang, M., Matinlinna, J.P., 2012. E-glass fiber reinforced composites in dental applications. Silicon 4, 73 78. Available from: https://doi.org/10.1007/s12633-011-9075-x. Zhou, H., Lee, J., 2011. Nanoscale hydroxyapatite particles for bone tissue engineering. Acta Biomater. 7, 2769 2781. Available from: https://doi.org/10.1016/J.ACTBIO.2011.03.019. Zhu, X., Chen, J., Scheideler, L., Altebaeumer, T., Geis-Gerstorfer, J., Kern, D., 2004. Cellular reactions of osteoblasts to micron- and submicron-scale porous structures of titanium surfaces. Cells Tissues Organs 178, 13 22. Available from: https://doi.org/ 10.1159/000081089. Zinger, O., Anselme, K., Denzer, A., Habersetzer, P., Wieland, M., Jeanfils, J., et al., 2004. Time-dependent morphology and adhesion of osteoblastic cells on titanium model surfaces featuring scale-resolved topography. Biomaterials 25, 2695 2711. Available from: https://doi.org/10.1016/J.BIOMATERIALS.2003.09.111.
Graphene to improve the physicomechanical properties and bioactivity of the cements 22 Vinicius Rosa1,2, Francisco Javier Rodrı´guez-Lozano3 and Kyung-san Min4 1 Faculty of Dentistry, National University of Singapore, Singapore, Singapore, 2Centre for Advanced 2D Materials and Graphene Research Centre, National University of Singapore, Singapore, Singapore, 3School of Dentistry, University of Murcia, Murcia, Spain, 4 School of Dentistry, Chonbuk National University, Jeonju, South Korea Chapter Outline 22.1 Introduction 599 22.1.1 Graphene and its derivatives 599 22.2 Graphene to improve cementitious materials 22.3 Conclusion 609 References 609 22.1 600 Introduction 22.1.1 Graphene and its derivatives Graphene is a single atomic sheet of conjugated sp2 carbon atoms arranged in a honeycomb pattern (Geim and Novoselov, 2007). It has high electron mobility and thermal conductivity (Nair et al., 2008; Neto et al., 2009; Khare et al., 2007). Graphene films are atomically thin, flexible, and light (0.77 mg/m2) . They present very high elastic modulus and ultimate tensile strength. Moreover, graphene films have large specific surface areas (2630 m2/g) that can be chemically modified to provide new capabilities (Zhang et al., 2013b; Castro Neto et al., 2009; Bunch et al., 2008; Dreyer et al., 2010; Geim and Novoselov, 2007). Pristine graphene films can be produced in a relatively pure form by chemical vapor deposition (CVD), graphene oxide (GO) reduction, exfoliation of graphite, segregation from silicon carbide, and other methods (Faugeras et al., 2008; Zhou et al., 2009; Lotya et al., 2009). In the traditional CVD process a sacrificial substrate (Cu or Ni) is placed in a reaction chamber where precursor gases (H2 and CH4) are burned at elevated temperatures ( . 800 C). The gaseous pyrolysis induces the nucleation of carbon atoms on the sacrificial substrate which grows in large domains producing the graphene film (Lopez and Mittemeijer, 2004; Li et al., 2009a,b; Xie et al., 2015b). While the CVD process allows the production of films Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00022-0 Copyright © 2019 Elsevier Ltd. All rights reserved.
600 Advanced Dental Biomaterials with high quality and purity, the chemical methods (e.g., exfoliation) emerge as viable routes to produce graphene-based sheets on a large scale at low cost (Zheng and Kim, 2015; Morin et al., 2017; Geim and Novoselov, 2007). Graphene has two derivatives, namely, GO and reduced GO (rGO). The first is a highly oxidized and hydrophilic form of graphene (Dreyer et al., 2010; Zheng and Kim, 2015). GO has several oxygen-containing groups (e.g., OH, COOH) that allow the material to be dispersed in organic solvents, water, different matrices, and be combined with several molecules and compounds (Dreyer et al., 2010; Goenka et al., 2014; Liu et al., 2013a). The most widely used method to produce GO is the liquid phase exfoliation of graphite (Perera et al., 2012; Rosa et al., 2016). This low-cost method has been used for several years and relies on the charge transfer between the solvent molecules and graphite layers to induce the separation of graphene sheets with several micrometers in lateral size (Stankovich et al., 2007; Dreyer et al., 2010). Finally, rGO can be produced by removal (reduction) of the oxygen-containing groups from GO by several techniques (Goenka et al., 2014; Dreyer et al., 2010). 22.2 Graphene to improve cementitious materials The term “bioactive cement” is used for a variety of cements with distinct chemical compositions (e.g., hydroxyapatite, calcium silicate, β-tricalcium phosphate, and others) that can induce and increase the mineralization processes in cells. These materials often release calcium ions and increase the production of calcium hydroxides. Moreover, they can induce the formation of an interfacial layer between the cement and host tissue and the formation of apatite crystals over the surface of the materials (Parirokh et al., 2018). Bioactive cements are widely used in vital pulp therapies (direct pulp capping, indirect pulp capping, pulpotomy), apexification, regenerative endodontics, perforation repair, root canal filling, root-end filling, treatment of root fractures, filling of bone defects, and cementation of prostheses (Main et al., 2004; Caron et al., 2014; Parirokh et al., 2018; Torabinejad et al., 2018). One of the most classical bioactive cements is the mineral trioxide aggregate (MTA), first introduced by Torabinejad in the 1990s. This is a derivative of Portland cement and was approved by the Food and Drug Administration in 1997 (Parirokh and Torabinejad, 2010a). The properties of MTA (including its biocompatibility, bioactivity, osteoconductivity) have fostered a race to develop new bioceramics or bioactive cements incorporating MTA (Torabinejad et al., 2018). Some of these are Biodentine, BioRoot RCS, calcium-enriched mixture cement, Endo-CPM, Endocem, EndoSequence, EndoBinder, EndoSeal MTA, GuttaFlow Bioseal, iRoot, MicroMega MTA, MTA Bio, MTA Fillapex, MTA HP Repair, MTA Plus, NeoMTA Plus, OrthoMTA, Quick-Set, RetroMTA, Tech Biosealer, and TheraCal LC (Torabinejad et al., 2018). One of the main advantages of bioceramics is their ability to dissociate calcium ions. Nonetheless, the materials differ dramatically in the amounts of calcium that are released and in the kinetics of this process. For instance, Biodentine releases
Graphene to improve the physicomechanical properties and bioactivity of the cements 601 approximate 95 ppm of calcium ions in the first 3 hours decreasing to 18 ppm at 24 hours (Gandolfi et al., 2015), while MTA releases approximately 5 and 25 ppm for the same time points (Natu et al., 2015). Although the values of calcium release may change according to the methodology used, the literature agrees on the potential of these materials to alkalinize the environment and increase the pH to values around 11 and above (Natu et al., 2015; Prati and Gandolfi, 2015; Dubey et al., 2017). Together, the calcium release and alkaline potential can increase the phosphatase activity and the expression of several genes related to mineralization processes such as osteocalcin, bone sialoprotein, and dentine sialophosphoprotein (Hakki et al., 2009; Min et al., 2009; Maeda et al., 2010; Natu et al., 2015). Despite these key features, these materials can present some essential clinical disadvantages. The main shortcoming may include long setting time, the presence of toxic elements in the material composition, the potential for discoloration, difficult handling characteristics, and high material cost (Parirokh and Torabinejad, 2010b; Natu et al., 2015). Considering the setting time, it can be as long as 18 minutes for MTA while the final set of this type of material can take up to 170 minutes (Natu et al., 2015; Gandolfi et al., 2009). Besides, some of these cements may present high solubility (up to 4.6% of weight loss for Biodentine after 28 days) and low values for hardness (26.9 VHN for ProRoot MTA) (Kaup et al., 2015). Moreover, the materials are prone to lose compressive strength upon contamination with the moisture of blood and modest pushout bond strength (Sheykhrezae et al., 2017; Dubey et al., 2017). The potential for tooth discoloration associated with the use of biocements, especially MTA, is a clear clinical disadvantage (Torabinejad et al., 2018). Recent investigations have shown that bismuth oxide is the major culprit of this discoloration. Consequently, manufacturers have developed alternative bioactive cement in a similar composition that should not cause tooth discoloration. In this context NeoMTA Plus contains tantalum oxide as radiopacifier instead of bismuth oxide, avoiding the discoloration (Tomas-Catala et al., 2017, 2018). Another limitation of bioactive endodontic cements is the difficulty of removing the material from the root canals during retreatment. Although some new bioactive endodontic sealers such as MTA Fillapex, EndoSequence BC sealer, MTA Plus, and iRoot SP can be more easily removed (Torabinejad et al., 2018), the presence of residual filling material after retreatment remains as a concern (Oltra et al., 2017). Thus the development of materials and strategies that can improve these aspects without compromising the bioactivity of cements is of high interest. In recent years several approaches have been developed to improve the properties of bioactive cements. Some of them focus on the modification of the liquids that are used to mix the powder particles. For instance, the addition of propylene glycol has been shown to increase the flowability and improve the handling of the materials, but it increases setting time and decreases hardness (Natu et al., 2015). Liquid extract of Casearia sylvestris increased the antimicrobial properties of MTA but interfered in some of its physicochemical properties (Natu et al., 2015; Cavenago et al., 2017). Also, thickening agents instead of water in injectable calcium silicate endodontic sealers are commonly used to create gel-like pastes that improve their handling (Ha et al., 2017). In the same line new
602 Advanced Dental Biomaterials MTA-based with high-plasticity materials have been introduced. The new strategy relies on the replacement of the distilled water by a liquid containing water Check and an organic plasticizer that provides higher plasticity, improving handling and placement of the material to the repair site (Cintra et al., 2017). Other strategies focus on the modification of the powders by the addition of fractions of new components to it. For instance, the addition of wollastonite (CaSiO3), a raw calcium metasilicate composed of calcium oxide and silicon dioxide, shortened the setting time but decreased the compressive strength of a MTA cement (Flores-Ledesma et al., 2017). The presence of MgO or ZnO improved the biological properties (e.g., cell viability) of glass ionomer cements but failed to increase compressive strength (Kim et al., 2015). The addition of zirconia to hydroxyapatite increased the fracture toughness and flexural strength by almost three times higher compared to the unmodified materials (Kong et al., 2010). The insertion of niobium (Nb2O5) oxide in bioactive cements can also increase the mineralization (Gomes-Cornelio et al., 2017). Also, the incorporation of prereacted glass ionomer fillers in root repair cements demonstrated significant antibacterial effects against endodontic pathogens, multiple ion-releasing abilities, and significantly shorter initial and final setting times in comparison to MTA (Yassen et al., 2016). Regardless of the achievements, there are several opportunities to further improve the properties of bioactive cements. Recently, nanomaterials have been explored to improve several physical and mechanical properties of industrial and biological cements. For instance, the addition of multiwalled carbon nanotubes (CNTs) can improve both the compressive strength and modulus of rupture of cements (Musso et al., 2009). The addition of 0.02 wt.% of CNTs to cement mortar increased its compressive strength by 11% (Mohsen et al., 2017). Improvements in these properties were observed when CNT and carbon nanofibers increased the compressive strengths of ordinary Portland cement mortar by 154% and 217% compared to the unmodified material (Yazdani and Mohanam, 2016). The addition of 0.25 wt.% CNTs to Portland cement increased in 55% the flexural strength of the set material (from 10.4 to 16.2 MPa) (Mohsen et al., 2017). The addition of multiwalled CNTs (0.25%) to type I Portland cement resulted in a composite with compressive and tensile strength higher than the control (20% and 40%, respectively) (Jang et al., 2016). Similar improvements can be observed with the addition of single-walled CNTs which increased the hardness by the hydration of cements at an early age (Makar et al., 2005). In terms of setting reaction different alternatives have been shown to effectively speed up the setting reaction while decreasing the setting time. For instance, nano-SiO2 added to Portland cement (42.5 grade) induces the early precipitation of calcium silicate hydrate shortening the induction period. Notably, the initial setting time decreased significantly with the presence of 5 wt.% of nano-SiO2 in the mix (Qing et al., 2007). All these enhancements observed in the past with different nanomaterials have led researchers to explore graphene family materials to improve the physical, mechanical, and biological properties of cements. The literature is vast on the enhancements provided by GO to improve the properties of cementitious materials. For instance, the presence of 0.03 wt.% of GO increased the tensile and compressive
Graphene to improve the physicomechanical properties and bioactivity of the cements 603 strength of Portland cement by 53% and 46%, respectively (Gong et al., 2015). Similar enhancements have been observed with the addition of 0.05 wt.% GO to cement that increased the flexural strength by 59% and the compressive strength by 33% (Pan et al., 2012). In fact very small quantities of GO (e.g., 0.01 wt.%) can significantly increase both the compressive and flexural strengths of Portland cements by 23% and 57%. Nevertheless, the values are even higher when the concentration rises to 0.06 wt.% resulting in increases of 72% in the compressive strength and 63% in the flexural strength after 7 days (Lv et al., 2014). The improvements provided by GO-based materials are not limited to Portland cements. Bioceramics, such as hydroxyapatite, β-tricalcium phosphate, and others, also experience improvements in their properties when modified with the different forms of graphene. One of the most explored bioactive materials for bone regeneration is hydroxyapatite. It has high osteoconductivity potential and tissue compatibility (Jang et al., 2014; LeGeros, 2002). The commercially available versions of hydroxyapatite can induce bone repair in clinical settings and increase bone to implant contact (Ogose et al., 2006; Jang et al., 2014; Tonetto et al., 2015b). Nonetheless, this material can be very brittle, difficult to shape, and present slow resorption rates (Porter et al., 2009). Notably, several studies have shown that graphene family materials can overcome some of these disadvantages. One of the potential uses of graphene is to improve the physical properties of hydroxyapatite-based materials. The presence of 2 and 5 wt.% of GO in a hydroxyapatite coating can increase the adhesion of the coating to titanium from 1.5 to 2.7 and 3.3 MPa, respectively. Notably, the GO-modified hydroxyapatite also presented higher corrosion resistance compared to the unmodified version of the material (Liu et al., 2014a). Increased corrosion resistance has also been observed in ternary GO chitosan hydroxyapatite (GO CS HA) coating. The corrosion current density of titanium decreased from 6823 to 248 nA/cm2 after being coated with chitosan hydroxyapatite and to 13.5 nA/cm2 when protected by GO CS HA (Shi et al., 2016). It is also possible to improve the mechanical properties of several bioactive cements. For instance, GO nanoflakes (0.5%) added a scaffold of hydroxyapatite, and gelatin increased both the compressive strength (from 37 to 52 MPa) and the yield strength (from 12 to 47 MPa) (Nair et al., 2015). Also, the addition of 0.5 wt.% graphene to 58S bioglass promoted a twofold increase in the compressive strength (from 23.6 to 48.7 MPa) and increased the fracture toughness from 1.4 to 1.9 MPa m1/2 (Gao et al., 2014). The addition of 1 wt.% rGO to hydroxyapatite increases its fracture toughness by more than 200% and elastic modulus by 47% (Yi et al., 2013). Similarly, enhancements were observed by the addition of 1 wt.% of graphene nanosheets (GNS) to hydroxyapatite that nearly doubled the fracture toughness and increased hardness in almost 50% (Zhang et al., 2013a). The addition of 3 wt.% of GNS to EndocemZr increased the hardness from 8 to 15 KHN (Fig. 22.1) (Dubey et al., 2017). Similarly, 1 wt.% rGO increased the hardness of the hydroxyapatite from 322 to 425 HV (Baradaran et al., 2014). Finally, the addition of 1.5 wt.% rGO increased both the elastic modulus and fracture toughness of hydroxyapatite by 40% and 86%, respectively (Zanin et al., 2013).
604 Advanced Dental Biomaterials Figure 22.1 The addition of given concentrations GNS to BIO and ECZ (top panels) can decrease the setting time (A) and increase the hardness of the materials (B). Nonetheless, the GNS negatively affects the pushout bond strength of ECZ to dentine (C). BIO, Biodentine; ECZ, EndocemZR; GNS, graphene nanosheets. Source: Adapted from Dubey, N., Rajan, S.S., Bello, Y.D., Min, K.-S., Rosa, V., 2017. Graphene nanosheets to improve physico-mechanical properties of bioactive calcium silicate cements. Materials, 10, 606. These enhancements in mechanical properties are likely to be a result of the crack deflection, bridging, and sheet pullout through promoted by the different forms of graphene inside the cementitious matrix that delays or prevents the crack propagation (Baradaran et al., 2014; Li et al., 2016). For instance, the pullout energy of graphene from hydroxyapatite varies from 3 to 40 J/m, which is higher than the fracture energy of bulk hydroxyapatite (1 J/m). This forces the cracks to propagate within the hydroxyapatite matrix and not along the graphene/hydroxyapatite interface. In addition, graphene induces the formation of smaller hydroxyapatite crystals that could potentially result in a fine-grain strengthening mechanism (Liu et al., 2013b). The changes in properties of the bioactive cements by the addition of graphene family materials are dependent on the concentration of the materials that are added to the cements. The addition of 0.5 wt.% of GO into 45S5 Bioglass increases the fracture toughness from B0.5 to 0.8 while the addition of 1 wt.% increases it further to 1.2 MPa m1/2 (Li et al., 2016). It deserves to be noted that the enhancements in mechanical properties are dose-dependent, and the increase of concentrations of graphene above a certain threshold does not translate into further enhancements in the set materials (Lv et al., 2013, 2014). It has been shown that the mechanical properties of ordinary Portland cement type 42.5R decrease when GO is added to a concentration higher than 0.03 wt.% (Fig. 22.2). Whereas 0.03% increases the tensile strength from 2.47 to 2.93 MPa, the addition of 0.06% GO nanosheets
Figure 22.2 The addition of GO nanosheets increases the mechanical properties of Portland cement in a dose-dependent manner: increases above 0.03% result in decreased tensile and flexural strength. GO, Graphene oxide. Source: Lv S.,Ting S., Liu J., and ZhouQ., Use of graphene oxide nanosheets to regulate the microstructure of hardened cement paste to increase its strength and toughness, Cryst Eng Comm 16, 2014, 8508.
606 Advanced Dental Biomaterials decreases it to 2.4 MPa. Similar behavior has been observed for the flexural strength that increases from 8.5 to 9.6 MPa for 0.03% and decreases to 7.2 with 0.06% (Lv et al., 2013). Similarly, GNS in high concentrations (5 and 7 wt.%) decreased the hardness of both Biodentine and EndocemZr compared to concentrations equal to or lower than 3 wt.% (Dubey et al., 2017). Moreover, carbon-based materials influence the formation of hydration products by accelerating the rate of hydration processes, particularly at the early stages (Manzur and Yazdani, 2014; Lv et al., 2014). Hence, the addition of these materials to cement can impact positively or negatively on the time taken to set the cements. It has been observed that GNS (1 and 3 wt.%) can decrease the setting time of both Biodentine and EndocemZr. Interestingly, the addition of 5 and 7 wt.% of nanosheets results in setting times that are similar to the unmodified cements (Dubey et al., 2017). This may be related to the plausible agglomeration of the carbonaceous materials around the cement grains that results in partial hydration (Siddique and Mehta, 2014). Besides the improvements in physicomechanical properties, it is also desirable to improve the bioactivity of cements. This is important since it can speed up the formation of bone and enhance the amount of the tissue formed. These two aspects can enhance the development of alternative bone substitutes to overcome the main limitations related to autologous and allogenous grafts (e.g., high cost, batch-to-batch variability, and limited availability). In this regard several studies have shown that graphene and its derivatives can improve the bioactivity of several bioactive cements. Graphene and its derivatives have been shown to enhance the osteogenic differentiation of cells. Substrates and scaffolds made of pure graphene produced by CVD have been shown to promote the spontaneous osteogenic differentiation of several types of cells (e.g., bone marrow stem cells, periodontal ligament stem cells, and dental pulp stem cells). The material, in both 2D and 3D arrangements, increased the expression of RUNX2, collagen I, osteopontin, and osteocalcin at both genetic and protein level (Fig. 22.3). These are key markers involved in the early and final stages of the osteoblastic differentiation (Xie et al., 2015a, 2017). Recently, a substrate made of GO has been shown to increase the expression of several genes commonly expressed by cells that secrete mineralized matrix such as DSPP, DMP-1, RUNX2, and collagen I (Fig. 22.4) (Rosa et al., 2016). Notably, the differentiation was further magnified in the presence of osteogenic induction media in the presence of graphene compared to when the chemical inductors are used alone (Nayak et al., 2011; Lee et al., 2011; Dubey et al., 2015). For these reasons, graphene has been extensively investigated to increase the biological properties of cementitious materials. Graphene and its derivatives have been combined with scaffolds to increase their potential for osteogenesis. Collagen sponge-form scaffolds modified with GO can adsorb considerable amounts of albumin and lysozyme and induce the formation of bone in vivo. When the GO-modified scaffolds were implanted into bone sockets, they promoted the formation of new bone while the control scaffold presented lesser bone formation and was mostly filled with connective tissue (Nishida et al., 2016). These notable effects were further explored to improve scaffolds and particles made of biocements. For example, the combination of GO to a
Graphene to improve the physicomechanical properties and bioactivity of the cements 607 Figure 22.3 CVD-grown graphene induces the osteoblastic differentiation of stem cells. Cells on graphene films (2DGp) or scaffolds (3DGp) presented high expression of osteocalcin (in green). Glass substrate (Gl) was used as a control for 2DGp (Xie et al., 2015b). CVD, Chemical vapor deposition. gelatin/hydroxyapatite scaffold has increased the alkaline phosphatase activity and the protein expression of osteopontin compared to the unmodified scaffold (Nair et al., 2015). Bone marrow mesenchymal stem cells (MSCs) cultured in β-tricalcium phosphate scaffolds combined with GO flakes presented increased alkaline phosphatase activity and higher expression of genes related to osteogenic differentiation, probably via the activation of the canonical Wnt signalling pathway (Wu et al., 2015). Scaffolds made of GO/calcium phosphate nanocomposites induced higher levels of calcification, alkaline phosphatase activity, and increased the protein expression of osteocalcin compared to cells treated with GO or calcium phosphate alone (Tatavarty et al., 2014). Graphene can also be combined with bioceramics and cements to increase the bioactivity of these materials. For instance, colloidal dispersion of rGO-coated hydroxyapatite increased the protein expression of osteopontin and osteocalcin calcium nodule deposition and alkaline phosphatase activity of MSCs compared to rGO or hydroxyapatite alone (Lee et al., 2015a). The addition of GNS to EndocemZr and Biodentine increased the mineralization potential of dental pulp stem cells
608 Advanced Dental Biomaterials Figure 22.4 (A and B) Dental pulp stem cells can efficiently adhere and proliferate on GO-based substrate. (C) The material also increased the expression of several genes commonly expressed by mineral-secreting cells. Source: Adapted from Rosa, V., Xie, H., Dubey, N., Madanagopal, T.T., Rajan, S.S., Morin, J.L.P., et al., 2016. Graphene oxide-based substrate: physical and surface characterization, cytocompatibility and differentiation potential of dental pulp stem cells. Dent. Mater. 32, 1019 1025. (Dubey et al., 2017). GO-modified β-tricalcium phosphate scaffold induced higher bone formation in critical-sized calvarial defects created in rabbits compared to the defects filled with the unmodified scaffold. The bone volume/total volume ratios were 44% for the GO-modified scaffold and 30% for β-tricalcium phosphate scaffold used as control (Wu et al., 2015). Similarly, rGO/hydroxyapatite grafts increased the bone density from 26% (hydroxyapatite alone) to 52%. The increment was even higher when compared to the value obtained for the untreated control (17.7%) after 4 weeks (Lee et al., 2015b). Hydroxyapatite/GNS composites increased the proliferation of osteoblast cells when used as a coating on titanium substrates. Interestingly, the random incorporation of curly graphene sheets allowed for higher protein absorption (fibronectin) and provided active sites for cell binding (Liu et al., 2014b).
Graphene to improve the physicomechanical properties and bioactivity of the cements 609 The enhancements on the bioactivity of materials modified with graphene have different origins. For hydroxyapatite, it has been observed that the presence of GNS induced the formation of uniform apatite layers that are thicker than hydroxyapatite alone. The GNS change the spatial distribution of the crystals that grow closely or even inside the porous hydroxyapatite, while in the GNS/hydroxyapatite composites, they grow preferentially on the surface of material (Zhang et al., 2013a). Furthermore, GO increases the nucleation and crystallization of hydroxyapatite. This process results in GO/hydroxyapatite composites with dense hydroxyapatite crystals (Gao et al., 2015). Notably, the presence of graphene nanofillers can increase the formation of thick apatite surface of hydroxyapatite under simulated body fluid (Zhang et al., 2013a). The changes in the compositions and characteristics of hydroxyapatite lead to different outcomes that eventually can increase the bioactivity of the composites. For instance, graphene/hydroxyapatite composites are less stable than hydroxyapatite alone. This leads to the fast dissolution of calcium ions from the composites (Janković et al., 2015). The lower crystallinity of graphene-modified materials can also increase the dissolution and the release of ions from the composites (Mohandes and Salavati-Niasari, 2014a). Similarly, composites made of hydroxyapatite/GO with chitosan or gelatin can release more calcium ions compared to pure hydroxyapatite particles (Mohandes and SalavatiNiasari, 2014a,b). These can have a biological impact since these ions can induce and enhance the osteoblastic differentiation (Natu et al., 2015; Jung et al., 2010). 22.3 Conclusion The different forms of graphene can be functionalized and combined with several biomolecules and biomaterials, increasing their potential to improve bioactive cements. In this area, graphene has been used as a physical additive or to modify the overall chemical composition of cementitious materials. The resultant materials often present improved physical and mechanical properties and can present certain characteristics that allow for higher bioactivity. Despite these exciting achievements, it is still necessary to evaluate the potential risks and long-term performance of the graphene-modified cements. Nevertheless, the findings obtained so far attest to the potential of graphene to improve the physical, chemical, mechanical, and biological properties of cements. References Baradaran, S., Moghaddam, E., Basirun, W.J., Mehrali, M., Sookhakian, M., Hamdi, M., et al., 2014. Mechanical properties and biomedical applications of a nanotube hydroxyapatite-reduced graphene oxide composite. Carbon N.Y. 69, 32 45. Bunch, J.S., Verbridge, S.S., Alden, J.S., Van Der Zande, A.M., Parpia, J.M., Craighead, H. G., et al., 2008. Impermeable atomic membranes from graphene sheets. Nano Lett. 8, 2458 2462. Caron, G., Azerad, J., Faure, M.O., Machtou, P., Boucher, Y., 2014. Use of a new retrograde filling material (Biodentine) for endodontic surgery: two case reports. Int. J. Oral Sci. 6, 250 253.
610 Advanced Dental Biomaterials Castro Neto, A.H., Guinea, F., Peres, N.M.R., Novoselov, K.S., Geim, A.K., 2009. The electronic properties of graphene. Rev. Mod. Phys. 81, 109 162. Cavenago, B.C., Del Carpio-Perochena, A.E., Ordinola-Zapata, R., Estrela, C., Garlet, G.P., Tanomaru-Filho, M., et al., 2017. Effect of using different vehicles on the physicochemical, antimicrobial, and biological properties of white mineral trioxide aggregate. J. Endod. 43, 779 786. Cintra, L.T.A., Benetti, F., De Azevedo Queiroz, I.O., De Araujo Lopes, J.M., Penha De oliveira, S.H., Sivieri Araujo, G., et al., 2017. Cytotoxicity, biocompatibility, and biomineralization of the new high-plasticity MTA material. J. Endod. 43, 774 778. Dreyer, D.R., Park, S., Bielawski, C.W., Ruoff, R.S., 2010. The chemistry of graphene oxide. Chem. Soc. Rev. 39, 228 240. Dubey, N., Bentini, R., Islam, I., Cao, T., Castro Neto, A.H., Rosa, V., 2015. Graphene: a versatile carbon-based material for bone tissue engineering. Stem Cells Int. 2015, 1 12. Dubey, N., Rajan, S.S., Bello, Y.D., Min, K.-S., Rosa, V., 2017. Graphene nanosheets to improve physico-mechanical properties of bioactive calcium silicate cements. Materials 10, 606. Faugeras, C., Nerrière, A., Potemski, M., Mahmood, A., Dujardin, E., Berger, C., et al., 2008. Few-layer graphene on SiC, pyrolitic graphite, and graphene: a Raman scattering study. Appl. Phys. Lett. 92, 011914. Flores-Ledesma, A., Barcelo Santana, F., Bucio, L., Arenas-Alatorre, J.A., Faraji, M., Wintergerst, A.M., 2017. Bioactive materials improve some physical properties of a MTA-like cement. Mater. Sci. Eng. C Mater. Biol. Appl. 71, 150 155. Gandolfi, M.G., Iacono, F., Agee, K., Siboni, F., Tay, F., Pashley, D.H., et al., 2009. Setting time and expansion in different soaking media of experimental accelerated calciumsilicate cements and ProRoot MTA. Oral Surg. Oral. Med. Oral Pathol. Oral Radiol. Endod. 108, e39 e45. Gandolfi, M.G., Siboni, F., Botero, T., Bossù, M., Riccitiello, F., Prati, C., 2015. Calcium silicate and calcium hydroxide materials for pulp capping: biointeractivity, porosity, solubility and bioactivity of current formulations. J. Appl. Biomater. Funct. Mater. 13, 43 60. Gao, C., Liu, T., Shuai, C., Peng, S., 2014. Enhancement mechanisms of graphene in nano58S bioactive glass scaffold: mechanical and biological performance. Sci. Rep. 4, 4712. Gao, F., Xu, C., Hu, H., Wang, Q., Gao, Y., Chen, H., et al., 2015. Biomimetic synthesis and characterization of hydroxyapatite/graphene oxide hybrid coating on Mg alloy with enhanced corrosion resistance. Mater. Lett. 138, 25 28. Geim, A.K., Novoselov, K.S., 2007. The rise of graphene. Nat. Mater. 6, 183 191. Goenka, S., Sant, V., Sant, S., 2014. Graphene-based nanomaterials for drug delivery and tissue engineering. J. Control. Release 173, 75 88. Gomes-Cornelio, A.L., Rodrigues, E.M., Salles, L.P., Mestieri, L.B., Faria, G., GuerreiroTanomaru, J.M., et al., 2017. Bioactivity of MTA Plus, Biodentine and an experimental calcium silicate-based cement on human osteoblast-like cells. Int. Endod. J. 50, 39 47. Gong, K., Pan, Z., Korayem, A.H., Qiu, L., Li, D., Collins, F., et al., 2015. Reinforcing effects of graphene oxide on Portland cement paste. J. Mater. Civil Eng. 27, A4014010. Ha, W.N., Nicholson, T., Kahler, B., Walsh, L.J., 2017. Mineral trioxide aggregate—a review of properties and testing methodologies. Materials (Basel) 10. Hakki, S.S., Bozkurt, S.B., Hakki, E.E., Belli, S., 2009. Effects of mineral trioxide aggregate on cell survival, gene expression associated with mineralized tissues, and biomineralization of cementoblasts. J. Endod. 35, 513 519. Jang, C.H., Cho, Y.B., Choi, C.H., Jang, Y.S., Jung, W.K., Lee, J.K., 2014. Comparison of osteoconductivity of biologic and artificial synthetic hydroxyapatite in experimental mastoid obliteration. Acta Otolaryngol. 134, 255 259.
Graphene to improve the physicomechanical properties and bioactivity of the cements 611 Jang, S.H., Kawashima, S., Yin, H., 2016. Influence of carbon nanotube clustering on mechanical and electrical properties of cement pastes. Materials (Basel) 9, 220. Janković, A., Eraković, S., Mitrić, M., Matić, I.Z., Juranić, Z.D., Tsui, G.C.P., et al., 2015. Bioactive hydroxyapatite/graphene composite coating and its corrosion stability in simulated body fluid. J. Alloys Compd. 624, 148 157. Jung, G.Y., Park, Y.J., Han, J.S., 2010. Effects of HA released calcium ion on osteoblast differentiation. J. Mater. Sci. Mater. Med. 21, 1649 1654. Kaup, M., Schäfer, E., Dammaschke, T., 2015. An in vitro study of different material properties of Biodentine compared to ProRoot MTA. Head. Face. Med. 11. Khare, R., Mielke, S.L., Paci, J.T., Zhang, S., Ballarini, R., Schatz, G.C., et al., 2007. Coupled quantum mechanical/molecular mechanical modeling of the fracture of defective carbon nanotubes and graphene sheets. Phys. Rev. B 75, 075412. Kim, D.A., Abo-Mosallam, H., Lee, H.Y., Lee, J.H., Kim, H.W., Lee, H.H., 2015. Biological and mechanical properties of an experimental glass-ionomer cement modified by partial replacement of CaO with MgO or ZnO. J. Appl. Oral Sci. 23, 369 375. Kong, Y.-M., Kim, S., Kim, H.-E., Lee, I.-S., 2010. ChemInform abstract: reinforcement of hydroxyapatite bioceramic by addition of ZrO2 coated with Al2O3. ChemInform 31. Lee, W.C., Lim, C.H.Y.X., Shi, H., Tang, L.A.L., Wang, Y., Lim, C.T., et al., 2011. Origin of enhanced stem cell growth and differentiation on graphene and graphene oxide. ACS Nano 5, 7334 7341. Lee, J.H., Shin, Y.C., Jin, O.S., Kang, S.H., Hwang, Y.S., Park, J.C., et al., 2015a. Reduced graphene oxide-coated hydroxyapatite composites stimulate spontaneous osteogenic differentiation of human mesenchymal stem cells. Nanoscale 7, 11642 11651. Lee, J.H., Shin, Y.C., Lee, S.M., Jin, O.S., Kang, S.H., Hong, S.W., et al., 2015b. Enhanced osteogenesis by reduced graphene oxide/hydroxyapatite nanocomposites. Sci. Rep. 5, 18833. Legeros, R.Z., 2002. Properties of osteoconductive biomaterials: calcium phosphates. Clin. Orthop. Relat. Res. 395, 81 98. Li, X., Cai, W., An, J., Kim, S., Nah, J., Yang, D., et al., 2009a. Large-area synthesis of high-quality and uniform graphene films on copper foils. Science 324, 1312 1314. Li, X., Cai, W., Colombo, L., Ruoff, R.S., 2009b. Evolution of graphene growth on Ni and Cu by carbon isotope labeling. Nano. Lett. 9, 4268 4272. Li, Z., Khun, N.W., Tang, X.Z., Liu, E.J., Khor, K.A., 2016. Mechanical, tribological and biological properties of novel 45S5 Bioglasss composites reinforced with in situ reduced graphene oxide. J. Mech. Behav. Biomed. Mater. 65, 77 89. Liu, J., Cui, L., Losic, D., 2013a. Graphene and graphene oxide as new nanocarriers for drug delivery applications. Acta Biomater. 9, 9243 9257. Liu, Y., Huang, J., Li, H., 2013b. Synthesis of hydroxyapatite reduced graphite oxide nanocomposites for biomedical applications: oriented nucleation and epitaxial growth of hydroxyapatite. J. Mater. Chem. B 1, 1826. Liu, H., Cheng, J., Chen, F., Hou, F., Bai, D., Xi, P., et al., 2014a. Biomimetic and cellmediated mineralization of hydroxyapatite by carrageenan functionalized graphene oxide. ACS Appl. Mater. Interfaces 6, 3132 3140. Liu, Y., Dang, Z., Wang, Y., Huang, J., Li, H., 2014b. Hydroxyapatite/graphene-nanosheet composite coatings deposited by vacuum cold spraying for biomedical applications: inherited nanostructures and enhanced properties. Carbon N.Y. 67, 250 259. Lopez, G., Mittemeijer, E., 2004. The solubility of C in solid Cu. Scr. Mater. 51, 1 5. Lotya, M., Hernandez, Y., King, P.J., Smith, R.J., Nicolosi, V., Karlsson, L.S., et al., 2009. Liquid phase production of graphene by exfoliation of graphite in surfactant/water solutions. J. Am. Chem. Soc. 131, 3611 3620.
612 Advanced Dental Biomaterials Lv, S., Ma, Y., Qiu, C., Sun, T., Liu, J., Zhou, Q., 2013. Effect of graphene oxide nanosheets of microstructure and mechanical properties of cement composites. Construct. Build. Mater. 49, 121 127. Lv, S., Ting, S., Liu, J., Zhou, Q., 2014. Use of graphene oxide nanosheets to regulate the microstructure of hardened cement paste to increase its strength and toughness. CrystEngComm 16, 8508. Maeda, H., Nakano, T., Tomokiyo, A., Fujii, S., Wada, N., Monnouchi, S., et al., 2010. Mineral trioxide aggregate induces bone morphogenetic protein-2 expression and calcification in human periodontal ligament cells. J. Endod. 36, 647 652. Main, C., Mirzayan, N., Shabahang, S., Torabinejad, M., 2004. Repair of root perforations using mineral trioxide aggregate: a long-term study. J. Endod. 30, 80 83. Makar, J., Margeson, J., Luh, J., 2005. Carbon nanotube/cement composites—early results and potential application. In: Proceedings of the Third International Conference on Construction Materials: Performance, Innovations and Structural Implications. August 22 24. Vancouver, B.C., pp. 1 10. Manzur, T., Yazdani, N., 2014. Optimum mix ratio for carbon nanotubes in cement mortar. KSCE J. Civil Eng. 19, 1405 1412. Min, K.S., Yang, S.H., Kim, E.C., 2009. The combined effect of mineral trioxide aggregate and enamel matrix derivative on odontoblastic differentiation in human dental pulp cells. J. Endod. 35, 847 851. Mohandes, F., Salavati-Niasari, M., 2014a. Freeze-drying synthesis, characterization and in vitro bioactivity of chitosan/graphene oxide/hydroxyapatite nanocomposite. RSC Adv. 4, 25993 26001. Mohandes, F., Salavati-Niasari, M., 2014b. In vitro comparative study of pure hydroxyapatite nanorods and novel polyethylene glycol/graphene oxide/hydroxyapatite nanocomposite. J. Nanoparticle Res. 16, 1 12. Mohsen, M.O., Taha, R., Abu Taqa, A., Shaat, A., 2017. Optimum carbon nanotubes’ content for improving flexural and compressive strength of cement paste. Construct. Build. Mater. 150, 395 403. Morin, J.L.P., Dubey, N., Decroix, F.E.D., Luong-Van, E.K., Castro Neto, A.H., Rosa, V., 2017. Graphene transfer to 3-dimensional surfaces: a vacuum-assisted dry transfer method. 2D Mater. 4, 025060. Musso, S., Tulliani, J.-M., Ferro, G., Tagliaferro, A., 2009. Influence of carbon nanotubes structure on the mechanical behavior of cement composites. Compos. Sci. Technol. 69, 1985 1990. Nair, R., Blake, P., Grigorenko, A., Novoselov, K., Booth, T., Stauber, T., et al., 2008. Fine structure constant defines visual transparency of graphene. Science 320, 1308. Nair, M., Nancy, D., Krishnan, A.G., Anjusree, G.S., Vadukumpully, S., Nair, S.V., 2015. Graphene oxide nanoflakes incorporated gelatin-hydroxyapatite scaffolds enhance osteogenic differentiation of human mesenchymal stem cells. Nanotechnology 26, 161001. Natu, V.P., Dubey, N., Loke, G.C., Tan, T.S., Ng, W.H., Yong, C.W., et al., 2015. Bioactivity, physical and chemical properties of MTA mixed with propylene glycol. J. Appl. Oral Sci. 23, 405 411. Nayak, T.R., Andersen, H., Makam, V.S., Khaw, C., Bae, S., Xu, X., et al., 2011. Graphene for controlled and accelerated osteogenic differentiation of human mesenchymal stem cells. ACS Nano 5, 4670 4678. Nishida, E., Miyaji, H., Kato, A., Takita, H., Iwanaga, T., Momose, T., et al., 2016. Graphene oxide scaffold accelerates cellular proliferative response and alveolar bone healing of tooth extraction socket. Int. J. Nanomed. 11, 2265 2277.
Graphene to improve the physicomechanical properties and bioactivity of the cements 613 Ogose, A., Kondo, N., Umezu, H., Hotta, T., Kawashima, H., Tokunaga, K., et al., 2006. Histological assessment in grafts of highly purified beta-tricalcium phosphate (OSferion) in human bones. Biomaterials 27, 1542 1549. Oltra, E., Cox, T.C., Lacourse, M.R., Johnson, J.D., Paranjpe, A., 2017. Retreatability of two endodontic sealers, EndoSequence BC Sealer and AH Plus: a micro-computed tomographic comparison. Restor. Dent. Endod. 42, 19 26. Pan, Z., Duan, W., Li, D., Collins, F., 2012. Graphene Oxide Reinforced Cement and Concrete Patent App. PCT/AU2012/001,582. Parirokh, M., Torabinejad, M., 2010a. Mineral trioxide aggregate: a comprehensive literature review—Part I: Chemical, physical, and antibacterial properties. J. Endod. 36, 16 27. Parirokh, M., Torabinejad, M., 2010b. Mineral trioxide aggregate: a comprehensive literature review—Part III: Clinical applications, drawbacks, and mechanism of action. J. Endod. 36, 400 413. Parirokh, M., Torabinejad, M., Dummer, P.M.H., 2018. Mineral trioxide aggregate and other bioactive endodontic cements: an updated overview—Part I: Vital pulp therapy. Int. Endod. J. 51, 177 205. Perera, S.D., Mariano, R.G., Vu, K., Nour, N., Seitz, O., Chabal, Y., et al., 2012. Hydrothermal synthesis of graphene-TiO2 nanotube composites with enhanced photocatalytic activity. ACS Catal. 2, 949 956. Porter, J.R., Ruckh, T.T., Popat, K.C., 2009. Bone tissue engineering: a review in bone biomimetics and drug delivery strategies. Biotechnol. Prog. 25, 1539 1560. Prati, C., Gandolfi, M.G., 2015. Calcium silicate bioactive cements: biological perspectives and clinical applications. Dent. Mater. 31, 351 370. Qing, Y., Zenan, Z., Deyu, K., Rongshen, C., 2007. Influence of nano-SiO2 addition on properties of hardened cement paste as compared with silica fume. Construct. Build. Mater. 21, 539 545. Rosa, V., Xie, H., Dubey, N., Madanagopal, T.T., Rajan, S.S., Morin, J.L.P., et al., 2016. Graphene oxide-based substrate: physical and surface characterization, cytocompatibility and differentiation potential of dental pulp stem cells. Dent. Mater. 32, 1019 1025. Sheykhrezae, M.S., Meraji, N., Ghanbari, F., Nekoofar, M.H., Bolhari, B., Dummer, P.M.H., 2017. Effect of blood contamination on the compressive strength of three calcium silicate-based cements. Aust. Endod. J. 44, 255 259. Shi, Y.Y., Li, M., Liu, Q., Jia, Z.J., Xu, X.C., Cheng, Y., et al., 2016. Electrophoretic deposition of graphene oxide reinforced chitosan-hydroxyapatite nanocomposite coatings on Ti substrate. J. Mater. Sci. Mater. Med. 27, 48. Siddique, R., Mehta, A., 2014. Effect of carbon nanotubes on properties of cement mortars. Construct. Build. Mater. 50, 116 129. Stankovich, S., Dikin, D.A., Piner, R.D., Kohlhaas, K.A., Kleinhammes, A., Jia, Y., et al., 2007. Synthesis of graphene-based nanosheets via chemical reduction of exfoliated graphite oxide. Carbon N. Y. 45, 1558 1565. Tatavarty, R., Ding, H., Lu, G., Taylor, R.J., Bi, X., 2014. Synergistic acceleration in the osteogenesis of human mesenchymal stem cells by graphene oxide-calcium phosphate nanocomposites. Chem Commun (Camb) 50, 8484 8487. Tomas-Catala, C.J., Collado-Gonzalez, M., Garcia-Bernal, D., Onate-Sanchez, R.E., Forner, L., Llena, C., et al., 2017. Comparative analysis of the biological effects of the endodontic bioactive cements MTA-Angelus, MTA Repair HP and NeoMTA Plus on human dental pulp stem cells. Int. Endod. J. 50 (Suppl. 2), e63 e72. Tomas-Catala, C.J., Collado-Gonzalez, M., Garcia-Bernal, D., Onate-Sanchez, R.E., Forner, L., Llena, C., et al., 2018. Biocompatibility of new pulp-capping materials NeoMTA
614 Advanced Dental Biomaterials plus, MTA repair HP, and biodentine on human dental pulp stem cells. J. Endod. 44, 126 132. Tonetto, A., Lago, P.W., Borba, M., Rosa, V., 2015a. Effects of chrondro-osseous regenerative compound associated with local treatments in the regeneration of bone defects around implants: an in vivo study. Clin. Oral. Investig. 20, 1 8. Tonetto, A., Lago, P.W., Borba, M., Rosa, V., 2015b. Effects of chrondro-osseous regenerative compound associated with local treatments in the regeneration of bone defects around implants: an in vivo study. Clin. Oral. Investig. 20, 267 274. Torabinejad, M., Parirokh, M., Dummer, P.M.H., 2018. Mineral trioxide aggregate and other bioactive endodontic cements: an updated overview—Part II: Other clinical applications and complications. Int. Endod. J. 51, 284 317. Wu, C., Xia, L., Han, P., Xu, M., Fang, B., Wang, J., et al., 2015. Graphene-oxide-modified β-tricalcium phosphate bioceramics stimulate in vitro and in vivo osteogenesis. Carbon N. Y. 93, 116 129. Xie, H., Cao, T., Gomes, J.V., Castro Neto, A.H., Rosa, V., 2015a. Two and threedimensional graphene substrates to magnify osteogenic differentiation of periodontal ligament stem cells. Carbon N.Y. 93, 266 275. Xie, H., Cao, T., Gomes, J.V., Neto, A.H.C., Rosa, V., 2015b. Two and three-dimensional graphene substrates to magnify osteogenic differentiation of periodontal ligament stem cells. Carbon N.Y. 93, 266 275. Xie, H., Chua, M., Islam, I., Bentini, R., Cao, T., Viana-Gomes, J.C., et al., 2017. CVDgrown monolayer graphene induces osteogenic but not odontoblastic differentiation of dental pulp stem cells. Dent. Mater. 33, e13 e21. Yassen, G.H., Huang, R., Al-Zain, A., Yoshida, T., Gregory, R.L., Platt, J.A., 2016. Evaluation of selected properties of a new root repair cement containing surface prereacted glass ionomer fillers. Clin. Oral. Investig. 20, 2139 2148. Yazdani, N., Mohanam, V., 2016. Strength, bleeding and setting time of cement mortar with carbon nanotubes and nanofibers. Adv. Sci. Eng. Med. 8, 490 495. Yi, L., Jing, H., Hua, L., 2013. Synthesis of hydroxyapatite reduced graphite oxide nanocomposites for biomedical applications: oriented nucleation and epitaxial growth of hydroxyapatite. J. Mater. Chem. B 1, 1826 1834. Zanin, H., Saito, E., Marciano, F.R., Ceragioli, H.J., Campos Granato, A.E., Porcionatto, M., et al., 2013. Fast preparation of nano-hydroxyapatite/superhydrophilic reduced graphene oxide composites for bioactive applications. J. Mater. Chem. B 1, 4947. Zhang, L., Liu, W., Yue, C., Zhang, T., Li, P., Xing, Z., et al., 2013a. A tough graphene nanosheet/hydroxyapatite composite with improved in vitro biocompatibility. Carbon N. Y. 61, 105 115. Zhang, L., Zhang, F., Yang, X., Long, G., Wu, Y., Zhang, T., et al., 2013b. Porous 3D graphene-based bulk materials with exceptional high surface area and excellent conductivity for supercapacitors. Sci. Rep. 3, 1378. Zheng, Q.,, Kim, J.-K.,, 2015. Synthesis, Structure, and Properties of Graphene and Graphene Oxide. Springer, New York, pp. 29 94. Zhou, M., Wang, Y., Zhai, Y., Zhai, J., Ren, W., Wang, F., et al., 2009. Controlled synthesis of large-area and patterned electrochemically reduced graphene oxide films. Chemistry 15, 6116 6120.
Biomaterials for maxillofacial prosthetic rehabilitation 23 Waqas Tanveer Maxillofacial Prosthetic Service, Department of Prosthodontics, Faculty of Dentistry, Mahidol University, Bangkok, Thailand Chapter Outline 23.1 23.2 23.3 23.4 Highlights 616 Historical background 616 Ideal properties of maxillofacial material 617 Search for ideal materials for maxillofacial rehabilitation 617 23.4.1 23.4.2 23.4.3 23.4.4 23.4.5 23.4.6 Acrylic resins (1940 60) 617 Polyvinylchloride and copolymer 619 Chlorinated polyethylene 619 Polyurethane elastomers (1970 90) 619 Thermoset urethane elastomers 619 Silicones (1960 70) 620 23.5 Silicones 620 23.5.1 Polymer structures 620 23.6 Classification of maxillofacial silicones 622 23.6.1 Classification of silicones according to application 622 23.7 Types of maxillofacial silicones 23.7.1 23.7.2 23.7.3 23.7.4 625 Most common room temperature vulcanizing silicones 625 Medical grade liquid silicone elastomers 627 Recommendations 627 Medical grade VerSiTal silicone elastomers 628 23.8 M-511 platinum silicone rubber 23.8.1 Silicone fluids 629 23.8.2 Properties 629 23.8.3 Types of silicone fluids 628 629 23.9 Primers 632 23.10 Soft liners and tissue conditioners 633 23.10.1 Soft liner 633 23.10.2 Coe-Comfort and Coe-Soft 633 23.10.3 Sculpturing clays and waxes 635 23.11 Coloring agents 23.11.1 23.11.2 23.11.3 23.11.4 635 Colored flocking 635 Intrinsic stains 636 Extrinsic colors 636 Acetoxy silastic adhesives 636 23.12 Skin adhesives References 639 638 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00024-4 Copyright © 2019 Elsevier Ltd. All rights reserved.
616 23.1 G G G G G G G G G Advanced Dental Biomaterials Highlights Introduction History of maxillofacial prosthetics Ideal requirements of maxillofacial materials Silicones Primers Soft liners and tissue conditioners Waxes Coloration Adhesives According to the Glossary of Prosthodontics Treatment (GPT), maxillofacial prosthodontics is defined as “The branch of prosthodontics concerned with the restoration and replacement of stomatognathic and craniofacial structures with prostheses that may or may not be removed on a regular or elective basis” (GPT) (Keith and Steven, 2017). Maxillofacial prosthodontics is a subgroup of prosthodontics, which is concerned with the rehabilitation of congenital and acquired defects of the craniofacial region with prostheses, which intends to restore the defects for optimal esthetics and function. They are divided into two major categories: intraoral prostheses and extraoral prostheses. The intraoral prostheses include surgical obturators, interim and definitive obturators, palatal lift and speech aid prostheses, palatal drop prostheses, mandibular guidance prostheses, tongue, jaw prostheses, and conventional dentures which tend to restore esthetics, speech, and functions. While in extraoral prostheses, oculars, auricular, orbital, nasal, mid-facial, and somatic prostheses (replaces a body part like fingers and hand) are included. There are various materials, which can be used to fabricate prostheses to obturate the maxillofacial defects. 23.2 Historical background The prosthetic rehabilitation of craniofacial defects has not been well documented in historical literature. There is some evidence that early Chinese and Egyptians had used wood, metals, and resins to restore the maxillofacial defects; however, the refined techniques for restoring maxillofacial defects were introduced after World War II with introduction of acrylic resins. The introduction of the osseointegration concept and improvement in physical and mechanical properties of silicones have brought revolution in the field of maxillofacial rehabilitation due to the predictable and long-lasting, esthetic, and functional outcome. The historical background and gradual advancement in maxillofacial prosthetic materials is briefly summarized in Table 23.1.
Biomaterials for maxillofacial prosthetic rehabilitation 617 Table 23.1 Historical background of maxillofacial prosthetic materials. Period Investigator Development Before 1600 Ambroise Pare Prosthetic nose made from silver, artificial ear made from paper and leather Artificial nose made from gold Facial prosthesis with paper and silver Nasal defect restoration by using enameled porcelain Combined nasal and palatal prosthesis Nasal prosthesis fabricated with ceramic Nasal and auricular prosthesis made with vulcanite rubber Nasal prosthesis made of vulcanized rubber Facial prosthesis made with gelatin-glycerin Vulcanized latex and water-soluble dyes 1600 1800 1800 1900 1900 40 1940 60 1960 70 1970 90 1990 to present 23.3 Tycho Brahe Pierre Fauchard William Morton Kingsley Claude Martin Upham Ottofy, Baird, and Baker Bercowitsch Kazanjian, Bulbulian, Clarke US Navy Barnhart Lontz Gonzalez Turner Udagama and Drane Antonucci and Stansbury Gettleman Artificial eyes made from acrylic resin Silicone material with acrylic resin Modified polysiloxane elastomers Polyurethane elastomers Isophorone polyurethane Silastic medical grade adhesives New acrylic resins Polyphosphazenes Ideal properties of maxillofacial material The successful completion of maxillofacial prosthesis depends on multiple factors such as the physical, mechanical, and biological properties of materials as well as processing conditions (Beumer et al., 2011). The required properties of maxillofacial prosthetic materials are listed in Tables 23.2 and 23.3. 23.4 Search for ideal materials for maxillofacial rehabilitation 23.4.1 Acrylic resins (1940 60) Acrylic resin has been used in cases when the least movement of the tissue bed is expected during function. It has various advantages including color stability, easy to process, can be relined and repaired, has good strength, can be fabricated with feather margin, and a good shelf-life of about 2 years. However, this material lacks flexibility or the ability to absorb water or adaptability to adjacent soft tissues (Khindria et al., 2009; Udita et al., 2010). Acrylic copolymers with elastic properties have disadvantages like poor edge strength and durability when exposed to
618 Advanced Dental Biomaterials Table 23.2 Ideal properties of maxillofacial materials. Ideal properties of maxillofacial prosthetic materials Processing properties Biological properties Physical and mechanical properties Dimensionally stable Should not affect supporting tissues Should be nonallergic to tissues Nontoxic and nonirritant Should be chemically inert Easy to stain Easy to manipulate Able to reproduce fine details Have stable composition No access to microbial colonization Odorless Should have enough working time Stable properties at higher temperature Long-term color stability Nontoxic while handling Should be easy to clean Colorants should be stable within materials Able to tolerate cyclic loading Able to maintain the surface hardness Resistant to abrasion Low specific gravity Low surface tension upon curing High tensile strength to resist tearing Optimum elastic properties Table 23.3 Required mechanical and physical properties of maxillofacial prosthetic materials (Lewis and Castleberry, 1980). Desirable physical and mechanical characteristics Values Tear strength Tensile strength Elastic modulus Elongation percent Glass transition temperature Heat distortion temperature Critical surface tension Coefficient of friction Hardness Water absorption 30 100 psi 1000 2000 psi 50 250 psi 400% 800% ,0 C .120 C 30 45 dyn/cm 0.4 0.6 25 35 Shore A scale None sunlight. The prosthesis often turns tacky and prone to dust collection and stains. The new generation of acrylic monomers, macromeres, and oligomers are chemical, thermal, and UV initiated; therefore, they can overcome the shortcomings of conventional acrylic copolymers (Udita et al., 2010).
Biomaterials for maxillofacial prosthetic rehabilitation 619 23.4.2 Polyvinylchloride and copolymer These copolymers are clear and odorless materials and have been used commonly for maxillofacial prosthetic applications with the advantages of being adaptable, flexible, suitable for intrinsic, and extrinsic staining. Plasticizers are inserted to incorporate elastomeric effects at room temperature. Other constituents include ultraviolet stabilizers for color stability and cross-linking agents for additional strength. 23.4.3 Chlorinated polyethylene The processing of chlorinated polyethylene involves high heat curing pigmented sheets in metal molds. Dow Chemicals’ chlorinated elastomer is an industrial grade thermoplastic elastomer. It is biocompatible, noncarcinogenic, and less toxic than thermosetting silicone materials. Chlorinated polyethylene elastomer can be a suitable alternative to silicones for extraoral maxillofacial prosthesis when cost of prosthesis is a limiting factor. 23.4.4 Polyurethane elastomers (1970 90) This material is composed of aliphatic di-isocyanate groups, polyol groups (mixture of polyesters), and organotin catalyst for completion of the polymerization process. The proportion of these groups dictate the Shore A hardness of these polyurethane elastomers; therefore, their intended application varies accordingly, since the maxillofacial prosthetic materials require appropriate flexibility and adaptability with adjacent structures. Epithane-3 and Calthane are the only polyurethanes currently available for fabricating facial prostheses. They comprise three components, roomtemperature vulcanizing materials (Huber and Stephan, 2002). These materials can be used to rehabilitate soft tissue defects, as they have the inherent flexibility needed to adapt with tissues for stability and esthetics. These possess environmental stability, low modulus, require no addition of plasticizers, and have higher tear strength, optimum ultimate strength, and elongation. Intrinsic coloration is possible with these materials; therefore, they are amenable to maxillofacial processing techniques. The disadvantages include lack of compatibility with adhesive systems, low color stability, and moisture sensitivity resulting in gas bubbles during processing, due to the presence of di-isocyanate. Hence, the reaction should be carried out in dry atmosphere, as carbon dioxide would be produced, resulting in porous elastomer (Robert and John, 2002). 23.4.5 Thermoset urethane elastomers Thermoset urethane elastomers are produced through the introduction of primary chemical cross-linkers. If reactants are combined in stoichiometric ratios and reactions are preferentially catalyzed, a known controlled morphology can be developed. According to Yu et al. (1980), the aging process affects polyurethane significantly.
620 Advanced Dental Biomaterials 23.4.6 Silicones (1960 70) Silicones are known as polydimethylsiloxane, the most successful maxillofacial prosthetic material to date. The silicones are popular materials due to their better physical and mechanical properties over other materials (excellent tear, tensile, and peel strength). Furthermore, they can be stained intrinsically and extrinsically to provide more esthetic and lifelike appearance. When appropriately vulcanized, silicones elastomers can resist deposition of organic materials that may lead to bacterial colonization, therefore, with simple cleaning methods, these materials are relatively safe and hygienic compared to other materials (Sharif et al., 2010). 23.5 Silicones In general silicones exist in five different structures (Fig. 23.1). By changing the structure of silicone, many different types of silicones can be produced. 23.5.1 Polymer structures (Deanin, 1972) 1. Linear polymers: Volatile and nonvolatile fluids (dimethicone). 2. Ring polymers: Volatile fluids (cyclomethicone). 3. Branched polymers: The ends of branches provide additional opportunities to connect with other chains of polymers. 4. Cross-linked polymers: Gels, elastomers, sealants, rubber. 5. Resins: Structures that create three-dimensional films upon curing. This type of silicone has industrial and commercial application. The silicones are synthetic materials, which consist of long chains of polymer units. Silicones have different chemical and physical properties that make them more useful than other polymers, and they keep these desirable properties under different environmental conditions (Dow Corning Corporation). Unlike organic polymers, silicone’s structure is made up of alternating chains of oxygen and silicon atoms. Mostly organic or methyl groups are attached to sides of silicone atoms (Fig. 23.2). Figure 23.1 Basic polymer structures (Deanin, 1972).
Biomaterials for maxillofacial prosthetic rehabilitation 621 Figure 23.2 Basic silicone unit (Colas and Curtis, 2005). Figure 23.3 Basic steps in silicone polymer synthesis (Colas and Curtis, 2005). By modifying the length of this oxygen silicone chain, scientists can produce different forms of silicones, such as fluids, resins, or viscous elastomers. These longchain synthetic silicones are used in various products, including waxes, repellants, polishes, nonstick coatings, and electrical insulation (Dow Corning Corporation). Silicone does not exist naturally; therefore, it is completely synthetic in nature. It is often bound with oxygen in materials like sand, granite, and quartzite. The highgrade quartz is reduced to elemental silicone by reacting in an electric furnace in a process called carbothermic smelting. The next step in the procedure is producing methylchloride by condensing methanol with hydrochloric acid (Dow Corning Corporation). Generally, two methyl groups (CH3) attach to each silicone atom to comprise each molecule of dimethyldichlorosilane, the principal unit of silicone. Methylchloride is then allowed to flow through the fluid form of silicone powder; this reaction is accelerated by a copper-based catalyst and produces complex units of methylchlorosilanes. The multiple methylchlorosilane groups are separated by distillation reaction and then condensed to dimethyldichlorosilane. When water is added to dimethyldichlorosilane (principal chlorosilane), the two (water and dimethylchlorosilane) react to disilanols and hydrochloric acid. This reaction occurs as oxygen has more affinity for silicone while chlorine has more affinity for hydrogen (Fig. 23.3). The final step is catalyzed by hydrochloric acid, which further
622 Advanced Dental Biomaterials Table 23.4 Composition of silicone elastomer (Colas and Curtis, 2005). Contents Examples Silicone polymer Fillers Catalyst/vulcanizing agent Polydimethylsiloxane Silica Platinum salt Dichlorobenzyl peroxide Stannous octoate Barium, tin Hydromethylsiloxane Opacifiers Cross-linking agent condenses disilanols into polydimethylsiloxanes and forms the backbone of silicone atoms bounded by oxygen atoms (Dow Corning Corporation). Dimethyldichlorosiloxane, which takes part in a catalyzed reaction, is the most frequently used material for maxillofacial prosthetic rehabilitation. It forms a polymer upon reaction with water. These polymers appear as translucent, white fluids and are watery in consistency. The extent of viscosity can be assessed by the length of the polymer chain (Dow Corning Corporation). Most rubbery consistency silicones are manipulated with fillers, which provide the required strength. Furthermore, additives provide different colors to translucent silicones. Vulcanizing agents as well as antioxidants are used to convert the raw product from plastic to a rubbery consistency, during processing. The long chain of polymer creates a network, which is difficult to disengage. This strong network makes the silicone resistant to adverse changes upon exposure to ultraviolet radiations. The general composition of silicone elastomer is shown in Table 23.4. 23.6 Classification of maxillofacial silicones 23.6.1 Classification of silicones according to application (Beumer et al., 2011) 1. 2. 3. 4. Implant grade Medical grade Food grade Industrial grade 23.6.1.1 Implant grade These materials need extensive testing and should fulfill FDA requirements to be used for medical purposes. According to literature, silicone breast implants have a controversial outcome regarding the success and safety of interstitial implants placement; therefore, these materials need further research.
Biomaterials for maxillofacial prosthetic rehabilitation 623 23.6.1.2 Medical grade These materials were approved for external use only. Their use in health care is widespread, mainly due to their biocompatibility and biodurability. Silicone use for maxillofacial prosthetics started in 1943, but the medical use started in 1953. The first silicone was investigated and developed by Dow Corning; they further developed a separate center to study medical applications of silicone materials in 1963. Later, silicones because popular and the material of choice for facial prostheses due to the physical properties similar to the skin texture (Fig. 23.4). Adverse reactions can be caused by contact of silicone prostheses with skin, but these reactions are very rare. 23.6.1.3 Food grade These materials meet FDA rules and regulations for food content. Many manufacturers claim that food grade silicones are safe to be used as maxillofacial prosthesis. VST-50 silicone material is a food grade silicone manufactured by Factor II Inc., the United States. Food grade silicones are more economical than medical grade silicones and might be used as an alternative, but their properties need to be evaluated before regular use in the fabrication of maxillofacial prosthesis 23.6.1.4 Industrial grade These materials have industrial application and their use in maxillofacial prosthetics is not documented. Figure 23.4 Medical grade silicones.
624 Advanced Dental Biomaterials 23.6.1.5 Maxillofacial silicones Maxillofacial silicones can be divided into two major groups according to their vulcanization reactions (Beumer et al., 2011; Colas and Curtis, 2005) (Table 23.5). 1. Room-temperature vulcanization (RTV) silicone 2. High-temperature vulcanization (HTV) silicone The vulcanization reaction of polymers includes the bonding of separate polymer chains. The ties among the long chains of polymer units are called cross-linking of polymers. The process through which cross-linking of polymer chains takes place is called vulcanization. The reaction can occur either with or without the use of heat, depending on the catalytic and/or cross-linking agents. The silicone elastomers used for maxillofacial prosthetics (medical grade) contain fillers and specific vulcanizing agents; however, they lack the huge variety of additives used in the organic rubber industry. 23.6.1.6 Room temperature vulcanizing silicone elastomer There are two types of RTV silicones (Mitra et al., 2014): 1. Cross-linking by condensation reaction 2. Cross-linking by addition reaction Cross-linking by condensation reaction (Mitra et al., 2014) They contain active groups such as silanols (hydroxyl-terminated polysiloxanes). Cross-linking by this method requires a cross-linking agent, for example, tetraethyl silicate, and a catalyst, for example, dibutyltin dilaurate, for example, Silastic Medical Adhesive type A (Dow Corning), where methyl triacetoxy silane (II) will be used as the cross-linking agent. However, cross-linking requires water molecules to hydrolyze the silane and produces acetic acid (an irritant) as a by-product. These types of silicones have limited use, mostly used to fix extrinsic staining. Disadvantages (Mitra et al., 2014) 1. Releases by-products 2. Vulcanization time is excessively long 3. Includes degradation reaction such as hydrolysis Table 23.5 Properties of the two types of silicone materials (Mitra et al., 2014). S. no. Properties HTV RTV 1. 2. 3. Ultimate tensile strength (MPa) Maximum elongation (%) Pants tear strength (dyn/ cm 3 106) 4. Dynamic modulus (MPa) 5.86 441 Does not tear but stretches, as in tensile elongation 4.66 4.20 445 Does not tear but stretches, as in tensile elongation 2.12 HTV, High temperature vulcanization; RTV, room temperature vulcanization.
Biomaterials for maxillofacial prosthetic rehabilitation 625 4. Relatively lower tear strength 5. Unable to maintain edge resistance Cross-linking by addition reaction (Mitra et al., 2014) The reactions usually involve the addition of silyl hydride groups ( SiH) to vinyl groups (CH2QCH ) attached with the silicone by the help of a platinum catalyst. These silicones are not strictly room-vulcanized silicones. The curing of these silicones requires exposure of dry heat at 150 C for an hour. These silicone elastomers have enhanced tear strength as compared to the first type of RTV silicones. Disadvantages (Mitra et al., 2014) 1. Very hydrophobic 2. No extrinsic coloration 3. Curing of the materials may be inhibited by: amines, sulfur, nitrogen oxides 23.6.1.7 High-temperature vulcanizing silicone elastomer The synthesis of heat-vulcanized silicone includes use of a di-organo-polysiloxane such as poly (dimethyl siloxane). When the liquid form of silicone is heated with benzoyl peroxide activator, a reaction results between one of the methyl radicals in the chain and an identical group in the adjacent chain. It results in cross-linking between two polymeric chains occur and ultimately yields benzoic acid as a byproduct. HTV silicone has been manufactured in three different consistencies, soft, medium, and hard. It is usually colorless or white in color with opaque appearance and varying degree of consistency (Dow Corning Corporation; Colas and Curtis, 2005). Advantages (Mitra et al., 2014): 1. 2. 3. 4. Adequate tear strength High tensile strength (5.87 MPa) Excellent thermal, color, and chemical stability Adequate percentage of elongation Disadvantages (Mitra et al., 2014): 1. 2. 3. 4. Poor esthetics due to opacity Less elasticity Low edge strength Technique sensitive 23.7 Types of maxillofacial silicones 23.7.1 Most common room temperature vulcanizing silicones Many materials have been tried to replace the lost hard and soft tissues of the body, including wood, stones, metals, and ceramics. After World War II, acrylic was introduced to rehabilitate the ocular defects. The greatest revolution in maxillofacial
626 Advanced Dental Biomaterials prosthetics rehabilitation came with the introduction of maxillofacial silicones. These maxillofacial silicones were considered superior to conventionally used acrylic resins for facial prostheses due to their flexibility, biocompatibility, ease of processing, and ability to adapt with adjacent soft tissues providing a lifelike appearance. Maxillofacial silicone MDX4-4210 is a translucent, two-component silicone (10:1 base to catalyst ratio by weight). It was introduced in 1970 and has been the most common silicone for facial prostheses. Silastic MDX4-4210 biomedical grade elastomer is a room temperature vulcanizing silicone, supplied as pourable twocomponent silicone, which vulcanizes upon mixing with catalyst, resulting in a translucent silicone elastomer. The vulcanization of silicone can be accelerated by a dry heat oven without significantly affecting the physical properties. The maxillofacial silicone A-2186 (Fig. 23.5) was introduced in 1986 for the fabrication of facial prosthesis. This silicone has been claimed to have comparable properties as MDX4-4210. It is a translucent silicone, composed of two parts silicone (10:1 ratio by weight for base to catalyst). The rapid vulcanization form of A2186 was introduced in 1987 with more platinum content to accelerate the curing of silicone. According to a study by Sanchez, A-2186 had greater tensile strength, tear resistance, and a higher percentage of elongation. A-2186 silicone also proved to be softer at the surface than the MDX 4-4210. The combination of these physical properties makes A-2186 silicone a better choice than the more conventionally used MDX 4-4210 for the fabrication of maxillofacial prostheses. The properties of the two most common maxillofacial silicones, MDX4-4210, and A-2186 are shown in Table 23.6. Figure 23.5 Maxillofacial silicone A-2186.
Biomaterials for maxillofacial prosthetic rehabilitation 627 Table 23.6 Comparative properties of two most common maxillofacial silicones (Platinum silicone A-2186; MDX4-4210). Properties Silicone A-2186 MDX4-4210 Color Consistency Viscosity Specific gravity Working time Durometer Shore A Tensile (psi) Elongation (%) Translucent Pourable 90 Pa s 1.10 2h 29 900 650 Translucent Pourable 60 1.11 3 30 730 470 Table 23.7 Properties of different types of liquid silicone rubber silicone elastomers. Silicones Durometer (Shore A) Tensile strength (psi) Elongation (%) Modulus 100% (psi) Tear strength (ppi) Appearance A-221-05 A-223-30 A-223-40 A-225-50 A-225-60 A-225-70 5 31 41 50 59 68 450 1400 1230 1225 1300 1300 980 790 615 570 480 450 15 90 190 295 350 530 60 195 255 265 250 240 Clear, thixotropic Clear, paste-like Clear, paste-like Clear, paste-like Clear, paste-like Clear, paste-like 23.7.2 Medical grade liquid silicone elastomers Liquid silicone rubber (LSR) is a colorless, translucent clear paste. This silicone is composed of part-A and part-B that are mixed together in equal proportions, 1:1 ratio by weight. The LSR platinum-catalyzed addition-cured silicone appears optically clear and is available in different types depending on the mechanical properties (Table 23.7). Conventionally, silicone A-221-05 is used for breast and delicate soft tissues prostheses due to very low Shore A hardness and modulus along with high elongation properties, which makes it highly soft and flexible. Silicones A-223-30 and A-223-40 are recommended for facial prostheses and somatic prostheses (fingers, hand, and foot prostheses), as the facial prosthesis needs to have Shore A hardness in the range of 30 50. Silicones with higher Shore A hardness, modulus, and low elongation properties are generally used to replace the hard tissues and molds fabrication. 23.7.3 Recommendations 05—Facial—breast prosthetics 30 40—Facial, finger, and hand prosthetics 50 70—Molds fabrication and restoration of hard tissue defects
628 Advanced Dental Biomaterials 23.7.4 Medical grade VerSiTal silicone elastomers The VST groups of silicones are translucent, composed of part-A and part-B, that are mixed in the ratio of 10:1 by weight (Fig. 23.6). These silicones have low viscosity; however, the viscosity can be increased by using thixotropic agents, depending upon the application. The VST groups of silicones are room temperature vulcanizing, but the vulcanization reaction can be accelerated by exposure to high temperature. These platinum-based addition-cured silicones are commonly used to fabricate silicone facial prostheses. The VST silicones are available in different types according to their vulcanization time and Shore A hardness. (Table 23.8) The silicone VST-30 has a short working and curing time; therefore, it requires expertise while mixing intrinsic stains according to patient’s skin tone. The curing times mentioned in Table. 23.9 are according to RTV. 23.8 M-511 platinum silicone rubber M-511 was specially formulated for the fabrication of maxillofacial prostheses. It is the silicone elastomer of choice at many clinics and hospitals throughout England. This silicone is ideal for implant-retained facial prostheses, as it has been claimed to bond well with methyl methacrylate (acrylic) resin. The recommended base to catalyst mixing ratio is 10:1, by weight. Upon vulcanization this silicone attains 25 30 Shore A hardness, which is considered appropriate to be used for the fabrication of facial silicone prosthesis. The M-511 silicone elastomer has 1 hour of working time. Silicone M-511 can be used with following components to modify its properties: G G G M-510 gel catalyst to fabricate gel-filled prosthesis M-513 softening agent to modify Shore A hardness (Shore A 15 20) M-514 antislump agent (thixotropic agent) Figure 23.6 VST silicone elastomers for maxillofacial prosthetics.
Biomaterials for maxillofacial prosthetic rehabilitation 629 Table 23.8 Properties of different types of VST silicone materials. Properties VST-30 VST-50 VST-50F VST-50HD Translucent Shore A Cure time RTV Pot life Tear strength (ppi) Tensile strength (psi) Elongation (%) Viscosity A Viscosity B Yes 23 25 30 10 13 min 115 985 610 12,000 10,000 Yes 30 Overnight 90 min 130 1035 660 12,000 10,000 Yes 28 2 3h 30 min 120 1095 625 12,000 10,000 Yes 38 Overnight 2h 110 950 550 15,000 10,000 Table 23.9 Shore A hardness of VST silicones. Product Cure Shore A hardness VST-30 VST-50 VST-50F VST-50HD 25 30 min Overnight 2 3h Overnight 23 30 28 38 Note: The base of all VST silicone materials is similar. They vary according to their crosslinking agents. 23.8.1 Silicone fluids These silicone fluids are supplied as thinning agents for all silicone elastomers and are composed of straight-chained 100% polydimethylsiloxane. These agents have been recommended to be used at no more than 20% by weight of silicone, during silicone mixing to improve the working time (Table 23.10). 23.8.2 Properties G G G G G G G Optimum resistance to low and high temperatures Good dielectric properties Low surface tension High compressibility Absence of aging upon exposure to atmospheric changes Better oxidation resistance Minimal change in viscosity with temperature 23.8.3 Types of silicone fluids G G G A-313 functional silicone fluid (platinum compatible) A-318 OH thinner A-244 silicone fluid
630 Advanced Dental Biomaterials Table 23.10 Research studies about maxillofacial silicones. S. no. Author (year) Study conducted Materials used Conclusion 1. Andreopoulos and Evangelatou (1994) Silica fiber fillers (short aramid fibers, glass fibers, high modulus polyethylene fibers) Instead of ultrahighmodulus fibers, silica fillers should be used 2. Lai (1996) Effect of fillers on the mechanical properties (tensile strength and modulus) of poly (dimethyl siloxane) rubber (C-50, Bayer AG, Leverkusen, Germany) maxillofacial material Invention of polysiloxane materials including MPDS to reduce the hydrophobicity of silicones Modification of polydimethyl-siloxanes. Thermal initiator— lauroyl peroxide preferred over benzoyl peroxide. Filler— amorphous silica. Cross-linker— TEGDMA. Thermal curing of the polysiloxane MPDS was conducted in a stainless steel mold to minimize the effects of oxygen inhibition MPDS has following advantages over MDX44210: Very long working time Extended shelf-life Higher tear and tensile strength Incorporation of methacrylate group in the silicones reduces the hydrophobicity of the silicones enhances the adhesive bond strength to the nonsilicone-based adhesives. However, the ultimate elongation is slightly lower than that of MDX4-4210 Tensile strength improved with increasing silica fillers up to 35% significantly 3. Andreopoulos et al. (1998) 4. Lai et al. (2002) 5. Aziz et al. (2003b) Effect of reinforcement of silica powder up to 50% concentration on the mechanical properties of silicone maxillofacial materials Introduction of new organosilicone MPDSMF 606 Silica powder up to 50% concentration Effect of different parameters on the development of new improved maxillofacial material C-50 Silica fillers—R104, R106, R202, R972, R974, and R812. Silica filler concentrations— 0 25 wt.%. Crosslinker concentration— 0.12 0.72 g (per 10 g of base polymer) Ratio of high and low molecular weight polymers Modified MPDS and comparing the properties with A-2186 G G G The hardness is similar to A-2186, whereas tear, tensile strength, ultimate elongation, and adhesive bond to nonsilicones are higher than A-2186. Suitable for use in fabrication of clinical prostheses Increasing the R812 silica filer concentration from 15 to 20 wt.% and having cross-linker of 0.28% was associated with significant increase in tear strength (Continued)
Biomaterials for maxillofacial prosthetic rehabilitation 631 Table 23.10 (Continued) S. no. Author (year) Study conducted Materials used Conclusion 6. Karayazgan et al. (2003) Effect of tulle in increasing the tear strength of maxillofacial prosthesis at the edges Tulle 7. Aziz et al. (2003a) Analysis of the properties of commonly used silicone rubber materials Cosmesil HC and Nusil (MED-4920), Prestige (Premium facial and body elastomer), Factor II (A-2186) 8. Gunay et al. (2008) Tulle 9. Mirabedini et al. (2008) Investigation on the incorporation of tulle in silicone maxillofacial prosthesis Effect of TiO2 on the mechanical and adhesion properties of RTV silicone elastomer coatings Tulle resulted in having a more stable and tearresistant silicone prosthesis, which is less likely to deform during application or removal of adhesives, cosmetics, or cleaning agents None of the commercially available materials have adequate properties. Cosmesil HC and Nusil (MED-4920) had adequate mechanical properties along with good translucency but have high viscosity. Prestige (premium facial and body elastomer) was a soft material; however, together with Cosmesil St, it had extremely high water absorption and low tears strength Factor II (A-2186) had well all round mechanical properties and had the highest tear strength Tensile and tear strengths were significantly higher 10. Sharif et al. (2010) Effect of POSS on the mechanical properties, of maxillofacial material (Factor II A2000) TiO2 fillers—5.0, 10.0, 15.0, and 20.0 wt.% POSS Tensile stress, modulus, and abrasion resistance increased with the increase in filler content up to 15.9 wt.%. Hardness was increased with increasing filler contents up to 25.0 wt. %, but on the contrary, elongation to fracture was decreased with increasing filler content The POSS loading had a significant effect on the tear and tensile properties of the materials (Continued)
632 Advanced Dental Biomaterials Table 23.10 (Continued) S. no. Author (year) Study conducted Materials used Conclusion 11. Liu et al. (2013) Effect of silicone elastomers filled with hollow microspheres Hollow microspheres were mixed with MDX44210 silicone elastomer using Q7-9180 silicone fluid as diluents. Volume fractions of microspheres were 0%, 5%, 15%, and 30% v/v (volume ratio to the total volume of MDX44210 and microspheres) were used The microspheres dispersed well in the matrix. Shock absorption was the greatest by the 5% v/v composite and decreased with increasing concentrations of microspheres. Other properties except elongation at break decreased with increasing concentration. Tear strength of all composites was markedly lower than that of pure silicone elastomer. Has good biocompatibility. Novel light and soft material with good flexibility, biocompatibility, and properties at 5% holds a promising prospect for clinical application as maxillofacial prosthesis MPDS, Methacryloxyalkyl-terminated polydimethylsiloxanes; POSS, polyhedral silsesquioxane; RTV, room temperature vulcanization. 23.9 Primers The primers used for maxillofacial silicones are usually platinum-cured, composed of one component. These are designed to improve the adhesion of addition-cured silicones to various substrates, including metals, glass, ceramics, some plastics, and other silicone materials. The platinum primers are transparent liquids with viscosity of 1.0 cps and specific gravity 0.77 (Table 23.11). The platinum primers contain VM&P Naphtha as the main solvent. A uniform thin film is required for best bonding results. They should be applied after thorough cleaning and degreasing of the surface of substrate with a chlorinated solvent 1:1:1 trichloroethane acetone or MEK. After the substrate dries, a uniform thin coat of primer is applied at room temperature with 50% relative humidity. The primer coat should be allowed to dry for a minimum of 30 minutes. The primer sets by exposure to moisture and higher humidity. The heavy whitish haze or chalky appearance is indicative of a thick layer of primer, which can adversely affect the bond strength; therefore, the primer should be cleaned and reapplied. The most common primers used for maxillofacial silicones are A-330G, A-304, A-306, and G-611 (A-306; A-330 G).
Biomaterials for maxillofacial prosthetic rehabilitation 633 Table 23.11 Properties of platinum primers. 23.10 Properties Platinum primers Color Viscosity (cps) Flash point Solvent Specific gravity Nonvolatile content Drying time at 77 F Translucent 1.0 48 F VM&P Naphtha 0.77 5 30 min Soft liners and tissue conditioners 23.10.1 Soft liner Soft liner material is used to reline the acrylic base temporarily for rebasing (relining) and tissue conditioning. It is available as a two-component (Fig. 23.7) powder and liquid system, which when mixed at the recommended ratio, forms a paste-like consistency. Advantages: Fast application No heat irritation Easy to trim Good retention Smooth surface minimizes unwanted food debris retention Simple to remove from the denture at the end of the treatment This material can be used in following conditions: To reline immediate acrylic denture base after extraction to reline and improve retention. To rebase the ill-fitting dentures to improve the adaptation with denture bearing areas. To reline the denture base temporarily to improve the health of inflamed tissues. After surgical resection of tumors, the bulb portion of immediate surgical obturators can be reline postoperatively to close the fresh surgical defect. 5. To reline the denture base in fresh surgical sites to push the skin graft toward the tissue bed. 1. 2. 3. 4. 23.10.2 Coe-Comfort and Coe-Soft These are resilient, self-curing reline materials that are used when the patient requires a soft, temporary liner in upper or lower dentures. Polymerization takes place in or out of the mouth in approximately 15 minutes. The reline normally lasts for approximately 3 months. Coe-Comfort and Coe-Soft materials are supplied as two parts, powder and liquid (Fig. 23.8), which are mixed together to obtain a paste-like consistency.
634 Figure 23.7 Soft liner for denture base acrylic resin. Figure 23.8 Coe-Soft and Coe-Comfort reline materials. Advanced Dental Biomaterials
Biomaterials for maxillofacial prosthetic rehabilitation 635 Figure 23.9 Sculpturing wax for facial wax pattern. These are used in the following conditions: 1. 2. 3. 4. 5. Ideal for postsurgical applications; placement of immediate denture. As functional impression materials during rebasing of dentures. As tissue conditioners to recover the health of tissues. To perform the functional molding of bulb portion of speech aid prosthesis. Postoperatively to reline the obturators to close the oral communication for the purpose of improved healing and prevention from infection. 23.10.3 Sculpturing clays and waxes Modeling clays for use in maxillofacial prosthetics are available for sculpturing and carving for face prosthesis. These clays are sulfur-free to prevent contamination of molds before packing in platinum vulcanizing silicones. These are available in different colors and hardnesses (soft, medium, hard) to carve according to different depth and layers (Fig. 23.9). Waxes, on the other hand, are easy to sculpt and need less artistic skills than clays due to the advantage of being easy to melt and adjust during trial procedures. 23.11 Coloring agents 23.11.1 Colored flocking The skin is composed of different layers with different chroma value hues for everyone; therefore it is important to match the skin tone individually as much as possible to give a lifelike appearance. Flocking coloring agents are fine hair-like nylon or rayon flocks of 0.75 mm length (Fig. 23.10), which are added into maxillofacial silicone and painted on to the mold with stains to give the superficial skin
636 Advanced Dental Biomaterials Figure 23.10 Rayon flocking for silicone elastomers (Flocking). tone. Flocking along with stains mimic the shade of melanomas, sebaceous glands, and tiny blood vessels beneath the skin layers. 23.11.2 Intrinsic stains The intrinsic colors are added in maxillofacial silicone during mixing of base and catalyst according to the individual skin tone. These stains are created by the addition of cosmetic pigments within silicone cross-linking fluids to make silicone stains. These stains chemically bind with addition-cured platinum silicones, therefore lasting longer in optimum conditions. Intrinsic stains have the major role in fabricating an esthetically pleasing facial prosthesis. These are available from basic colors to premixed average skin colors (Fig. 23.11). 23.11.3 Extrinsic colors Extrinsic staining is useful if the intrinsic staining is not enough or to incorporate some pigmentation to simulate the natural skin tone. Extrinsic stains are available in dry earth powder pigments and silicone thixotropic paste forms (Fig. 23.12). These stains are applied with appropriate-sized painting brushes after thorough cleaning of prosthesis with acetone or alcohol to remove the greasy debris and dust. 23.11.4 Acetoxy silastic adhesives The silicone adhesives are condensation-cured silicones, supplied as onecomponent, translucent silicones (Fig. 23.13) with low slump property. The acetoxy
Biomaterials for maxillofacial prosthetic rehabilitation 637 Figure 23.11 Silicone intrinsic colors compatible with platinum-cured silicones (Functional Intrinsic Skin Colors). Figure 23.12 Thixotropic silicone extrinsic stains (Extrinsic colors). silicone adhesives cure at room temperature upon exposure to atmospheric moisture. During the curing process, the silicone adhesives release acetic acid vapor as a by-product. After curing, the silicone elastomer has a texture, appearance, and composition like most other conventional silicone polydimethylsiloxane elastomers. The initial curing takes place within 20 minutes, while the complete curing with optimum physical properties are achieved after 24 120 hours depending on the atmospheric moisture and humidity. These medical adhesives are used to bond acrylic resin or metals to platinumcured silicone prostheses. The primary use of medical adhesives in maxillofacial prosthetics is to fix the external staining of a silicone facial prosthesis by applying a thin coating. The silicone prosthesis appears shiny upon fixation of stains with medical adhesives; therefore, solutions like MD-564 can be used to matte the appearance of the final silicone prosthesis.
638 Advanced Dental Biomaterials Figure 23.13 Acetoxy silastic silicone adhesive and dispersion liquid (Extrinsic sealing kit). Table 23.12 Properties of acetoxy silicone adhesives used for maxillofacial prostheses. Properties Medical adhesive A-100 A-564 Appearance Specific gravity Durometer hardness Tensile strength (psi) Elongation (%) Translucent 1.06 38 480 450 Translucent 1.08 28 780 650 These adhesives can also be used to bond the dislodged acrylic housing containing retentive elements to the vulcanized silicone prosthesis for repair purposes, and repair of small voids and tears in vulcanized silicone by mixing the adhesive with intrinsic stains and applying a thin layer on chair-side. This is the time-saving alternative, which can be utilized temporarily for an emergency repair (Table 23.12). 23.12 Skin adhesives Maxillofacial prostheses are retained by various methods, including craniofacial implants, skin adhesives, anatomical undercuts, and mechanical retention by means of spectacles. Craniofacial implants are considered the best option for optimum retention of facial prosthesis; however, radiotherapy and systemic health are the relative contraindications for implants placement.
Biomaterials for maxillofacial prosthetic rehabilitation 639 Figure 23.14 Daro water-based skin adhesive for retention of silicone facial prosthesis. The skin adhesives are supplied as silicone based and water based. The silicone adhesives are difficult to clean from silicone prostheses compared to water-based adhesives due to the compatibility of silicone adhesives with a silicone prosthesis. Therefore, patients have shown more compliance toward water-based adhesives. These are applied in a thin layer over the margins of the silicone prosthesis. The white appearance of adhesive changes to a clear layer upon leaving at room temperature or exposure to compressed air. The water-based adhesives are available in different strengths as their application varies according to the weight and size of silicone prosthesis. The edge adhesives are available to be used on the thin margins of a silicone prosthesis, which are more prone to tear with regular strength adhesives upon insertion and removal of the prosthesis. These adhesives provide minimal bond strength, therefore protecting the thin margins of the prosthesis. The edge adhesives are used to seal the margins and blend them with adjacent skin (Fig. 23.14). References A-306: Platinum Primer. Product Information. Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/Primer_Silicone_Bond_Elastomer_p/a-306. htm.. A-330 G: Platinum Primer. Product Information, Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/Gold_Platinum_Primer_Silicone_ Bonding_Elastomer_p/a-330-gold.htm..
640 Advanced Dental Biomaterials Andreopoulos, A.G., Evangelatou, M., 1994. Evaluation of various reinforcements for maxillofacial silicone elastomers. J. Biomater. Appl. 8 (4), 344 360. Andreopoulos, A.G., Evangelatou, M., Tarantili, P.A., 1998. Properties of maxillofacial silicone elastomers reinforced with silica powder. J. Biomater. Appl. 13 (1), 66 73. Aziz, T., Waters, M., Jagger, R., 2003a. Analysis of the properties of silicone rubber maxillofacial prosthetic materials. J. Dent. 31, 67 74. Aziz, T., Waters, M., Jagger, R., 2003b. Development of a new poly (dimethylsiloxane) maxillofacial prosthetic material. J. Biomed. Mater. Res. B: Appl. Biomater. 65 (2), 252 261. Beumer, J., Marunick, M.T., Esposito, S.J., 2011. Prosthodontic and Surgical Management of Cancer-related Acquired, and Congenital Defects of Head and Neck, third ed. Quintessence Pub. Co Inc.. Colas, A., Curtis, J., 2005. Silicone Biomaterials: History and Chemistry in Silicone Biomaterials: History and Chemistry & Medical Applications of Silicones. Dow Corning, pp. 80 86. Deanin, R.D., 1972. Polymer Structure, Properties, and Applications. Cahners, Boston, MA, p. 496. Dow Corning Corporation. Product Information, Dow Corning. Available from: ,http:// www.dowcorning.com/content/discover/discoverchem.. Extrinsic Colors. Product Information, Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/Extrinsic_Coloration_Individual_Colors_p/ fe-extrinsic-colors.htm.. Extrinsic Sealing Kit. Product Information. Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/Extrinsic_Sealing_Kit_p/564-kit.htm.. Flocking. Product Information. Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/H_Flocking_Individual_Colors_p/h-100.htm.. Functional Intrinsic Skin Colors. Product Information. Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/ Functional_Intrinsic_Skin_Colors_p/fi-sk.htm.. Gunay, Y., Kurtoglu, C., Atay, A., Karayazgan, B., Gurbuz, C.C., 2008. Effect of tulle on the mechanical properties of a maxillofacial silicone elastomer. Dent. Mater. J. 27 (6), 775 779. Huber, H., Stephan, P.S., 2002. Materials and techniques in maxillofacial prosthodontic rehabilitation. Oral Maxillofacial Surg. Clin. N. Am. 14, 73 93. Karayazgan, B., Gunay, Y., Evlioglu, G., 2003. Improved edge strength in a facial prosthesis by incorporation of tulle. J. Prosthet. Dent. 90 (6), 526 529. Keith, J.F., Steven, M.M., 2017. The glossary of prosthodontic terms 9th ed. J. Prosthet. Dent. 117 (5), 6 92. Khindria, S.K., Bansal, S., Kansal, M., 2009. Maxillofacial prosthetic materials—review article. J. Indian Prosthodont. Soc. 9 (1), 2 5. Lai J.H., 1996. Maxillofacial Prosthetic Materials, US Patent No. 5, 584, 886. Lai, J.H., Wang, L.L., Ko, C.C., DeLong, R.L., Hodges, J.S., 2002. New organosilicon maxillofacial prosthetic materials. Dent. Mater. 18, 281 286. Lewis, D.H., Castleberry, D.J., 1980. An assessment of recent advances in external maxillofacial materials. J. Prosthet. Dent. 43 (4), 426 432. Liu, Q., Shao, L.Q., Xiang, H.F., Zhen, D., Zhao, N., Yang, S.G., et al., 2013. Biomechanical characterization of a low density silicone elastomer filled with hollow microspheres for maxillofacial prostheses. J. Biomater. Sci. Polym. Ed. 24 (11), 1378 1390.
Biomaterials for maxillofacial prosthetic rehabilitation 641 MDX4-4210 Silicone Elastomer. Product Information. Available from: ,http://www.lookpolymers.com/pdf/Multibase-SILASTIC-MDX4-4210-Dow-Corning-Biomedical-GradeElastomer.pdf.. Mirabedini, S.M., Mohseni, M., Pazokifard, S.H., Esfandeh, M., 2008. Effect of TiO2 on the mechanical and adhesion properties of RTV silicone elastomer coatings. Colloids Surf. A: Physicochem. Eng. Asp. 317, 80 86. Mitra, A., Choudhary, S., Garg, H.J., 2014. Maxillofacial prosthetic materials—an inclination towards silicones. J. Clin. Diagn. Res. 8 (12), 8 13. Platinum Silicone A-2186. Product Information. Factor II, Inc., Forest Avenue, Lakeside, AZ. Available from: ,http://www.factor2.com/Silicone_Elastomer_RTV_Facial_Prosthetic_ p/a-2186.htm.. Robert, G.C., John, M.P., 2002. Restorative Dental Materials, eleventh ed. Mosby, Inc, Philadelphia, PA. Sharif, A.M., Alvin, G.W., Deborah, J.R., Scott, R.S., 2010. Maxillofacial materials reinforced with various concentrations of polyhedral silsesquioxanes. J. Dent. Biomech. 2010, 1 6. Udita, S.M., Karthik, K.S., Sudhakara, V.M., 2010. Maxillofacial prostheric materials—past & present trends. J. Acquir. Immune Defic. Syndr. 1 (2), 26 30. Yu, R., Koran, A., Craig, R.G., 1980. Physical properties of maxillofacial elastomers under conditions of accelerated aging. J. Dent. Res. 59 (6), 1041 1047.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 24 Sukumaran Anil1, Elna Paul Chalisserry2, Seung Yun Nam2,3 and Jayachandran Venkatesan4 1 Department of Dentistry, Hamad Medical Corporation, Doha, Qatar, 2Interdisciplinary Program of Marine-Biomedical, Electrical and Mechanical Engineering, Center for Marine-Integrated Biomedical Technology (BK21 Plus), Pukyong National University, Busan, South Korea, 3Department of Biomedical Engineering, Pukyong National University, Busan, South Korea, 4Yenepoya Research Centre, Yenepoya University, Mangalore, India Chapter Outline 24.1 Introduction 644 24.1.1 Scaffolds for bone tissue engineering 645 24.1.2 Functions and features of scaffolds 646 24.1.3 Classification of biomaterials 646 24.2 Natural biomaterials 24.2.1 24.2.2 24.2.3 24.2.4 24.2.5 24.2.6 24.2.7 24.2.8 647 Collagen 647 Fibrin 648 Alginate 649 Silk 650 Hyaluronate 651 Chitosan 651 Agarose 653 Elastin 653 24.3 Synthetic biomaterials 654 24.3.1 Polyethyleneglycol 654 24.3.2 Poly-e-caprolactone 654 24.3.3 Polyglycolic acid 655 24.4 Bioceramics 655 24.4.1 24.4.2 24.4.3 24.4.4 Tricalcium phosphate 655 Hydroxyapatite 656 Tricalcium phosphate/hydroxyapatite biphasic ceramics (biphasic calcium phosphate) 657 Bioactive glasses 657 24.5 Metals 658 24.5.1 Biodegradable metal scaffolds 658 24.5.2 Titanium 659 24.5.3 Zirconia 660 Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00025-6 Copyright © 2019 Elsevier Ltd. All rights reserved.
644 Advanced Dental Biomaterials 24.6 Bioactive restorative materials 661 24.6.1 Mineral trioxide aggregate 661 24.6.2 Biodentine 662 24.7 Three-dimensional printed scaffolds 24.8 Conclusion 664 References 664 24.1 663 Introduction The craniofacial skeleton is complex and serves important functional demands. Regeneration of craniofacial structures requires a sound understanding of complex developmental processes, physiology, and the remodeling characteristics (Tevlin et al., 2014). Regenerative tissue engineering deals with generating a biologic substitute capable of sustaining itself and integrating with functional native tissue from living cells and biocompatible scaffolds. Cells, scaffolds, and growth-stimulating signals are generally referred to as the tissue engineering triad, the key components of engineered tissues (Chan and Leong, 2008). Scaffolds serve as space-holders, provide a temporary support structure for the tissue that they are intended to replace, create a substrate for cells to attach, grow, proliferate, migrate, and differentiate. They also serve as a delivery vehicle for cells and facilitate their retention and distribution in the region where new tissue growth is desired (Hsu et al., 2013). Biomaterials are generally used as biocompatible scaffold systems that allow for the migration, proliferation, and differentiation of either resident or externally delivered cells, which are used to promote new bone formation. A wide variety of biomaterials have been used for craniofacial bone augmentation. An ideal scaffold should be biocompatible, biodegradable, and their degradation rates should match the healing rate of new tissues (Keane and Badylak, 2014; Jafari et al., 2017). The various factors that contribute to tissue response by the scaffolds are shape and size of the implant, chemical reactivity, the mechanism rate and byproducts of degradation, and ability of the material to promote cell material interactions (Schulte et al., 2000). Biodegradable and nonbiodegradable materials have been utilized in bone tissue engineering. These materials should be chosen based on their in vivo degradation rates with the idea that as the scaffold is degraded it should gradually be replaced by newly formed tissue. When the biomaterials are implanted in the body, cell biomaterial interactions depend on the surface characteristics of biomaterials such as chemical, topological, and macro microarchitectured conditions. Naturally derived and synthetic scaffold materials have been used to exploit the regenerative capacities of host tissues or transplanted cells. Scaffolds are defined as three-dimensional porous solid biomaterials designed to perform some or all of the following functions: (1) promote cell biomaterial interactions, cell adhesion, and extra cellular matrix (ECM) deposition; (2) permit sufficient transport of gases, nutrients, and regulatory factors to allow cell survival, proliferation, and
Biomaterials for craniofacial tissue engineering and regenerative dentistry 645 differentiation; (3) biodegrade at a controllable rate that approximates the rate of tissue regeneration under the culture conditions of interest; and (4) provoke a minimal degree of inflammation or toxicity in vivo. The tissue engineered construct for cartilage and bone repair and replacement poses a number of specific requirements on scaffold materials including biocompatibility, osteoconduction or induction, temporary mechanical support, controlled degradation, and adequate interstitial fluid flow (Elisseeff et al., 2005). Polymeric biomaterial scaffolds provide the structural support for cell attachment and subsequent tissue development. The best scaffold for an engineered tissue should imitate the (ECM) of the target tissue in its native state. Nevertheless, the multiple functions, the complex composition, and the dynamic nature of ECM in native tissues make it difficult to mimic exactly (Giannitelli et al., 2014). Bone grafts are selected based on its characteristics such as osteoinductiveness, osteoconductiveness, mechanical stability, quality of transplantable bone, and preservation techniques. Osteoconductive grafts function as a scaffold for native bone cells and stimulate progenitor cells to differentiate into osteoblasts. The application of bone grafts is based on three biological drivers: the utilization of a bone graft that will bring new bone-forming cells into the defect, the ability of the scaffold to induce bone formation, and the presence of factors that may induce osteoblastic proliferation and differentiation. 24.1.1 Scaffolds for bone tissue engineering Bone tissue engineering utilizes scaffolds to deliver biofactors including cells, genes, and proteins to generate bone, blood vessel formation and maturation into the construct. The desirable characteristics of a scaffold for bone tissue engineering should be three primary functions to ensure successful treatment of bone defects (Arvidson et al., 2011). 1. Provide the correct anatomic geometry to define and maintain the space for tissue regeneration. 2. Provide temporary mechanical load bearing within the tissue defect. 3. Enhance the regenerative capability of the chosen bio factor; a balance to a regenerative capacity. Achieving stiffness and strength equivalent to bone tissue requires minimally porous scaffolds, whereas the enhanced delivery of biofactors requires highly connected porous scaffolds that allow cell migration, vascularization, and connective tissue formation within scaffolds. The scaffold should fully integrate into the surrounding bone providing the initial three-dimensional (3D) structure which can allow the cells to adhere, proliferate, differentiate, and lay down extracellular matrix. A good scaffold is biocompatible, biodegradable, and porous for it to be used for construction of a biological tissue. This means that even the breakdown products of the scaffold should be nontoxic to the body. It should also be resorbable to allow the growing bone to take its place gradually and it should have porous 3D structures to accommodate implanted cells.
646 Advanced Dental Biomaterials The scaffold for bone regeneration should accelerate osteoinductivity and increase cell migration, proliferation, and release growth factors. Immune-potent activity by such scaffolds is acceptable. Scaffold should have strong water uptake and water-binding capacity as well as water delivery characteristics in order to absorb the inflammatory mediators and cellular elements in its architecture, preserve them, and slowly deliver the beneficial growth factors and proinflammatory mediators in order to provide an optimum environment during different stages of bone healing (Moshiri and Oryan, 2012). Scaffold should have the capability to align the callus and establish the continuity in the defect area. Assembling the growth factors to tissue-engineered scaffolds results in enhanced healing capability of the injured area (Oryan et al., 2013). 24.1.2 Functions and features of scaffolds 1. Architecture: Scaffolds should provide void volume for vascularization, new tissue formation, and remodeling so as to facilitate host tissue integration upon implantation. 2. Cyto and tissue compatibility: Scaffolds should provide support for either extraneously introduced or endogenous cells to attach, grow, and differentiate during both in vitro culture and in vivo implantation. 3. Bioactivity: Scaffolds may interact with the cellular components of the engineered tissues actively to facilitate and regulate their activities. 4. Mechanical property: Scaffolds provide mechanical and shape stability to the tissue defect. The intrinsic mechanical properties of the biomaterials used for scaffolding or their postprocessing properties should match that of the host tissue. 24.1.3 Classification of biomaterials A scaffold material must adequately reproduce the physical and chemical properties of natural bone in order to promote the attachment, proliferation, and differentiation of both seeded osteoprogenitor cells and surrounding recipient tissues. The 3D structure of a scaffold is of critical importance for the adequate bone development and for the cellular interactions (Kale et al., 2000). An overwhelming variety of substances are employed by researchers seeking to generate a scaffolding material that possesses all the desired traits of an ideal scaffold. Three individual groups of biomaterials—ceramics, synthetic polymers, and natural polymers—are used in the fabrication of scaffolds for tissue engineering. There has been widespread use of ceramic scaffolds, such as hydroxyapatite (HA) and tricalcium phosphate (TCP), for bone regeneration applications. Ceramic scaffolds are typically characterized by high mechanical stiffness, very low elasticity, and a hard-brittle surface. From a bone perspective, they exhibit excellent biocompatibility due to their chemical and structural similarity to the mineral phase of native bone. The interactions of osteogenic cells with ceramics are important for bone regeneration as ceramics are known to enhance osteoblast differentiation and proliferation (Ambrosio et al., 2001).
Biomaterials for craniofacial tissue engineering and regenerative dentistry 647 Biomaterials can be classified broadly as natural biomaterials or synthetic biomaterials. Natural biomaterials can further be classified as protein-based biomaterials and polysaccharide-based biomaterials. G G Protein-based biomaterials include collagen, fibrin, and silk. Polysaccharide-based biomaterials include chitosan (CS), alginate, and hyaluronan. Synthetic biomaterials include polymer-based biomaterials, peptide-based biomaterials, and ceramic-based biomaterials. G G G G Polymer-based biomaterials include PLGA and polyethyleneglycol (PEG). Peptide-based biomaterials include short amino acids and self-assembling peptides. Ceramic-based biomaterials include bioactive glass and hydroxyapatite. Composite biomaterials. Biomaterials used in tissue engineering and regeneration can be classified based on their original sources into different categories: G G G G Human-derived biomaterials such as human acellular dermal matrix; Animal-derived biomaterials including processed bovine bone materials and porcinebased collagen membranes; Other natural biomaterials such as polysaccharide-based biomaterials (cellulose, chitin/ CS); Synthetic biomaterials such as polymers, composites, and ceramic-based dental biomaterials. 24.2 Natural biomaterials 24.2.1 Collagen Collagen is one of the two major components of the bone, making up 89% of the organic matrix and 32% of the volumetric composition of bone (Kang et al., 2013). Because of the biomimetic properties, collagen-based biomaterials are the most highly investigated for bone regeneration (Stoppel et al., 2015; Glowacki and Mizuno, 2008). Collagen provides strength and structural stability to tissues in the body, such as the skin, blood vessels, tendon, cartilage, and bone. Native collagen and denatured collagen (gelatin), alone or in combination with other natural and synthetic polymeric fibers as well as ceramics, have been assessed for their inherent scaffold characteristics (Kruger et al., 2013). Collagen biomaterial has poor mechanical properties and swells readily when implanted in vivo due to its high hydrophilicity. Therefore collagen is commonly modified, cross-linked, or mixed with other components (polymers or ceramics) in order to tailor the physicochemical and mechanical properties of the scaffold to the requirements of the final application. Collagen is used mainly used in the forms such as injectable hydrogels, membranes and films, sponges and scaffolds, and micro nanospheres (Ferreira et al., 2012). The cross-linked collagen scaffolds can
648 Advanced Dental Biomaterials exhibit minimal inflammatory as well as encapsulation responses and nonfibrotic cellular growth (Ahmed et al., 2008). The major advantages of using collagen of natural origin are its profusion, low index of immunological reaction, and its capacity to form fibers from soluble preparations which are similar to those found in natural tissues (Kadler, 2004). Collagen is resorbable with high swelling ability, low antigenicity, cytocompatibility, and tissue regeneration potential (Abou Neel et al., 2012). Collagen-bioactive ceramic composites can be fabricated in many forms such as 3D scaffolds, hydrogels, or dry powders (Villa et al., 2015; Laydi et al., 2013; Zhang et al., 2012). Collagen/CS/β-TCP (CCTP) based tissue engineering scaffolds were tested (Bian et al., 2016). In order to improve the regeneration ability of scaffold, raloxifene-loaded poly lactic-co-glycolic acid (PLGA) microsphere was embedded in the CCTP scaffold. To overcome the lack of neovascularization in bone tissue engineering, cobalt ion-incorporated bioactive glass/collagen glycosaminoglycan (GAG) scaffolds were developed which enhanced the expression of vascular endothelial growth factor by stabilizing the hypoxia inducible transcription factor (Quinlan et al., 2015). Induced pluripotent stem cell derived mesenchymal stem cells cultured on collagen/hydroxy apatite (HAP)/CS nanofibers showed enhanced osteogenic differentiation (Xie et al., 2015). A collagen network was incorporated into porous calcium phosphate ceramic by vacuum infusion, which was then coated with HAP by biomimetic mineralization. This three-level hierarchical CaP-collagen-HAP scaffold demonstrated better mechanical strength together with rapid and greater bone formation than normal CaP scaffolds when implanted ectopically in the dorsal muscles of rabbits (Zhou et al., 2014). Collagen CS gels reinforced with bioactive glass nanoparticles exhibited thermosensitive behavior. This gel can be injected in the body and is shown to have a great potential in bone tissue engineering on account of its relative ease of use and minimal surgical procedure (Moreira et al., 2016). Collagen hydrogels have also been used to test the ability of mesenchymal stem cells (MSCs) from different sources to undergo osteogenic differentiation (Schneider et al., 2010). Several other materials are incorporated into collagen in order to enhance mechanical properties and enhanced bone matrix interface strength. 3D collagen scaffolds were used to culture a wide variety of stem cells for different tissue engineering applications. 24.2.2 Fibrin Fibrin has captured the interest of bone tissue engineers due to its excellent biocompatibility, controllable biodegradability, and ability to deliver cells and biomolecules. Fibrin is one of the preferred biomaterials for tissue engineering applications. Fibrin gels promote cell migration, proliferation, and matrix synthesis through the incorporation of platelet-derived growth factors and transforming growth factor. Fibrin is particularly appealing because its precursors, fibrinogen, and thrombin, which can be derived from the patient’s own blood, enable the fabrication of completely autologous scaffolds (Noori et al., 2017). Addition of fibrin may also be used to enhance the behavior of cells on other natural polymers. Conjunction of
Biomaterials for craniofacial tissue engineering and regenerative dentistry 649 fibrin on CS/nano-β-TCP composites markedly improved cellular functions on scaffolds including MSC attachment, density, proliferation, differentiation, and mineralization (Siddiqui and Pramanik, 2015). Studies have shown that fibrin alone is incapable of healing bone defects, and so, addition of other materials and/or biomolecules could yield better outcomes in tissue engineering (Linsley et al., 2016). The combination of fibrin with osteoconductive ceramics that allows the incorporation of growth factors, drugs, and metallic ions may enable the fabrication of multifunctional scaffolds to promote bone tissue remodeling (Kim et al., 2014). Hence fibrin is an attractive biomaterial for bone tissue engineering because of its excellent biocompatibility, biodegradability, intrinsic bioactivity, and other unique characteristics. 24.2.3 Alginate Alginate is a hydrogel comprising 1,4-linked b-D-mannuronic acid and a-L-guluronic acid, typically derived from brown seaweed and also bacteria. It has been extensively investigated and used for many biomedical applications, due to its biocompatibility, low toxicity, relatively low cost, and simple gelation by addition of divalent cations such as Ca21, Mg21, Ba21, and Sr21 (Wee and Gombotz, 1998). Alginate is a copolymer of D-mannuronic acid (M) and L-guluronic acid (G). The gelation of alginate occurs as blocks of guluronic acid bound to other G blocks via divalent cations, usually calcium ions (Wong, 2004). The modification of the chemical structure of alginate by combining it with other biopolymers can further extend the properties of this versatile material (Caterson et al., 2002). Alginate is readily processable for 3D scaffolding materials such as hydrogels, microspheres, microcapsules, sponges, foams and fibers, nanoparticles, and multilayers at physiological conditions ensuring the preservation of cell viability and function (Sun and Tan, 2013). Alginate allows efficient cell penetration into matrix and cell encapsulation. The alginate/hydroxyapatite composite scaffolds were prepared by internal gelation followed by a freeze drying procedure to obtain a porous structure. The nanoparticles can be prepared in the presence of a lactose-modified CS, which is adsorbed on the scaffolds and can be used as temporary resorbable bone implants (Marsich et al., 2013). Two different types of polymer scaffolds, that is, CS alginate and CS alginate with fucoidan, were developed by a freeze drying method, and each was characterized as a bone graft substitute (Venkatesan et al., 2014). CS alginate hybrid scaffolds displayed improved mechanical strength and structural stability and were shown to stimulate new bone formation and rapid vascularization. Alginate microparticle and microfiber-aggregated scaffolds were produced through the aggregation method. Such a porous structure allowed vascularization, oxygenation and cell migration, adhesion, and proliferation, which are biological events that are fundamental for bone tissue regeneration (Valente et al., 2012). An injectable and biodegradable scaffold based on oxidized alginate microbeads encapsulating periodontal ligaments and gingival mesenchymal stem cells was found to be a promising strategy for bone tissue engineering (Moshaverinia et al., 2012).
650 Advanced Dental Biomaterials A high porosity composite scaffold comprising alginate, CS, collagen, and hydroxyapatite without chemical cross-linking agent was used as a scaffold for regenerating bone tissue (Yu et al., 2013). Another composite comprising solid free-form fabricated polycaprolactone, bone morphogenetic protein (BMP-2), or bone formation peptide, and alginate is used for bone tissue regeneration (Kim et al., 2013b). The physicochemical and biological properties of hybrid materials fabricated from photocross-linked gelatin, alginate, and gelatin/alginate-based hydrogels combined with silica particles were studied by Lewandowska-Lancucka et al. (2017). It was confirmed that the addition of silica particles to gelatin as well as gelatin/alginatebased matrices increased the elastic modulus values when compared to pristine hydrogels and are promising candidates for bioactive scaffolds in tissue engineering. One of the disadvantages of the material are the inability to control its degradation rate in vivo and its low viscoelasticity, although this can be improved by increased cross-linking or addition of other substances, such as HA (Yuan et al., 2011). Alginate and alginate/HA mixtures have been used in bone and cartilage tissue engineering (Coluccino et al., 2016; Sarker et al., 2015; Wang et al., 2016). 24.2.4 Silk Silk fibroin (SF) is a natural fibrous protein and has been shown to be a potential biomaterial for a number of biomedical applications because of its excellent biocompatibility, nontoxicity, diverse physical characteristics, and ability towards cell attachment and proliferation. Silk, a natural polypeptide having very good biological properties either alone or in combination with other suitable materials, can be fabricated as scaffolds for tissue engineering applications. SF films and fibers exhibit good ability to support the adhesion and growth of various cells and hence they are proposed as scaffolds for skin and bone regeneration (Ribeiro et al., 2018; Li et al., 2018a). SFs are characterized as natural amphiphilic block copolymers composed of hydrophobic and hydrophilic blocks which combine together to give rise to the elasticity and toughness to SF (Zhang et al., 2009). Novel SF and carboxymethyl cellulose composite nanofibrous scaffold was developed to investigate its ability to nucleate bioactive nano-sized calcium phosphate by biomineralization for bone tissue engineering application. The developed scaffold has proved to be a novel and excellent candidate for bone tissue engineering with significant improvement in physicochemical, mechanical, and biological properties in comparison to the gelatin and pure SF nanofibrous scaffolds (Singh et al., 2016). SF/CS-based magnetic scaffolds showed optimal characteristics such as superparamagnetic behavior with decreased phosphate-buffered saline (PBS) uptake and degradation and, hence, can be used for bone tissue engineering applications (Aliramaji et al., 2017). Xie et al. (2016) developed novel bioscaffolds for bone regeneration, based on the combination of SF and calcium polyphosphate (CPP). It was found that Bombyx mori SF (BMSF) and Antheraea pernyi SF (APSF), prepared and applied into CPP scaffolds with glutaraldehyde cross-linking forms of the BMSF/CPP and APSF/CPP bioceramics, could be suitable for bone tissue engineering.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 651 Silk, when mixed with hydroxyapatite, serves as a convenient material for bone and dental tissue growth of extracellular matrix (Melke et al., 2016). Silk gel material is able to create a lasting 3D soft tissue augmentation; hence, it is useful in periodontal and maxillofacial therapies. Hexafluoroisopropanol (HFIP)-based silk is slower in degradation and supports soft dental pulp formation better than aqueousbased silk. Tooth bud cells seeded on HFIP silk scaffolds, with or without incorporated arginine glycine aspartate (RGD) peptides, have been shown to form mineralized tissue indicating the usefulness of these scaffolds in osteodentine formation (Zhang et al., 2011). Micron-sized silk fibers have been incorporated as a reinforcement in compact fiber composite to produce high strength biomaterial that would serve as load bearing bone grafts. These have favored human bone derived stem cells (hBMSC) differentiation and formation of bone-like tissue suggesting their use for bone engineering applications (Mandal et al., 2012). 24.2.5 Hyaluronate Hyaluronic acid (HA), a nonsulfated GAG, present in the extracellular matrix of connective tissues offers excellent potential as a scaffold for tissue regeneration. The poor mechanical strength and rapid in vivo degradation rate can be controlled by cross-linking and chemical modification of the polymers (Ouasti et al., 2011). Hyaluronic acid is capable of specific cell interaction via the CD4 receptor which promotes wound healing and induces chondrogenesis. HA has been chemically and physically incorporated into various tissue engineering scaffold matrices (Kim et al., 2013a). Hyaluronic acid (HA) hydrogels have been widely applied in the biomedical field because of their good biocompatibility and structural diversity. They possess some advantages, including facilitating proliferation of seeded cells and wound healing, good antiinflammatory ability, and promoting intracellular signaling (Collins and Birkinshaw, 2013). Porous hyaluronic acid scaffolds used for BMP-2 delivery for bone growth showed that the BMP-2 were continuously released for controlled times in an active form from the scaffolds (Kim and Valentini, 2002). Solid free-form fabrication of polylactic-co-glycolic acid grafted HA/PEG scaffolds has successfully delivered BMP-2 in vivo with controllable release from the scaffold for up to a month. Histological analyses and staining after implantation in rats revealed active bone regeneration, and the BMP-2 released from the scaffold was thought to contribute in enhanced bone regeneration (Park et al., 2011). A composite of alginate and hyaluronic acid gel showed improved physical, mechanical, and biological properties (Ganesh et al., 2013). The RGD peptides in hyaluronic acid hydrogel enhance cellular attachment, cellular spreading, and proliferation. Hyaluronic acid gels are injectable and hence serve as suitable scaffolds for pulp regeneration (Inuyama et al., 2010). 24.2.6 Chitosan CS a linear polysaccharide, derived from partial deacetylation of chitin, is a highly versatile biomaterial from crustacean exoskeletons. CS evokes minimal
652 Advanced Dental Biomaterials foreign-body response and fibrous encapsulation. CS is especially attractive as a bone scaffold material as it supports the attachment and proliferation of boneforming osteoblast cells as well as the formation of mineralized bone matrix (Seol et al., 2004). Pure CS is a suitable substrate for adhesion and proliferation of osteoblasts as well as matrix formation and mineralization, but the strength and structural stability of pure CS scaffolds is inadequate for bone tissue engineering applications. Studies have shown that modified CS scaffolds exhibit osteoconductivity in vivo in surgically created bone defects (Muzzarelli et al., 1994). CS can be processed in multiple ways to produce a variety of 3D scaffolds with different pore structures for use in bone tissue engineering. It can also be combined with different materials including ceramics and polymers to yield composite scaffolds with superior mechanical and biological properties (Levengood and Zhang, 2014). A major limitation in using CS as a standalone polymer for tissue engineering applications is its faster degradation. To overcome this problem, usually CS is either modified by adding other biocompatible polymers or by cross-linking with suitable cross linkers. Blending other polymers/bioactive materials in CS matrices resulted in functionally improved materials with respect to physicochemical and biological properties, and hence, CS-based nanofibers have wide relevance for their use in the field of bone tissue engineering (Balagangadharan et al., 2017). Modified CS, such as imidazole-modified CS and methylpyrrolidinone CS, are osteoconductive in vivo thereby promoting bone regeneration in surgically created bone defects (Kim et al., 2002). Importantly, CS is hydrophilic; that it should support adhesion and proliferation of cells. In vitro studies have demonstrated that CS promotes the adhesion and proliferation of osteogenic cells and mesenchymal stem cells. Osteogenic cells cultured on CS deposit extracellular matrix which becomes mineralized to yield bone tissue. In addition, CS has also been implicated in promoting osteogenic differentiation of mesenchymal stem cells (Yang et al., 2009; Mathews et al., 2011). CS combined with bioactive ceramics have a predominant role in bone tissue engineering as these composites exhibit tailored physical, biological, and mechanical properties favorable for bone regeneration as well as predictable degradation behavior. The presence of micro- or nanoscale calcium phosphate particles in CS scaffolds provides important functionalities, but the incorporation or immobilization of a second or even third functional component can further improve bone tissue engineering scaffolds’ structural stability and osteogenic response and thereby its clinical use. The incorporation of nanohydroxyapatite (nHAp) in the CS matrix by in situ nucleation in the presence of its precursor would provide distinct nanophase distribution of HAp particles and also improve the mechanical strength of the composite compared to simple mixing of nHAp in CS. In the case of fibrous scaffolds, addition of binders or synthetic polymers helps in the easy incorporation of nHAp into the organic solution (Shalumon et al., 2013; Wang et al., 2014). CS/nHAp prepared by cross-linking with glutaraldehyde when implanted to the calvarial bone defect showed new bone formation which was histologically confirmed as bioactive and neovascularized (Ma et al., 2011).
Biomaterials for craniofacial tissue engineering and regenerative dentistry 653 Composite CS scaffolds, which incorporate other biocompatible polymers and/or micro- or nanoscale calcium phosphates, usually possess mechanical and biological properties superior to pure CS scaffolds. Mesenchymal stem cells introduced into a defect site may either actively participate in bone formation by synthesizing mineralized matrix or secrete trophic factors that activate endogenous cells. Heparin can be immobilized on CS scaffolds as a way to exploit naturally occurring noncovalent interactions between growth factors and extracellular matrix components. 24.2.7 Agarose Agarose is a natural polysaccharide polymer having unique characteristics that give reason to consider it for tissue engineering applications. It is one of the two principal components of agar and is purified from agar by removing agaropectin. Due to their soft tissue like mechanical properties and biocompatibility, agarose gels have been investigated as potential scaffolds for neural and cartilage tissue engineering (Zarrintaj et al., 2018). Hydroxyapatite agarose composites accelerate the new bone formation more than pure HA (Iwai et al., 2015). Biomimetic scaffold containing agarose and nHAp was used with dental pulp stem cells and periosteal stem cells to regenerate the bone (Annibali et al., 2013). Biodegradable, biocompatible, and injectable gel based on HA/agarose was developed for orthopedic, oral, and maxillofacial surgery (Watanabe et al., 2007). 24.2.8 Elastin Elastin is a key structural protein in extracellular matrix of all mammals which is responsible for the elastic action of various tissues in the body. Tropoelastin is the secreted soluble precursor of elastin. Although less abundant than collagen, elastin is present in large amounts within highly elastic tissues (Muiznieks and Keeley, 2013). Incorporation of elastin in biomaterials is especially significant when its elasticity or biological effects can be exploited. Major structural contributors to vertebrate ECM are the fibrous proteins collagen and elastin. Collagen provides tissues with essential tensile strength, enabling resistance to plastic deformation and rupture, while elastin imparts the properties of extensibility and reversible recoil, enabling tissues to withstand repetitive mechanical stress (Gosline et al., 2002; Rauscher and Pomes, 2012). Elastin can also be conveniently used in association with collagen for tissue engineering applications. Rapid growth of preosteoblast cells was demonstrated in a novel elastin-like polypeptide (ELP) -collagen composite material prepared in the hydrogel form for bone tissue engineering (Amruthwar and Janorkar, 2013). Porous polycaprolactone scaffolds impregnated with bovine elastin demonstrated improved mechanical and biological responses (Annabi et al., 2011). Chondrocytes cultured on coacervated or enzymatically cross-linked ELP hydrogels were shown to promote chondrogenesis in vitro (Betre et al., 2006).
654 24.3 Advanced Dental Biomaterials Synthetic biomaterials 24.3.1 Polyethyleneglycol PEG, also known as polyethylene oxide or polyoxyethylene, is the most commercially important polyether. Polyethers are oligomers or polymers of ethylene oxide. PEG is a synthetic polymer and offers the advantages of nontoxicity, biocompatibility, low immunogenicity, and ability to undergo in vivo degradation. It is resistant to cellular and protein adsorption and thereby reduces recognition by immune system and its rejection. In its hydrogel form, it has been widely tested as a scaffold material for tissue regeneration. PEG hydrogel modified with cell adhesion RGD peptides improves cell adhesion, cell survival, and matrix synthesis within the 3D scaffold network (Burdick and Anseth, 2002). This polymer can also help to seal cell membranes after injury, making it useful for avoiding cell death. Hydrophilic PEG hydrogels can be made through a variety of cross-linking schemes to develop scaffolds with different degradation as well as release rates. The mechanical strength of PEG hydrogel is determined by the molecular weight, cross-linking, and concentration of polymers. By adding cell binding motifs like RGD peptides into the PEG hydrogel network, cell adhesion can be considered as a crucial feature of PEG-based hydrogel (Lee et al., 2015). PEG has been used in regeneration of several tissues, including spinal cord/nerve, cardiac tissue, bone, and skin (Engebretson and Sikavitsas, 2012). Properties such as good biocompatibility, nonimmunogenicity, resistance to protein adsorption, and cell adhesion of PEG have made them more suitable hydrophilic polymers in biomedical applications including bioconjugation, surface modification, drug delivery, and tissue engineering (Zhu, 2010). Bioactive molecules such as cell adhesion ligands, enzyme-sensitive peptides, and growth factors have been incorporated into PEG hydrogels, to simulate one or more ECM biofunctions such as cell adhesion, proteolytic degradation, and growth factor-binding. 24.3.2 Poly-e-caprolactone Poly-e-caprolactone (PCL) is a semicrystalline material with good mechanical properties. PCL belongs to a family of poly(α-hydroxyl esters) and is one of the most widely used biodegradable polyesters for medical applications because of its biocompatibility, biodegradability, and flexibility (Lin et al., 1999). PCL scaffolds were used in bone tissue engineering, either alone or combined with hydroxyapatite (HA) (Weisgerber et al., 2018). PCL scaffolds are excellent in longer term use as they degrade over 2 years (Pitt et al., 1981). Compared with other polyesters, PCL has slower degradation kinematics and the degradation products are harmlessly metabolized via the tricarboxylic acid cycle (Woodward et al., 1985). Scaffolds fabricated using PCL are more resistant to hydrolysis, and consequently are capable of supporting the viability, proliferation, and differentiation status of encapsulated cells. The mechanical and degradation characteristics of PCL facilitate long-term in vitro cell culture before implantation into sites of injury. PCL scaffolds thus maintain structural integrity during the in vitro culture period, while MSCs
Biomaterials for craniofacial tissue engineering and regenerative dentistry 655 differentiate and synthesize cartilaginous matrix. The electrospinned PCL nanofibrous scaffold supported multilineage differentiation of the human MSCs when cultured under a specific, differentiation-promoting culture condition (Yoshimoto et al., 2003). Electrospun composite scaffolds, made from PCL/gelatin/nHAp and PCL/gelatin, enhanced cell proliferation and odontogenic differentiation. It was shown that the PCL/HAp scaffold supported growth and osteogenic differentiation of all cells, suggesting the potential of 3D porous PCL/HAp scaffold as a good candidate material for bone tissue engineering (Chuenjitkuntaworn et al., 2016). 24.3.3 Polyglycolic acid Polyglycolic acid (PGA) is a rigid, thermoplastic, highly crystalline, aliphatic polyester, made by a ring-opening polymerization reaction of a cyclic glycolide. PGA is preferred in medical applications because its degradation products, lactic and glycolic acids, are biological metabolites. PGA has outstanding properties; however, its low solubility in organic solvents, faster degradation, and the high acidity of its degradation products have limited its biomedical applications (Nair and Laurencin, 2007). Though PGA has several applications in tissue engineering, its softness and inability to hold space make it less preferred in craniofacial defect repair. 24.4 Bioceramics Bioceramics have attracted great attention in the field of bone tissue engineering. Bioceramics are fully, partially, or noncrystalline ceramics that are designed for the regeneration and reconstruction of diseased parts of the body. Bioceramics can create strong bonds to tissues, thereby stimulating osteogenesis and improving bone graft incorporation and osseointegration (Baino et al., 2015). The dissolution rate and tissue interaction are influenced by the characteristics such as the composition, crystallinity, particle size, and porosity which can be controlled during the processing of the bioceramics (Valerio et al., 2004). Current forms of application in clinical use include solid blocks, powders, and granules for bone filling, coatings on metal joint prostheses, injectable formulations, and porous scaffolds. 24.4.1 Tricalcium phosphate The major representatives of this class are calcium phosphates that are among the most widely used crystalline ceramics for bone tissue regeneration. This is due to their exceptional properties that include (1) similarity, in terms of structure and chemical composition, to the mineral phase of bone, and (2) osteoconductivity, that is, the ability of providing a biocompatible interface along with bone migrates, and thus allows bonding to the host tissue without the formation of scar tissue (LeGeros, 2002). They are used in the form of porous ceramic pieces and granules to reconstruct all kinds of bone defects, from augmentation of alveolar ridge defects after a tooth extraction and before implant positioning to sinus reconstruction
656 Advanced Dental Biomaterials correction of various deformities and bone reconstruction following injury or disease (Liu and Lun, 2012). 24.4.2 Hydroxyapatite Hydroxyapatite has been used for a variety of biomedical applications, including matrices for drug release control and bone tissue engineering materials (Ginebra et al., 2006). Although the synthetic and natural HA differ in terms of physical microstructure, crystal size, and porosity, chemical similarities to bone, in both forms, account for the osteoconductive potential of HA (Fig. 24.1) (Ramesh et al., 2018). When compared to β-TCP, HA resorbs slowly and undergoes a little conversion to a bone-like material after implantation. There have been efforts toward developing HA-based bioceramic materials that have been doped with ions. Strontium-HA, magnesium-HA, and silicon-HA have been tested to improve mechanical and biological properties for bone tissue engineering applications (Li et al., 2018b; Panzavolta et al., 2018; Andres et al., 2018). HA is commonly used for bone repair, bone augmentation, as well as coating of implants. The low mechanical strength of normal HA ceramics restricts its use mainly to low loadbearing applications. To overcome these limitations nanocrystalline HA was Figure 24.1 3D-printed scaffold designs using synthetic biomaterials. 3D, Threedimensional.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 657 developed which showed improved sinterability and enhanced densification due to greater surface area (Zhou and Lee, 2011). 24.4.3 Tricalcium phosphate/hydroxyapatite biphasic ceramics (biphasic calcium phosphate) The biphasic calcium phosphate (BCP) concept is based on an optimum balance between the more stable phase (HA) and the more soluble phase (β-TCP). β-TCP is more rapidly replaced by bone than the hardly degradable HA, owing to its higher solubility (Petrov et al., 2001). Upon degradation of β-TCP through a process of dissolution and absorption, a normal bone structure in the regenerated bone can be achieved (Von Arx et al., 2001). The development of BCP ceramics (i.e., containing both HA and β-TCP) has provided materials in which bioactivity and degradation are controlled based on the ratios of the components (LeGeros et al., 2003). Several studies have shown that BCP granules have excellent biocompatibility and bioactivity and lead to new bone formation and degradation of the biomaterial (Valimaki et al., 2005). The rate of degradation or resorption of HA/TCP ceramics can be accelerated by increasing the amount of the more soluble phase, TCP. Studies have shown that BCPs with higher β-TCP ratio are expected to yield more replacement of biomaterial by new bone (Daculsi et al., 1990; Gauthier et al., 1998). BCP compounds containing approximately 60% of HA and 40% of β-TCP seemed to provide the optimal bone conductive properties (Rouvillain et al., 2009). Nery et al. (1992) reported optimal bone regeneration in surgically created bone defects with BCP containing HA/β-TCP ratio of 85/15. However, recently BCP products containing higher β-TCP ratios (BCP 60/40 and BCP 20/80) have been studied in prospect of its enhanced osteoconductive effect (Yang et al., 2014; Lee et al., 2013). 24.4.4 Bioactive glasses Bioactive glasses and glass ceramics, which are used in bone repair, are another type of bioceramics containing Ca, P, and Si ions with good mechanical strength. The original bioglass composition is 45% silica (SiO2), 24.5% calcium oxide (CaO), 24.5% sodium oxide (Na2O), and 6% phosphorous pentoxide (P2O5) in weight percentage (Hench and Wilson, 1984). Bioglasses are biocompatible, osteoconductive, and depending on their processing condition offer a porous structure which promotes their resorption and bone ingrowth (De Aza et al., 2003). Bioglass does not induce an inflammatory response, and the silica-based bioglass completely degrades in 6 months. Silicon has been found to promote differentiation of mesenchymal cells and bone formation. The SiO2 in combination with apatite is responsible for the inhibition of fibroblast proliferation at the bioimplant surface (Nandi et al., 2009). The borate-based bioglasses showed a faster degradation than silicabased bioglasses, but this degradation rate can be controlled by adjusting its composition. The phosphate-based bioglasses are a promising group of bioglasses for hard and soft tissue engineering, since their solubility can be controlled by manipulating
658 Advanced Dental Biomaterials the composition (Knowles, 2003). Bioglasses are quite brittle and present low mechanical strength and decreased fracture resistance. Thus their utilization should be selective or in association with other bone substitutes. Future research can be directed toward incorporating various other trace elements such as Zn, Cu, F, Mn, Sr, and B. These individual elements have been known for their osteogenic property and their incorporation in the compositional flexible glasses make them promising bioactive materials (Rahaman et al., 2011). 24.5 Metals Several metals are used in craniofacial repair and regeneration. Metals currently in use clinically include gold, stainless steel, cobalt chromium, and titanium. Most of the metals are inert alloplasts which won’t integrate with adjacent tissues or induce new bone formation. Pure titanium and some of its alloys are widely used as implant materials under load-bearing conditions in dentistry and orthopedics. Metals have been long used especially in load-bearing areas in the oral cavity. They are used as implants to replace lost teeth, reconstruction posttrauma/cancer as plates, mesh, or screws. 24.5.1 Biodegradable metal scaffolds Biodegradable metal scaffolds have shown an interesting mechanical property that was close to that of human bone with tailored degradation behavior. Metals that can degrade in the physiological environment, namely, biodegradable metals, are proposed as potential materials for hard tissue scaffolding. These biodegradable metals may have superior mechanical properties in comparison to biodegradable polymers (Yusop et al., 2012; Yazdimamaghani et al., 2017). Metals that degrade in the physiologic environment, such as magnesium alloys possessing mechanical properties that are very similar to those of bone, yet retain the ability to naturally degrade when placed within an aqueous type of environment, are considered promising scaffolds (Staiger et al., 2006). The porous architecture of Mg scaffold has been proven to play a significant role in cell growth and proliferation. Efforts to design polymer magnesium composites are ongoing. Theoretically, the polymer matrix benefits from magnesium incorporation, as magnesium may confer higher mechanical strength and fracture toughness while the polymer may prevent premature degradation (Mantripragada et al., 2013). Porous Fe and Fe-phosphorous alloys have been investigated as biodegradable bone replacements, and the results showed that addition of phosphorus increased compressive strength higher than that of pure Fe which is comparable to that of typical bone (Hermawan et al., 2010). These alloys showed also faster in vitro degradation than pure Fe, hence alloying Fe with phosphorous seems to be a promising way to optimize both mechanical and degradation properties of Fe, especially for bone scaffold (He et al., 2016).
Biomaterials for craniofacial tissue engineering and regenerative dentistry 659 Metallic nanoparticles have been incorporated into scaffolds with evidence of increased mechanical strength, increased cellular adhesion of osteoblasts and chondrocytes, and increased long-term osteoblast function, with notable improvements in collagen synthesis, alkaline phosphatase activity, and calcium deposition (Kim and Fisher, 2007; Tran and Webster, 2011). Biodegradable metals as tissue scaffolding materials have been viewed as alternative to polymers for hard tissue regeneration exploiting mostly their superior mechanical properties over biodegradable polymers. 24.5.2 Titanium Titanium is the most biocompatible and corrosion-resistant metal, its elasticity modulus corresponds to the elasticity modulus of the bone, and hence it is widely used in dentistry and orthopedics. Titanium and some titanium alloys are better received by human tissue when compared to the receptivity of other metal materials. According to the American Society for Testing and Materials, six types of titanium are listed as implant biomaterials. Depending on the various concentration of oxygen, nitrogen, hydrogen, iron, and carbon influencing the physical and mechanical properties, there are four grades of commercially pure titanium (cpTi) (Grades I IV) and two titanium (Ti) alloys (Grades V and VI)—Ti 6Al 4V and Ti 6Al 4V extra low interstitial alloys (Özcan and Hämmerle, 2012). Grades I and II are the most commonly used cpTi types for the production of metal ceramic fixed dental prosthesis. In cases where good mechanical characteristics are required, such as in bone screws and plates, Ti 6Al 4V alloy is being used (Özcan and Hämmerle, 2012; De Viteri and Fuentes, 2013). A superficial layer of titanium oxide that develops on the surface is responsible for corrosion-resistance as well as for the adhesion of glycoproteins in vivo, which is important for the biocompatibility. This oxide layer is maintained at physiological pH resisting ion formation tendency and has low reactivity with macromolecules making it highly resistant to corrosion (Sidambe, 2014). Osseointegration, a direct structural and functional connection, anchors the titanium dental implant with bone (Le Guehennec et al., 2007). Surface chemistry, topography, wettability, charge, surface energy, crystal structure, crystallinity, roughness, chemical potential, strain hardening, the presence of impurities, thickness of titanium oxide layer, and the presence of metal and nonmetal composites play a role in implant tissue interaction and osseointegration (Anil et al., 2011). Various approaches are employed to make implants more biomimetic from both biomechanical and biological perspectives. Alteration in titanium surface morphology is used to influence osteoblast attachment, differentiation, proliferation, and migration. Altering the macro-, micro-, and nanoscale surface topographies through plasma spraying, particle blasting, micromachining, grinding, polishing, and chemical methods such as acid etching, alkali etching, and anodization have resulted in higher bone implant contact (BIC) ratio and improved osseointegration (Subramani et al., 2018).
660 Advanced Dental Biomaterials Biochemical binding to bone can be achieved by coating with chemicals or biomaterials such as ceramics and polymers that are known to improve osteogenic properties. An in vivo study using micro arc oxidation method in phosphoric acid on titanium implants facilitated more chemical bonding sites for calcium ions during mineralization (Sul et al., 2002). Hydroxyapatite coating through various methods has shown precipitation of biological apatite nanocrystals with the incorporation of various proteins, which, in turn, promote cell adhesion, differentiation into osteoblast, and the synthesis of mineralized collagen, the extracellular matrix of bone tissue (Lavenus et al., 2010). Attempts have been made to improve and accelerate osseointegration by introducing bioactive factors such as known osteogenic drugs, antiinflammatory mediators, and growth factors to titanium surfaces. The effects of recombinant bone morphogenetic protein-2 (rhBMP-2) on the osseointegration of titanium implants have shown positive influences in experimental animal studies (Wikesjö et al., 2002). Tetracycline-HCl, an antimicrobial agent, kills microorganisms that may be present on the contaminated implant surface. Further, it inhibits collagenase activity and increases cell proliferation as well as attachment and bone healing (Herr et al., 2008). Oral administration of drugs such as simvastatin have proved to have a positive impact on osteogenesis; hence, Yang et al. (2011), when studying the effect of simvastatin-loaded porous implant surfaces, demonstrated accelerated osteogenic differentiation of preosteoblasts. However, there are esthetic issues due to the gray color of titanium so that in areas where soft tissue is not optimal, the dark color shines through the thin mucosa, which has led to the pursuit of esthetic replacements to titanium (Saini et al., 2015). Titanium foams fabricated by a new powder metallurgical process have bimodal pore distribution architecture, mimicking natural bone. The mechanical properties of the titanium foam with low relative densities of approximately 0.20 0.30 are close to those of human cancellous bone (Wen et al., 2011). The titanium foam after a simple thermochemical pretreatment process followed by immersion in a simulated body fluid showed good ability to form a bone-like apatite layer throughout the foam. The thermochemical pretreatments for inducing bone-like apatite formation are attractive due to their simplicity, cost savings, and effectiveness (Kapat et al., 2017). The titanium foams are anticipated to be promising implant materials for bone tissue engineering applications due to their excellent biomechanical properties and bioactivity. 24.5.3 Zirconia Zirconia, a tooth-colored material, was first introduced as endosseous implants for bone regeneration to replace mainly the impaired esthetic outcomes of dark grayish color of titanium, visible through the periimplant mucosa. These include unfavorable soft tissue conditions such as thin mucosal biotype, gingival recession, and/or in the anterior maxillary or mandibular incisor replacements. It was also claimed that ceramic implants could manage soft tissue better due to less plaque accumulation than titanium (Ozkurt and Kazazoglu, 2011; Hoffmann et al., 2008). Zirconium implants show lower inflammatory response and increased angiogenic factors
Biomaterials for craniofacial tissue engineering and regenerative dentistry 661 compared to titanium (Degidi et al., 2006). Zirconium dioxide in its pure phase is unstable and brittle with low shear strength. Hence, to improve the modulus of elasticity and stabilize the material, small amounts of aluminum oxide or yttrium oxide are added through the hot isostatic pressing method. Such additions also improve the toughness and crack propagation property. Of the two oxides, yttrium (yttriumstabilized tetragonal polycrystals) appears to offer improved advantages owing to its higher fracture resilience and higher flexure strength (Saini et al., 2015; Sennerby et al., 2005). Zirconium is used as an implant material by itself and zirconia particles as a coating material on titanium dental implants are also explored (Ozkurt and Kazazoglu, 2011). Huang et al. (2018) concluded that plasma-sprayed nanostructured-zirconium coating showed hierarchical surface morphology with better surface roughness and wettability than titanium implants. Histological analysis exhibited earlier and more condensed bone formation improving the BIC and thereby the osseointegration. Currently the majority of zirconium implants are produced as one-system implants. While microgaps between an implant and abutment is absent in such systems, the use of angular abutments to correct misalignment or secondary corrections to alter shape or vertical height is not possible as it may affect the fracture strength of zirconia (Cionca et al., 2017). Two-piece implant systems with screw-retained abutments are desirable for several reasons, although they are technically challenging because of the limitations of zirconia (Preis et al., 2016). The early failure rates of the zirconia implant systems developed and tested so far were generally higher compared with titanium implants. Technical failure as a result of fracture of the material is a sensitive issue and a critical factor for usability and acceptance in daily practice (Elnayef et al., 2017; Cionca et al., 2017). 24.6 Bioactive restorative materials The bioactivity of the dental restoratives is an important feature in the preservation of the hard dental tissues; therefore it is of utmost significance to find a material which can be used to decrease the amount of excavated dentine during preparation of the tooth. Several materials have been used as possible dentine substitutes in repairing teeth damaged by caries. Materials with clear bioactive properties seems to be the ideal solution for this application. Bioglasses have been accepted as mineralizing agents in caries prevention, as well as desensitizing agents in the treatment of hypersensitive teeth caused by opened dentinal tubules (Gjorgievska et al., 2012). 24.6.1 Mineral trioxide aggregate Mineral trioxide aggregate (MTA) is a biomaterial that has been investigated for applications in restorative dentistry. Its multiple applications include direct and
662 Advanced Dental Biomaterials indirect pulp capping, formation of apical plug, root-end filling, perforation repair, furcation repair, repair of resorptive defects, and management of immature apices (Rajasekharan et al., 2014). The main components of MTA are tricalcium oxide, tricalcium silicate, bismuth oxide, tricalcium aluminate, tricalcium oxide, tetracalcium aluminoferrite, and silicate oxide. In addition, there are a few other mineral oxides, which are responsible for the chemical and physical properties of MTA. Studies revealed that MTA materials are biocompatible with an acceptable in vivo biologic performance when used for root-end fillings, perforation repairs, pulp capping and pulpotomy, and apexification treatment (Roberts et al., 2008). 24.6.2 Biodentine Biodentine (BD), a new calcium silicate-based material, has been introduced as a dentine substitute to repair the damaged dentine. BD has a wide range of applications including endodontic repair (root perforations, apexification, resorptive lesions, and retrograde filling material in endodontic surgery) and pulp capping and can be used as a dentine replacement material in restorative dentistry. BD was originally developed for use in direct and indirect pulp capping as a single application dentine substitute without any cavity conditioning treatment. Biological studies of this material indicated that it may be safely applied directly to the pulp, since it induces the formation of a dentine-like matrix (Laurent et al., 2008). It was also confirmed that BD can be used as a dentine substitute under a composite for posterior restoration (Koubi et al., 2013). BD can be used as a direct posterior restorative material. The modified composition of the powder, the addition of setting accelerators and softeners, and a new predosed capsule formulation for use in a mixing device have largely improved the physical and handling properties of this material (Wang et al., 2011). The cement has a faster setting time than other related materials and also higher compressive strength. The components of BD are provided in the form of a capsule containing a powder and an ampule containing a liquid, which are combined by adding the liquid to the capsule and mixing with an amalgamator. According to the manufacturer, the setting time of BD is short, between 9 and 12 minutes, owing to the addition of an accelerator (CaCl2) to the liquid. The push-out bond strength of BD is similar to that of MTA (Alsubait et al., 2014). It also has excellent color stability (Valles et al., 2013) and is, therefore, more preferred than MTA for use in anterior teeth. Furthermore, similar to MTA, BD shows complete dentine bridge formation when used for pulp capping (Nowicka et al., 2013). Bioactive and biocompatible characteristics of the material were confirmed. BD when placed directly in contact with the pulp showed enhanced proliferation, migration, and adhesion of human dental pulp stem cells (Luo et al., 2014). Considering the superior physical and biologic properties, BD could be used in the field of endodontics, dental traumatology, restorative dentistry, and pediatric dentistry.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 24.7 663 Three-dimensional printed scaffolds 3D printing provides a platform that helps to fulfill the hallmarks of tissue engineering, namely, rehabilitation, reconstruction, and regeneration. The technique allows construction of individualized, patient-specific bone substitutes (Bauermeister et al., 2016). The complexity and the multicellular interactions of the craniofacial structures can be managed by the application of 3D-printed scaffolds. 3D printing is fundamentally based on the additive manufacturing technology where a 3D construct is fabricated by adding materials layer by layer. Computed tomography, magnetic resonance imaging, or X-rays are used to create a patient’s specific scaffolds. 3D bioprinting promises precise printing of various biomaterials (bio-ink) with complex 3D architectures, while being able to use multiple cell sources simultaneously and being customizable to patient-specific needs (Fig. 24.2). Various classes of bio-inks (polymer hydrogels, inert metals, ceramics, composites, and cell aggregates) may be used for 3D biomanufacturing of scaffolds to replace damaged tissue, guide in reconstruction through fixation devices, screws, custom cutting guides, as well as to create craniofacial tissue analogs for tissue regeneration. Based on the underlying printing principles, bioprinting can be classified into inkjet-based, extrusionassisted, and laser-assisted (Obregon et al., 2015; Nyberg et al., 2017; Derby, 2012). Depending on the biomaterial properties, the appropriate bioprinting system is selected. Inkjet 3D printers are capable of applying low-viscosity bio-inks using a thermal or piezoelectrical controlling system; laser-assisted printers can print cell Figure 24.2 Examples of various shapes of bioceramics for bone engineering applications: cylinder, block, root form, and granules as bone fillers.
664 Advanced Dental Biomaterials and biomaterial sources with various viscosities using laser source; and extrusion printers extrude high-viscosity and stiff polymeric sources at high temperatures (Amrollahi et al., 2016). Bone and cartilage regeneration can be achieved through scaffolds from extrusion-based printing systems as these areas demand high load bearing properties. In a clinical trial, bioprinted PCL tooth-shaped plugs have shown significant improvement in alveolar preservation (Goh et al., 2015). The possibility of combining two classes of materials, such as calcium sulfate hydrate and bioactive glass, was proven by Qi et al. Both in vitro and in vivo studies showed improved osteogenesis and thereby enhanced bone formation in calvarial defects (Qi et al., 2017). In the craniofacial region, the periodontal complex represents multiple tissues both soft and hard. Hence multiphase 3D printing using a composite (PCL-HA) with spatiotemporal arrangement and precise delivery of growth factor and BMP-2 led to the differentiation of distinctive progenitor cells of putative dentin/cementum, periodontal ligament (PDL), and alveolar bone complex (Lee et al., 2014). Similar attempts with biphasic scaffolds, cells, and gene delivery with tunable mechanical properties within the scaffold have been attempted in efforts toward bioengineering a whole tooth (Kim et al., 2010; Ikeda et al., 2009; Zhang et al., 2010). 3D-printed Grade V titanium dental implant with a novel dual-stemmed design, when tested in rabbits against conventional manufactured steel implants, showed better osteogenesis and comparable biocompatibility (Tedesco et al., 2017). Such advances promise more avenues for 3D printing in craniofacial regeneration. 24.8 Conclusion Craniofacial reconstruction has undergone tremendous expansion since the inception of the concept of tissue engineering. Various materials such as metals, ceramics, natural and synthetic polymers, and even their composites have been explored as scaffolds to promote maxillofacial tissue regeneration. The progress in this field is taking several routes, including cell biology, the development of novel scaffolds, fabrication methods, and characterization techniques. Regenerating maxillofacial structures is challenging and requires the recapitulation of the biological development of several tissues and interfaces. Stem cell therapy and engineering of irreversibly damaged tissues is progressing toward a reality. Further research should be focused on the development of materials which promote rapid tissue regeneration; the development of processing methods to give scaffolds with designed architecture; and the delivery of therapeutic molecules, such as antibiotics and growth factors. References Abou Neel, E.A., Salih, V., Revell, P.A., Young, A.M., 2012. Viscoelastic and biological performance of low-modulus, reactive calcium phosphate-filled, degradable, polymeric bone adhesives. Acta Biomater. 8, 313 320.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 665 Ahmed, T.A., Dare, E.V., Hincke, M., 2008. Fibrin: a versatile scaffold for tissue engineering applications. Tissue Eng. Part B Rev. 14, 199 215. Aliramaji, S., Zamanian, A., Mozafari, M., 2017. Super-paramagnetic responsive silk fibroin/ chitosan/magnetite scaffolds with tunable pore structures for bone tissue engineering applications. Mater. Sci. Eng. C Mater. Biol. Appl. 70, 736 744. Alsubait, S.A., Hashem, Q., Alhargan, N., Almohimeed, K., Alkahtani, A., 2014. Comparative evaluation of push-out bond strength of ProRoot MTA, bioaggregate and biodentine. J. Contemp. Dent. Pract. 15, 336 340. Ambrosio, A.M., Sahota, J.S., Khan, Y., Laurencin, C.T., 2001. A novel amorphous calcium phosphate polymer ceramic for bone repair: I. Synthesis and characterization. J. Biomed. Mater. Res. 58, 295 301. Amrollahi, P., Shah, B., Seifi, A., Tayebi, L., 2016. Recent advancements in regenerative dentistry: a review. Mater. Sci. Eng. C Mater. Biol. Appl. 69, 1383 1390. Amruthwar, S.S., Janorkar, A.V., 2013. In vitro evaluation of elastin-like polypeptide-collagen composite scaffold for bone tissue engineering. Dent. Mater. 29, 211 220. Andres, N.C., Sieben, J.M., Baldini, M., Rodriguez, C.H., Famiglietti, A., Messina, P.V., 2018. Electroactive Mg(2 1 )-hydroxyapatite nanostructured networks against drugresistant bone infection strains. ACS Appl. Mater. Interfaces 10, 19534 19544. Anil, S., Anand, P., Alghamdi, H., Jansen, J., 2011. Dental implant surface enhancement and osseointegration. Implant Dentistry—A Rapidly Evolving Practice. InTech. Annabi, N., Fathi, A., Mithieux, S.M., Martens, P., Weiss, A.S., Dehghani, F., 2011. The effect of elastin on chondrocyte adhesion and proliferation on poly(varepsilon-caprolactone)/elastin composites. Biomaterials 32, 1517 1525. Annibali, S., Cicconetti, A., Cristalli, M.P., Giordano, G., Trisi, P., Pilloni, A., et al., 2013. A comparative morphometric analysis of biodegradable scaffolds as carriers for dental pulp and periosteal stem cells in a model of bone regeneration. J. Craniofac. Surg. 24, 866 871. Arvidson, K., Abdallah, B.M., Applegate, L.A., Baldini, N., Cenni, E., Gomez-Barrena, E., et al., 2011. Bone regeneration and stem cells. J. Cell Mol. Med. 15, 718 746. Baino, F., Novajra, G., Vitale-Brovarone, C., 2015. Bioceramics and scaffolds: a winning combination for tissue engineering. Front. Bioeng. Biotechnol. 3, 202. Balagangadharan, K., Dhivya, S., Selvamurugan, N., 2017. Chitosan based nanofibers in bone tissue engineering. Int. J. Biol. Macromol. 104, 1372 1382. Bauermeister, A.J., Zuriarrain, A., Newman, M.I., 2016. Three-dimensional printing in plastic and reconstructive surgery: a systematic review. Ann. Plast. Surg. 77, 569 576. Betre, H., Ong, S.R., Guilak, F., Chilkoti, A., Fermor, B., Setton, L.A., 2006. Chondrocytic differentiation of human adipose-derived adult stem cells in elastin-like polypeptide. Biomaterials 27, 91 99. Bian, S., He, M., Sui, J., Cai, H., Sun, Y., Liang, J., et al., 2016. The self-crosslinking smart hyaluronic acid hydrogels as injectable three-dimensional scaffolds for cells culture. Colloids Surf., B: Biointerfaces 140, 392 402. Burdick, J.A., Anseth, K.S., 2002. Photoencapsulation of osteoblasts in injectable RGDmodified PEG hydrogels for bone tissue engineering. Biomaterials 23, 4315 4323. Caterson, E.J., Li, W.J., Nesti, L.J., Albert, T., Danielson, K., Tuan, R.S., 2002. Polymer/alginate amalgam for cartilage-tissue engineering. Ann. N.Y. Acad. Sci. 961, 134 138. Chan, B.P., Leong, K.W., 2008. Scaffolding in tissue engineering: general approaches and tissue-specific considerations. Eur. Spine J. 17 (Suppl. 4), 467 479. Chuenjitkuntaworn, B., Osathanon, T., Nowwarote, N., Supaphol, P., Pavasant, P., 2016. The efficacy of polycaprolactone/hydroxyapatite scaffold in combination with mesenchymal stem cells for bone tissue engineering. J. Biomed. Mater. Res. A 104, 264 271.
666 Advanced Dental Biomaterials Cionca, N., Hashim, D., Mombelli, A., 2017. Zirconia dental implants: where are we now, and where are we heading? Periodontol. 2000 73, 241 258. Collins, M.N., Birkinshaw, C., 2013. Hyaluronic acid based scaffolds for tissue engineering —a review. Carbohydr. Polym. 92, 1262 1279. Coluccino, L., Stagnaro, P., Vassalli, M., Scaglione, S., 2016. Bioactive TGF-beta1/HA alginate-based scaffolds for osteochondral tissue repair: design, realization and multilevel characterization. J. Appl. Biomater. Funct. Mater. 14, e42 e52. Daculsi, G., Passuti, N., Martin, S., Deudon, C., Legeros, R.Z., Raher, S., 1990. Macroporous calcium phosphate ceramic for long bone surgery in humans and dogs. Clinical and histological study. J. Biomed. Mater. Res. 24, 379 396. De Aza, P.N., Luklinska, Z.B., Santos, C., Guitian, F., De Aza, S., 2003. Mechanism of bone-like formation on a bioactive implant in vivo. Biomaterials 24, 1437 1445. Degidi, M., Artese, L., Scarano, A., Perrotti, V., Gehrke, P., Piattelli, A., 2006. Inflammatory infiltrate, microvessel density, nitric oxide synthase expression, vascular endothelial growth factor expression, and proliferative activity in peri-implant soft tissues around titanium and zirconium oxide healing caps. J. Periodontol. 77, 73 80. Derby, B., 2012. Printing and prototyping of tissues and scaffolds. Science 338, 921 926. De Viteri, V.S., Fuentes, E., 2013. Titanium and titanium alloys as biomaterials. Tribology— Fundamentals and Advancements. InTech. Elisseeff, J., Puleo, C., Yang, F., Sharma, B., 2005. Advances in skeletal tissue engineering with hydrogels. Orthod. Craniofac. Res. 8, 150 161. Elnayef, B., Lazaro, A., Suarez-Lopez Del Amo, F., Galindo-Moreno, P., Wang, H.L., Gargallo-Albiol, J., et al., 2017. Zirconia implants as an alternative to titanium: a systematic review and meta-analysis. Int. J. Oral. Maxillofac. Implants 32, e125 e134. Engebretson, B., Sikavitsas, V.I., 2012. Long-term in vivo effect of PEG bone tissue engineering scaffolds. J. Long Term Eff. Med. Implants 22, 211 218. Ferreira, A.M., Gentile, P., Chiono, V., Ciardelli, G., 2012. Collagen for bone tissue regeneration. Acta Biomater. 8, 3191 3200. Ganesh, N., Hanna, C., Nair, S.V., Nair, L.S., 2013. Enzymatically cross-linked alginic-hyaluronic acid composite hydrogels as cell delivery vehicles. Int. J. Biol. Macromol. 55, 289 294. Gauthier, O., Bouler, J.M., Aguado, E., Pilet, P., Daculsi, G., 1998. Macroporous biphasic calcium phosphate ceramics: influence of macropore diameter and macroporosity percentage on bone ingrowth. Biomaterials 19, 133 139. Giannitelli, S.M., Accoto, D., Trombetta, M., Rainer, A., 2014. Current trends in the design of scaffolds for computer-aided tissue engineering. Acta Biomater. 10, 580 594. Ginebra, M.P., Traykova, T., Planell, J.A., 2006. Calcium phosphate cements as bone drug delivery systems: a review. J. Controlled Release 113, 102 110. Gjorgievska, E., Nicholson, J.W., Grcev, A.T., 2012. Ion migration from fluoride-releasing dental restorative materials into dental hard tissues. J. Mater. Sci. Mater. Med. 23, 1811 1821. Glowacki, J., Mizuno, S., 2008. Collagen scaffolds for tissue engineering. Biopolymers 89, 338 344. Goh, B.T., Teh, L.Y., Tan, D.B., Zhang, Z., Teoh, S.H., 2015. Novel 3D polycaprolactone scaffold for ridge preservation—a pilot randomised controlled clinical trial. Clin. Oral. Implants Res. 26, 271 277. Gosline, J., Lillie, M., Carrington, E., Guerette, P., Ortlepp, C., Savage, K., 2002. Elastic proteins: biological roles and mechanical properties. Philos. Trans. R. Soc. Lond. B Biol. Sci. 357, 121 132.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 667 He, J., He, F.L., Li, D.W., Liu, Y.L., Yin, D.C., 2016. A novel porous Fe/Fe-W alloy scaffold with a double-layer structured skeleton: preparation, in vitro degradability and biocompatibility. Colloids Surf., B: Biointerfaces 142, 325 333. Hench, L.L., Wilson, J., 1984. Surface-active biomaterials. Science 226, 630 636. Hermawan, H., Dube, D., Mantovani, D., 2010. Degradable metallic biomaterials: design and development of Fe-Mn alloys for stents. J. Biomed. Mater. Res. A 93, 1 11. Herr, Y., Woo, J., Kwon, Y., Park, J., Heo, S., Chung, J., 2008. Implant surface conditioning with tetracycline-HCl: a SEM study. Key Engineering Materials. Trans Tech Publ, pp. 849 852. Hoffmann, O., Angelov, N., Gallez, F., Jung, R.E., Weber, F.E., 2008. The zirconia implantbone interface: a preliminary histologic evaluation in rabbits. Int. J. Oral. Maxillofac. Implants 23, 691 695. Hsu, E.L., Ghodasra, J.H., Ashtekar, A., Nickoli, M.S., Lee, S.S., Stupp, S.I., et al., 2013. A comparative evaluation of factors influencing osteoinductivity among scaffolds designed for bone regeneration. Tissue Eng. Part A 19, 1764 1772. Huang, Z., Wang, Z., Li, C., Yin, K., Hao, D., Lan, J., 2018. Application of plasma sprayed zirconia coating in dental implant: study in implant. J. Oral. Implantol. Ikeda, E., Morita, R., Nakao, K., Ishida, K., Nakamura, T., Takano-Yamamoto, T., et al., 2009. Fully functional bioengineered tooth replacement as an organ replacement therapy. Proc. Natl. Acad. Sci. U.S.A. 106, 13475 13480. Inuyama, Y., Kitamura, C., Nishihara, T., Morotomi, T., Nagayoshi, M., Tabata, Y., et al., 2010. Effects of hyaluronic acid sponge as a scaffold on odontoblastic cell line and amputated dental pulp. J. Biomed. Mater. Res. B Appl. Biomater. 92, 120 128. Iwai, S., Shimizu, H., Suzawa, Y., Akashi, M., Yura, Y., 2015. Hydroxyapatite agarose composite gels as a biochemical material for the repair of alveolar bone defects due to cleft lip and palate. J. Oral Maxillofac. Surg. Med. Pathol. 27, 637 644. Jafari, M., Paknejad, Z., Rad, M.R., Motamedian, S.R., Eghbal, M.J., Nadjmi, N., et al., 2017. Polymeric scaffolds in tissue engineering: a literature review. J. Biomed. Mater. Res. B Appl. Biomater. 105, 431 459. Kadler, K., 2004. Matrix loading: assembly of extracellular matrix collagen fibrils during embryogenesis. Birth Defects Res. C Embryo Today 72, 1 11. Kale, S., Biermann, S., Edwards, C., Tarnowski, C., Morris, M., Long, M.W., 2000. Threedimensional cellular development is essential for ex vivo formation of human bone. Nat. Biotechnol. 18, 954 958. Kang, B.J., Kim, Y., Lee, S.H., Kim, W.H., Woo, H.M., Kweon, O.K., 2013. Collagen I gel promotes homogenous osteogenic differentiation of adipose tissue-derived mesenchymal stem cells in serum-derived albumin scaffold. J. Biomater. Sci. Polym. Ed. 24, 1233 1243. Kapat, K., Srivas, P.K., Rameshbabu, A.P., Maity, P.P., Jana, S., Dutta, J., et al., 2017. Influence of porosity and pore-size distribution in Ti6Al4V foam on physicomechanical properties, osteogenesis, and quantitative validation of bone ingrowth by microcomputed tomography. ACS Appl. Mater. Interfaces 9, 39235 39248. Keane, T.J., Badylak, S.F., 2014. Biomaterials for tissue engineering applications. Semin. Pediatr. Surg. 23, 112 118. Kim, K., Fisher, J.P., 2007. Nanoparticle technology in bone tissue engineering. J. Drug Target. 15, 241 252. Kim, H.D., Valentini, R.F., 2002. Retention and activity of BMP-2 in hyaluronic acid-based scaffolds in vitro. J. Biomed. Mater. Res. 59, 573 584.
668 Advanced Dental Biomaterials Kim, I.S., Park, J.W., Kwon, I.C., Baik, B.S., Cho, B.C., 2002. Role of BMP, betaig-h3, and chitosan in early bony consolidation in distraction osteogenesis in a dog model. Plast. Reconstr. Surg. 109, 1966 1977. Kim, K., Lee, C.H., Kim, B.K., Mao, J.J., 2010. Anatomically shaped tooth and periodontal regeneration by cell homing. J. Dent. Res. 89, 842 847. Kim, I.L., Khetan, S., Baker, B.M., Chen, C.S., Burdick, J.A., 2013a. Fibrous hyaluronic acid hydrogels that direct MSC chondrogenesis through mechanical and adhesive cues. Biomaterials 34, 5571 5580. Kim, M., Jung, W.K., Kim, G., 2013b. Bio-composites composed of a solid free-form fabricated polycaprolactone and alginate-releasing bone morphogenic protein and bone formation peptide for bone tissue regeneration. Bioprocess Biosyst. Eng. 36, 1725 1734. Kim, B.S., Sung, H.M., You, H.K., Lee, J., 2014. Effects of fibrinogen concentration on fibrin glue and bone powder scaffolds in bone regeneration. J. Biosci. Bioeng. 118, 469 475. Knowles, J.C., 2003. Phosphate based glasses for biomedical applications. J. Mater. Chem. 13, 2395 2401. Koubi, G., Colon, P., Franquin, J.C., Hartmann, A., Richard, G., Faure, M.O., et al., 2013. Clinical evaluation of the performance and safety of a new dentine substitute, Biodentine, in the restoration of posterior teeth—a prospective study. Clin. Oral Investig. 17, 243 249. Kruger, T.E., Miller, A.H., Wang, J., 2013. Collagen scaffolds in bone sialoprotein-mediated bone regeneration. Sci. World J. 2013, 812718. Laurent, P., Camps, J., De Meo, M., Dejou, J., About, I., 2008. Induction of specific cell responses to a Ca(3)SiO(5)-based posterior restorative material. Dent. Mater. 24, 1486 1494. Lavenus, S., Louarn, G., Layrolle, P., 2010. Nanotechnology and dental implants. Int. J. Biomater. 2010, 915327. Laydi, F., Rahouadj, R., Cauchois, G., Stoltz, J.-F., De Isla, N., 2013. Hydroxyapatite incorporated into collagen gels for mesenchymal stem cell culture. Biomed. Mater. Eng. 23, 311 315. Lee, J.H., Ryu, M.Y., Baek, H.R., Lee, K.M., Seo, J.H., Lee, H.K., 2013. Fabrication and evaluation of porous beta-tricalcium phosphate/hydroxyapatite (60/40) composite as a bone graft extender using rat calvarial bone defect model. Sci. World J. 2013, 481789. Lee, C.H., Hajibandeh, J., Suzuki, T., Fan, A., Shang, P., Mao, J.J., 2014. Three-dimensional printed multiphase scaffolds for regeneration of periodontium complex. Tissue Eng. Part A 20, 1342 1351. Lee, T.T., Garcia, J.R., Paez, J.I., Singh, A., Phelps, E.A., Weis, S., et al., 2015. Lighttriggered in vivo activation of adhesive peptides regulates cell adhesion, inflammation and vascularization of biomaterials. Nat. Mater. 14, 352 360. LeGeros, R.Z., 2002. Properties of osteoconductive biomaterials: calcium phosphates. Clin. Orthop. Relat. Res. 81 98. Legeros, R.Z., Lin, S., Rohanizadeh, R., Mijares, D., Legeros, J.P., 2003. Biphasic calcium phosphate bioceramics: preparation, properties and applications. J. Mater. Sci. Mater. Med. 14, 201 209. Le Guehennec, L., Soueidan, A., Layrolle, P., Amouriq, Y., 2007. Surface treatments of titanium dental implants for rapid osseointegration. Dent. Mater. 23, 844 854. Levengood, S.L., Zhang, M., 2014. Chitosan-based scaffolds for bone tissue engineering. J. Mater. Chem. B 2, 3161 3184.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 669 Lewandowska-Lancucka, J., Mystek, K., Mignon, A., Van Vlierberghe, S., Latkiewicz, A., Nowakowska, M., 2017. Alginate- and gelatin-based bioactive photocross-linkable hybrid materials for bone tissue engineering. Carbohydr. Polym. 157, 1714 1722. Li, D.W., He, J., He, F.L., Liu, Y.L., Liu, Y.Y., Ye, Y.J., et al., 2018a. Silk fibroin/chitosan thin film promotes osteogenic and adipogenic differentiation of rat bone marrow-derived mesenchymal stem cells. J. Biomater. Appl. 32, 1164 1173. Li, N., Wang, Z., Zhang, Y., Zhang, K., Xie, J., Liu, Y., et al., 2018b. Curcumin-loaded redox-responsive mesoporous silica nanoparticles for targeted breast cancer therapy. Artif. Cells Nanomed. Biotechnol. 46, 921 935. Lin, W.-J., Flanagan, D.R., Linhardt, R.J., 1999. A novel fabrication of poly(ε-caprolactone) microspheres from blends of poly(ε-caprolactone) and poly(ethylene glycol)s. Polymer 40, 1731 1735. Linsley, C.S., Wu, B.M., Tawil, B., 2016. Mesenchymal stem cell growth on and mechanical properties of fibrin-based biomimetic bone scaffolds. J. Biomed. Mater. Res. A 104, 2945 2953. Liu, B., Lun, D.X., 2012. Current application of beta-tricalcium phosphate composites in orthopaedics. Orthop. Surg. 4, 139 144. Luo, Z., Li, D., Kohli, M.R., Yu, Q., Kim, S., He, W.X., 2014. Effect of Biodentine on the proliferation, migration and adhesion of human dental pulp stem cells. J. Dent. 42, 490 497. Ma, X., Wang, Y., Guo, H., Wang, J., 2011. Nano-hydroxyapatite/chitosan sponge-like biocomposite for repairing of rat calvarial critical-sized bone defect. J. Bioact. Compat. Polym. 26, 335 346. Mandal, B.B., Grinberg, A., Gil, E.S., Panilaitis, B., Kaplan, D.L., 2012. High-strength silk protein scaffolds for bone repair. Proc. Natl. Acad. Sci. U.S.A. 109, 7699 7704. Mantripragada, V.P., Lecka-Czernik, B., Ebraheim, N.A., Jayasuriya, A.C., 2013. An overview of recent advances in designing orthopedic and craniofacial implants. J. Biomed. Mater. Res. A 101, 3349 3364. Marsich, E., Bellomo, F., Turco, G., Travan, A., Donati, I., Paoletti, S., 2013. Nanocomposite scaffolds for bone tissue engineering containing silver nanoparticles: preparation, characterization and biological properties. J. Mater. Sci. Mater. Med. 24, 1799 1807. Mathews, S., Gupta, P.K., Bhonde, R., Totey, S., 2011. Chitosan enhances mineralization during osteoblast differentiation of human bone marrow-derived mesenchymal stem cells, by upregulating the associated genes. Cell Prolif. 44, 537 549. Melke, J., Midha, S., Ghosh, S., Ito, K., Hofmann, S., 2016. Silk fibroin as biomaterial for bone tissue engineering. Acta Biomater. 31, 1 16. Moreira, C.D., Carvalho, S.M., Mansur, H.S., Pereira, M.M., 2016. Thermogelling chitosan collagen bioactive glass nanoparticle hybrids as potential injectable systems for tissue engineering. Mater. Sci. Eng., C 58, 1207 1216. Moshaverinia, A., Chen, C., Akiyama, K., Ansari, S., Xu, X., Chee, W.W., et al., 2012. Alginate hydrogel as a promising scaffold for dental-derived stem cells: an in vitro study. J. Mater. Sci. Mater. Med. 23, 3041 3051. Moshiri, A., Oryan, A., 2012. Role of tissue engineering in tendon reconstructive surgery and regenerative medicine: current concepts, approaches and concerns. Hard Tissue 1, 11. Muiznieks, L.D., Keeley, F.W., 2013. Molecular assembly and mechanical properties of the extracellular matrix: a fibrous protein perspective. Biochim. Biophys. Acta 1832, 866 875.
670 Advanced Dental Biomaterials Muzzarelli, R.A., Mattioli-Belmonte, M., Tietz, C., Biagini, R., Ferioli, G., Brunelli, M.A., et al., 1994. Stimulatory effect on bone formation exerted by a modified chitosan. Biomaterials 15, 1075 1081. Nair, L.S., Laurencin, C.T., 2007. Biodegradable polymers as biomaterials. Prog. Polym. Sci. 32, 762 798. Nandi, S.K., Kundu, B., Datta, S., De, D.K., Basu, D., 2009. The repair of segmental bone defects with porous bioglass: an experimental study in goat. Res. Vet. Sci. 86, 162 173. Nery, E.B., Legeros, R.Z., Lynch, K.L., Lee, K., 1992. Tissue response to biphasic calcium phosphate ceramic with different ratios of HA/beta TCP in periodontal osseous defects. J. Periodontol. 63, 729 735. Noori, A., Ashrafi, S.J., Vaez-Ghaemi, R., Hatamian-Zaremi, A., Webster, T.J., 2017. A review of fibrin and fibrin composites for bone tissue engineering. Int. J. Nanomed. 12, 4937 4961. Nowicka, A., Lipski, M., Parafiniuk, M., Sporniak-Tutak, K., Lichota, D., Kosierkiewicz, A., et al., 2013. Response of human dental pulp capped with biodentine and mineral trioxide aggregate. J. Endod. 39, 743 747. Nyberg, E.L., Farris, A.L., Hung, B.P., Dias, M., Garcia, J.R., Dorafshar, A.H., et al., 2017. 3D-printing technologies for craniofacial rehabilitation, reconstruction, and regeneration. Ann. Biomed. Eng. 45, 45 57. Obregon, F., Vaquette, C., Ivanovski, S., Hutmacher, D.W., Bertassoni, L.E., 2015. Threedimensional bioprinting for regenerative dentistry and craniofacial tissue engineering. J. Dent. Res. 94, 143S 152SS. Oryan, A., Moshiri, A., Parizi Meimandi, A., Silver, I.A., 2013. A long-term in vivo investigation on the effects of xenogenous based, electrospun, collagen implants on the healing of experimentally-induced large tendon defects. J. Musculoskelet. Neuronal. Interact. 13, 353 367. Ouasti, S., Donno, R., Cellesi, F., Sherratt, M.J., Terenghi, G., Tirelli, N., 2011. Network connectivity, mechanical properties and cell adhesion for hyaluronic acid/PEG hydrogels. Biomaterials 32, 6456 6470. Özcan, M., Hämmerle, C., 2012. Titanium as a reconstruction and implant material in dentistry: advantages and pitfalls. Materials 5, 1528 1545. Ozkurt, Z., Kazazoglu, E., 2011. Zirconia dental implants: a literature review. J. Oral. Implantol. 37, 367 376. Panzavolta, S., Torricelli, P., Casolari, S., Parrilli, A., Fini, M., Bigi, A., 2018. Strontiumsubstituted hydroxyapatite-gelatin biomimetic scaffolds modulate bone cell response. Macromol. Biosci. 18, e1800096. Park, J.K., Shim, J.H., Kang, K.S., Yeom, J., Jung, H.S., Kim, J.Y., et al., 2011. Solid freeform fabrication of tissue-engineering scaffolds with a poly(lactic-co-glycolic acid) grafted hyaluronic acid conjugate encapsulating an intact bone morphogenetic protein 2/poly(ethylene glycol) complex. Adv. Funct. Mater. 21, 2906 2912. Petrov, O., Dyulgerova, E., Petrov, L., Popova, R., 2001. Characterization of calcium phosphate phases obtained during the preparation of sintered biphase Ca-P ceramics. Mater. Lett. 48, 162 167. Pitt, C.G., Gratzl, M.M., Kimmel, G.L., Surles, J., Schindler, A., 1981. Aliphatic polyesters II. The degradation of poly(DL-lactide), poly(epsilon-caprolactone), and their copolymers in vivo. Biomaterials 2, 215 220. Preis, V., Kammermeier, A., Handel, G., Rosentritt, M., 2016. In vitro performance of twopiece zirconia implant systems for anterior application. Dent. Mater. 32, 765 774.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 671 Qi, X., Pei, P., Zhu, M., Du, X., Xin, C., Zhao, S., et al., 2017. Three dimensional printing of calcium sulfate and mesoporous bioactive glass scaffolds for improving bone regeneration in vitro and in vivo. Sci. Rep. 7, 42556. Quinlan, E., Thompson, E.M., Matsiko, A., O’brien, F.J., Lopez-Noriega, A., 2015. Functionalization of a collagen-hydroxyapatite scaffold with osteostatin to facilitate enhanced bone regeneration. Adv. Healthcare Mater. 4, 2649 2656. Rahaman, M.N., Day, D.E., Bal, B.S., Fu, Q., Jung, S.B., Bonewald, L.F., et al., 2011. Bioactive glass in tissue engineering. Acta Biomater. 7, 2355 2373. Rajasekharan, S., Martens, L.C., Cauwels, R.G., Verbeeck, R.M., 2014. Biodentine material characteristics and clinical applications: a review of the literature. Eur. Arch. Paediatr. Dent. 15, 147 158. Ramesh, N., Moratti, S.C., Dias, G.J., 2018. Hydroxyapatite-polymer biocomposites for bone regeneration: a review of current trends. J. Biomed. Mater. Res. B Appl. Biomater. 106, 2046 2057. Rauscher, S., Pomes, R., 2012. Structural disorder and protein elasticity. Adv. Exp. Med. Biol. 725, 159 183. Ribeiro, V.P., Pina, S., Oliveira, J.M., Reis, R.L., 2018. Silk fibroin-based hydrogels and scaffolds for osteochondral repair and regeneration. Adv. Exp. Med. Biol. 1058, 305 325. Roberts, H.W., Toth, J.M., Berzins, D.W., Charlton, D.G., 2008. Mineral trioxide aggregate material use in endodontic treatment: a review of the literature. Dent. Mater. 24, 149 164. Rouvillain, J.L., Lavalle, F., Pascal-Mousselard, H., Catonne, Y., Daculsi, G., 2009. Clinical, radiological and histological evaluation of biphasic calcium phosphate bioceramic wedges filling medial high tibial valgisation osteotomies. Knee 16, 392 397. Saini, M., Singh, Y., Arora, P., Arora, V., Jain, K., 2015. Implant biomaterials: a comprehensive review. World J. Clin. Cases 3, 52 57. Sarker, A., Amirian, J., Min, Y.K., Lee, B.T., 2015. HAp granules encapsulated oxidized alginate-gelatin-biphasic calcium phosphate hydrogel for bone regeneration. Int. J. Biol. Macromol. 81, 898 911. Schneider, R.K., Puellen, A., Kramann, R., Raupach, K., Bornemann, J., Knuechel, R., Perez-Bouza, A., Neuss, S., 2010. The osteogenic differentiation of adult bone marrow and perinatal umbilical mesenchymal stem cells and matrix remodelling in three-dimensional collagen scaffolds. Biomaterials 31, 467 480. Schulte, M., Schultheiss, M., Hartwig, E., Wilke, H.J., Wolf, S., Sokiranski, R., et al., 2000. Vertebral body replacement with a bioglass-polyurethane composite in spine metastases—clinical, radiological and biomechanical results. Eur. Spine J. 9, 437 444. Sennerby, L., Dasmah, A., Larsson, B., Iverhed, M., 2005. Bone tissue responses to surfacemodified zirconia implants: a histomorphometric and removal torque study in the rabbit. Clin. Implant Dent. Relat. Res. 7 (Suppl. 1), S13 S20. Seol, Y.J., Lee, J.Y., Park, Y.J., Lee, Y.M., Young, K., Rhyu, I.C., et al., 2004. Chitosan sponges as tissue engineering scaffolds for bone formation. Biotechnol. Lett. 26, 1037 1041. Shalumon, K.T., Sowmya, S., Sathish, D., Chennazhi, K.P., Nair, S.V., Jayakumar, R., 2013. Effect of incorporation of nanoscale bioactive glass and hydroxyapatite in PCL/chitosan nanofibers for bone and periodontal tissue engineering. J. Biomed. Nanotechnol. 9, 430 440. Sidambe, A.T., 2014. Biocompatibility of advanced manufactured titanium implants—a review. Materials 7, 8168 8188.
672 Advanced Dental Biomaterials Siddiqui, N., Pramanik, K., 2015. Development of fibrin conjugated chitosan/nano β-TCP composite scaffolds with improved cell supportive property for bone tissue regeneration. J. Appl. Polym. Sci. 132. Singh, B.N., Panda, N.N., Mund, R., Pramanik, K., 2016. Carboxymethyl cellulose enables silk fibroin nanofibrous scaffold with enhanced biomimetic potential for bone tissue engineering application. Carbohydr. Polym. 151, 335 347. Staiger, M.P., Pietak, A.M., Huadmai, J., Dias, G., 2006. Magnesium and its alloys as orthopedic biomaterials: a review. Biomaterials 27, 1728 1734. Stoppel, W.L., Ghezzi, C.E., Mcnamara, S.L., Black III, L.D., Kaplan, D.L., 2015. Clinical applications of naturally derived biopolymer-based scaffolds for regenerative medicine. Ann. Biomed. Eng. 43, 657 680. Subramani, K., Mathew, R.T., Pachauri, P., 2018. Titanium surface modification techniques for dental implants—from microscale to nanoscale, Emerging Nanotechnologies in Dentistry, second ed. Elsevier. Sul, Y.T., Johansson, C.B., Kang, Y., Jeon, D.G., Albrektsson, T., 2002. Bone reactions to oxidized titanium implants with electrochemical anion sulphuric acid and phosphoric acid incorporation. Clin. Implant Dent. Relat. Res. 4, 78 87. Sun, J., Tan, H., 2013. Alginate-based biomaterials for regenerative medicine applications. Materials (Basel) 6, 1285 1309. Tedesco, J., Lee, B.E.J., Lin, A.Y.W., Binkley, D.M., Delaney, K.H., Kwiecien, J.M., et al., 2017. Osseointegration of a 3D printed stemmed titanium dental implant: a pilot study. Int. J. Dent. 2017, 5920714. Tevlin, R., Mcardle, A., Atashroo, D., Walmsley, G.G., Senarath-Yapa, K., Zielins, E.R., et al., 2014. Biomaterials for craniofacial bone engineering. J. Dent. Res. 93, 1187 1195. Tran, P.A., Webster, T.J., 2011. Selenium nanoparticles inhibit Staphylococcus aureus growth. Int. J. Nanomed. 6, 1553 1558. Valente, J.F.A., Valente, T.A.M., Alves, P., Ferreira, P., Silva, A., Correia, I.J., 2012. Alginate based scaffolds for bone tissue engineering. Mater. Sci. Eng., C 32, 2596 2603. Valerio, P., Pereira, M.M., Goes, A.M., Leite, M.F., 2004. The effect of ionic products from bioactive glass dissolution on osteoblast proliferation and collagen production. Biomaterials 25, 2941 2948. Valimaki, V.V., Yrjans, J.J., Vuorio, E., Aro, H.T., 2005. Combined effect of BMP-2 gene transfer and bioactive glass microspheres on enhancement of new bone formation. J. Biomed. Mater. Res. A 75, 501 509. Valles, M., Mercade, M., Duran-Sindreu, F., Bourdelande, J.L., Roig, M., 2013. Influence of light and oxygen on the color stability of five calcium silicate-based materials. J. Endod. 39, 525 528. Venkatesan, J., Bhatnagar, I., Kim, S.K., 2014. Chitosan-alginate biocomposite containing fucoidan for bone tissue engineering. Mar. Drugs 12, 300 316. Villa, M.M., Wang, L., Huang, J., Rowe, D.W., Wei, M., 2015. Bone tissue engineering with a collagen-hydroxyapatite scaffold and culture expanded bone marrow stromal cells. J. Biomed. Mater. Res. B Appl. Biomater. 103, 243 253. Von Arx, T., Cochran, D.L., Hermann, J.S., Schenk, R.K., Buser, D., 2001. Lateral ridge augmentation using different bone fillers and barrier membrane application. A histologic and histomorphometric pilot study in the canine mandible. Clin. Oral. Implants Res. 12, 260 269.
Biomaterials for craniofacial tissue engineering and regenerative dentistry 673 Wang, Z., Jiang, T., Sauro, S., Pashley, D.H., Toledano, M., Osorio, R., et al., 2011. The dentine remineralization activity of a desensitizing bioactive glass-containing toothpaste: an in vitro study. Aust. Dent. J. 56, 372 381. Wang, F., Zhang, Y.C., Zhou, H., Guo, Y.C., Su, X.X., 2014. Evaluation of in vitro and in vivo osteogenic differentiation of nano-hydroxyapatite/chitosan/poly(lactide-co-glycolide) scaffolds with human umbilical cord mesenchymal stem cells. J. Biomed. Mater. Res. A 102, 760 768. Wang, M.O., Bracaglia, L., Thompson, J.A., Fisher, J.P., 2016. Hydroxyapatite-doped alginate beads as scaffolds for the osteoblastic differentiation of mesenchymal stem cells. J. Biomed. Mater. Res. A 104, 2325 2333. Watanabe, J., Kashii, M., Hirao, M., Oka, K., Sugamoto, K., Yoshikawa, H., et al., 2007. Quick-forming hydroxyapatite/agarose gel composites induce bone regeneration. J. Biomed. Mater. Res. A 83, 845 852. Wee, S., Gombotz, W.R., 1998. Protein release from alginate matrices. Adv. Drug Deliv. Rev. 31, 267 285. Weisgerber, D.W., Milner, D.J., Lopez-Lake, H., Rubessa, M., Lotti, S., Polkoff, K., et al., 2018. A mineralized collagen-polycaprolactone composite promotes healing of a porcine mandibular defect. Tissue Eng. Part A 24, 943 954. Wen, C.E., Yamada, Y., Shimojima, K., Chino, Y., Hosokawa, H., Mabuchi, M., 2011. Novel titanium foam for bone tissue engineering. J. Mater. Res. 17, 2633 2639. Wikesjö, U.M., Sorensen, R.G., Kinoshita, A., Wozney, J.M., 2002. rhBMP-2/αBSMs induces significant vertical alveolar ridge augmentation and dental implant osseointegration. Clin. Implant Dent. Relat. Res. 4, 174 182. Wong, M., 2004. Alginates in tissue engineering. Methods Mol. Biol. 238, 77 86. Woodward, S.C., Brewer, P.S., Moatamed, F., Schindler, A., Pitt, C.G., 1985. The intracellular degradation of poly(epsilon-caprolactone). J. Biomed. Mater. Res. 19, 437 444. Xie, J., Lou, X., Wang, X., Yang, L., Zhang, Y., 2015. Electrospun nanofibers of hydroxyapatite/collagen/chitosan promote osteogenic differentiation of the induced pluripotent stem cell-derived mesenchymal stem cells. J. Controlled Release 213, e53. Xie, H., Gu, Z., Li, C., Franco, C., Wang, J., Li, L., et al., 2016. A novel bioceramic scaffold integrating silk fibroin in calcium polyphosphate for bone tissue-engineering. Ceram. Int. 42, 2386 2392. Yang, X., Chen, X., Wang, H., 2009. Acceleration of osteogenic differentiation of preosteoblastic cells by chitosan containing nanofibrous scaffolds. Biomacromolecules 10, 2772 2778. Yang, F., Zhao, S.F., Zhang, F., He, F.M., Yang, G.L., 2011. Simvastatin-loaded porous implant surfaces stimulate preosteoblasts differentiation: an in vitro study. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 111, 551 556. Yang, C., Unursaikhan, O., Lee, J.S., Jung, U.W., Kim, C.S., Choi, S.H., 2014. Osteoconductivity and biodegradation of synthetic bone substitutes with different tricalcium phosphate contents in rabbits. J. Biomed. Mater. Res. B Appl. Biomater. 102, 80 88. Yazdimamaghani, M., Razavi, M., Vashaee, D., Moharamzadeh, K., Boccaccini, A.R., Tayebi, L., 2017. Porous magnesium-based scaffolds for tissue engineering. Mater. Sci. Eng. C Mater. Biol. Appl. 71, 1253 1266. Yoshimoto, H., Shin, Y.M., Terai, H., Vacanti, J.P., 2003. A biodegradable nanofiber scaffold by electrospinning and its potential for bone tissue engineering. Biomaterials 24, 2077 2082.
674 Advanced Dental Biomaterials Yu, C.-C., Chang, J.-J., Lee, Y.-H., Lin, Y.-C., Wu, M.-H., Yang, M.-C., et al., 2013. Electrospun scaffolds composing of alginate, chitosan, collagen and hydroxyapatite for applying in bone tissue engineering. Mater. Lett. 93, 133 136. Yuan, Z., Nie, H., Wang, S., Lee, C.H., Li, A., Fu, S.Y., et al., 2011. Biomaterial selection for tooth regeneration. Tissue Eng. Part B Rev. 17, 373 388. Yusop, A.H., Bakir, A.A., Shaharom, N.A., Abdul Kadir, M.R., Hermawan, H., 2012. Porous biodegradable metals for hard tissue scaffolds: a review. Int. J. Biomater. 2012, 641430. Zarrintaj, P., Manouchehri, S., Ahmadi, Z., Saeb, M.R., Urbanska, A.M., Kaplan, D.L., et al., 2018. Agarose-based biomaterials for tissue engineering. Carbohydr. Polym. 187, 66 84. Zhang, Q., Yan, S., Li, M., 2009. Silk fibroin based porous materials. Materials 2, 2276 2295. Zhang, W., Ahluwalia, I.P., Yelick, P.C., 2010. Three dimensional dental epithelialmesenchymal constructs of predetermined size and shape for tooth regeneration. Biomaterials 31, 7995 8003. Zhang, W., Ahluwalia, I.P., Literman, R., Kaplan, D.L., Yelick, P.C., 2011. Human dental pulp progenitor cell behavior on aqueous and hexafluoroisopropanol based silk scaffolds. J. Biomed. Mater. Res. A 97, 414 422. Zhang, Y., Kong, D., Yokogawa, Y., Feng, X., Tao, Y., Qiu, T., 2012. Fabrication of porous hydroxyapatite ceramic scaffolds with high flexural strength through the double slipcasting method using fine powders. J. Am. Ceram. Soc. 95, 147 152. Zhou, H., Lee, J., 2011. Nanoscale hydroxyapatite particles for bone tissue engineering. Acta Biomater. 7, 2769 2781. Zhou, C., Ye, X., Fan, Y., Ma, L., Tan, Y., Qing, F., et al., 2014. Biomimetic fabrication of a three-level hierarchical calcium phosphate/collagen/hydroxyapatite scaffold for bone tissue engineering. Biofabrication 6, 035013. Zhu, J., 2010. Bioactive modification of poly(ethylene glycol) hydrogels for tissue engineering. Biomaterials 31, 4639 4656.
Applications of silver diamine fluoride in management of dental caries 25 Wei-Te Huang, Saroash Shahid and Paul Anderson Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom Chapter Outline 25.1 Introduction 675 25.2 Brief history 676 25.3 Clinical effects of silver diamine fluoride applications on caries management 676 25.3.1 Management of coronal caries in children 677 25.3.2 Management of coronal caries in adults 683 25.3.3 Management of root caries in the elderly 684 25.4 Cariostatic mechanism of silver diamine fluoride 685 25.4.1 Cariostatic effects of silver diamine fluoride on dental mineral 685 25.4.2 Cariostatic effects of silver diamine fluoride on cariogenic bacteria 690 25.4.3 Cariostatic effects of silver diamine fluoride on organic content of dentine 692 25.5 Safety of silver diamine fluoride treatment 25.6 Conclusion 694 References 694 25.1 692 Introduction Silver diamine fluoride (SDF; Ag[NH3]2F) is an alkaline (pHB10), colorless topical agent containing silver and fluoride, which has been used as a cariostatic agent for almost 50 years (Shah et al., 2014; Mei et al., 2017; Yamaga et al., 1972). SDF combines the remineralizing effect of fluoride and the antibacterial effect of silver, which together make SDF treatment more effective in controlling caries lesions than many other fluoride treatments, such as sodium fluoride varnish (Chibinski et al., 2017; Rosenblatt et al., 2009; Shah et al., 2014). Further, as Ag1 in SDF can be stabilized by forming a silver diamine complex, [Ag(NH3)2]1, the storage period of SDF is longer than that of AgF (Liu et al., 2012b). SDF has been regarded as an efficient, affordable, effective, and safe cariostatic agent, whose application in dental caries management complies with the concept of minimally invasive dentistry (MID) (Frencken et al., 2012; Murdoch-Kinch and McLean, 2003; Wilson, 2007). SDF treatment has shown substantial anticaries Advanced Dental Biomaterials. DOI: https://doi.org/10.1016/B978-0-08-102476-8.00023-2 Copyright © 2019 Elsevier Ltd. All rights reserved.
676 Advanced Dental Biomaterials efficacy comparable to restorative dentistry in both deciduous and permanent teeth (Horst et al., 2016; Zhao et al., 2017a). However, the black staining of the treated lesion is a major disadvantage of topical SDF treatment (Horst et al., 2016; Mei et al., 2016a). Thus potassium iodide (KI) has been used as a reducing agent to remove the discoloration from SDF (Ngo et al., 2002). Nonetheless, further studies are required to verify the cariostatic efficacy of this combination treatment (Knight et al., 2005). 25.2 Brief history Since the 1970s SDF has been accepted as a therapeutic agent with an anticaries function by the Central Pharmaceutical Council of the Ministry of Health and Welfare in Japan (Yamaga et al., 1972). However, the utilization of SDF is less common in Western countries (Gao et al., 2016). Recently SDF has drawn attention from dental researchers due to its effectiveness in preventing and arresting dental caries, and its noninvasive method of application (Frencken et al., 2012; Mei et al., 2017). Shortly after being approved for clinical use by the US Food and Drug Administration (FDA) in 2014, SDF entered the US market in 2015 (Horst et al., 2016). In 2016 a current dental terminology code was approved for SDF to be used for arresting dental caries (Horst et al., 2016). 25.3 Clinical effects of silver diamine fluoride applications on caries management MID has been advocated since the early 1990s, in order to keep the teeth functional for a whole lifetime. In contrast to conventional clinical treatment, the aim of MID is to reduce the need for tissue cutting and save as much healthy tissue as possible by the enhancement of cariostatic effects, and by minimally invasive restoration (Frencken et al., 2012; Murdoch-Kinch and McLean, 2003; Wilson, 2007). Even though fluoride has a pronounced remineralization ability, and is still the cornerstone of MID for dental caries management (Hellwig and Lennon, 2004), the fluoride concentration in the oral cavity is not sufficient to inhibit cariogenic bacterial activities (Tenuta and Cury, 2010). Recently there has been a growing trend in the use of SDF in dentistry due to its antibacterial properties and noninvasive way of application (Horst et al., 2016; Wilson, 2007; Frencken et al., 2012; Mei et al., 2016a). The application of SDF for caries control conforms to the World Health Organization’s Millennium Goals and the US Institute of Medicine’s criteria for 21st century medical care (Mei et al., 2013a). However, the black staining of the SDF-treated lesion is an esthetic problem (Fig. 25.1) compared with other fluoride treatments (Horst et al., 2016; Mei et al., 2016a). One solution to this problem is the use of additional KI in the application (Ngo et al., 2002). However, further studies are required to examine the
Applications of silver diamine fluoride in management of dental caries 677 Figure 25.1 Black staining of SDF-treated teeth (Chu et al., 2014). SDF, Silver diamine fluoride. effectiveness of this combination treatment. To date many clinical studies have been carried out to investigate the anticaries effects of SDF treatment on coronal caries in children and adults, and root caries in the elderly (Contreras et al., 2017), and these are summarized in Tables 25.1 and 25.2. 25.3.1 Management of coronal caries in children As SDF application is simple and not costly and is suitable to be used for children with special health care needs (Wright and White, 2017), and for children from underprivileged and remote areas, in order to treat early childhood caries (ECC) (Chu and Lo, 2008b; Contreras et al., 2017; Fung et al., 2013; Yee et al., 2009; Cheng, 2017; Gold, 2017). One review concluded that due to a large number of caries-affected preschool children and limited resources, 38 wt.% SDF is the best choice to be used for ECC in China (Chen et al., 2018). SDF treatment can be used as an alternative for restorative treatments. The use of 38 wt.% SDF for arresting cavities in preschool children has been reported to be safe (Milgrom et al., 2018). Further, clinical studies found that deciduous caries could be arrested with a single spot application with 38 wt.% SDF without restorative intervention (Chu et al., 2002; Yee et al., 2009). Compared to other noninvasive, cariostatic treatments, SDF treatment has higher caries-arresting efficacy. It was found that a single annual application of 38 wt.% SDF arrested and prevented more caries in deciduous teeth than three times weekly application of 5 wt.% sodium fluoride varnish (Lo et al., 2001). Further, another study found that topical application of 30 wt.% SDF led to a higher caries arrest rate in cavitated deciduous teeth than using atraumatic restorative technique (ART) with glass ionomer cement (GIC) (Dos Santos et al., 2012).
Table 25.1 Clinical effects of silver diamine fluoride (SDF) applications on coronal caries in deciduous and permanent teeth. References Cariostatic effects studied Subjects/dentitions Durations Groups compared Main findings McDonald and Sheiham (1994) Arrest 52 Children (29 years old)/ deciduous 18 Gotjamanos (1997) Arrest 356 Months Caries in progress (%): Gp 1: 46.5% Gp 2: 26.6% Gp 3: 5.2% Gp 4: 11.1% Gp 5: 53.0% Caries can be treated in a nontraumatic way SDF treatment for deep caries does not affect pulp vitality Lo et al. (2001) Arrest and prevention 55 Carious deciduous teeth (from 6 to 13 years old)/ deciduous 375 Children (35 years old)/deciduous Gp 1: SnF2 Gp 2: SDF/SnF2 Gp 3: Minimal preparation 1 SDF/ SnF2 1 composite resin Gp 4: Minimal preparation 1 composite resin Gp 5: no treatment 55 Carious dentine treated with 50.9% SDF and GIC Gp 1: annual 38% SDF 1 carious tissue removal Gp 2: annual 38% SDF Gp 3: 3-monthly NaF (5%) 1 carious tissue removal Gp 4: 3-monthly NaF (5%) Gp 5: no treatment Chu et al. (2002) Arrest 375 Children (35 years old)/deciduous 30 Mean no. of new carious surfaces/arrest carious surfaces: Gp 1: 0.4/2.8 Gp 2: 0.4/3.0 Gp 3: 0.8/1.7 Gp 4: 0.6/1.5 Gp 5: 1.2/1.0 1. 38% SDF is more effective than 5% NaF in preventing and arresting caries 2. Carious tissue removal is unnecessary for SDF treatment Mean no. of arrest carious surfaces: Gp 1: 2.5 Gp 2: 2.8 Gp 3: 1.5 Gp 4: 1.5 Gp 5: 1.3 SDF is effective in caries arrest in deciduous teeth Months 18 Months Months Gp 1: annual 38% SDF 1 carious tissue removal Gp 2: annual 38% SDF Gp 3: 3-monthly 5% NaF 1 carious tissue removal Gp 4: 3-monthly NaF (5%) Gp 5: no treatment
Llodra et al. (2005) Arrest and prevention 425 Children ( $ 6 years old)/deciduous and permanent first molars 36 Braga et al. (2009) Arrest 22 Children/permanent first molars 30 Yee et al. (2009) Arrest 976 Children (39 years old)/deciduous 24 Zhi et al. (2012) Arrest 212 Children (34 years old)/deciduous 24 Dos Santos et al. (2012) Liu et al. (2012a) Arrest 91 Children/deciduous 12 Months Gp 1: CTT Months Months Months Months Prevention 501 Children (mean age 9.1 years)/permanent first molars Gp 1: biannual 38% SDF Gp 2: no treatment ( 3 min) 24 Months Gp 2: 3-time weekly 10% SDF Gp 3: GIC sealant ( 3 min) Gp 1: 38% SDF/tannic acid Gp 2: 38% SDF Gp 3: 12% SDF Gp 4: no treatment ( 2 min) Gp 1: annual 38% SDF Gp 2: biannual 38% SDF Gp 3: annual GIC Gp 1: 30% SDF Gp 2: GIC ( 3 min) Gp 1: resin sealant Gp 2: biannual 5% NaF Gp 3: annual 38% SDF Gp 4: placebo control Mean no. of new carious surfaces (deciduous/ permanent first molars): Gp 1: 0.29/0.37 Gp 2: 1.43/1.06 SDF is effective for caries reduction in deciduous teeth and first permanent molars All treatments are effective in control occlusal incipient caries Mean no. of arrest carious surfaces: Gp 1: 2.2 Gp 2: 2.1 Gp 3: 1.5 Gp 4: 1.0 38% SDF is more effective than 12% SDF in caries arrest Caries arrest rates: Gp 1: 79% Gp 2: 91% Gp 3: 82% Increasing application frequency to biannual can enhance the caries arrest rate SDF was more effective than GIC (Relative risk (RR) 5 38.6%) for caries arrest Proportions of pit/fissure sites with dentine caries: Gp 1: 1.6% Gp 2: 2.4% Gp 3: 2.2% Gp 4: 4.6% All treatments are effective in preventing pit and fissure caries in permanent molars (Continued)
Table 25.1 (Continued) References Cariostatic effects studied Subjects/dentitions Durations Groups compared Main findings Monse et al. (2012) Prevention 704 Children (68 years old)/permanent first molars 18 One-time 38% SDF application cannot prevent caries development on permanent first molar, while GIC sealant can Chu et al. (2014) Arrest Duangthip et al. (2016) Arrest Gp 1: SDF (38%) Gp 2: GIC sealant Gp: no treatment With or without tooth-brushing program ( 1 min) Patient with rampant caries was treated with 38% SDF, followed by provisional crown restoration Gp 1: annual 38% SDF Gp 2: 3-weekly 38% SDF Gp 3: 3-weekly 5% NaF ( 10 s) Fung et al. (2016) Arrest Duangthip et al. (2017) Arrest A 14-year-old Chinese boy/ permanent (rampant caries) 304 Children (34 years old)/deciduous Months 18 Months 888 Children (34 years old)/deciduous 18 371 Children (34 years old)/deciduous 30 Months Months Gp 1: annual 12% SDF Gp 2: biannual 12% SDF Gp 3: annual 38% SDF Gp 4: biannual 38% SDF Gp 1: annual 38% SDF Gp 2: 3-weekly 38% SDF Gp 3: 3-weekly 5% NaF ( 10 s) Good prognosis with pain relief Caries arrest rates: Gp 1: 40% Gp 2: 35% Gp 3: 27% Annual or 3-weekly applications of 38% SDF is more effective in caries arrest than 3-weekly application of 5% NaF varnish Caries arrest rates: Gp 1: 50% Gp 2: 55% Gp 3: 64% Gp 4: 74% SDF is more effective in cares arrest at 38% than 12% and when applied biannually rather than annually Caries arrest rates: Gp 1: 48% Gp 2: 33% Gp 3: 34%
Fung et al. (2018) Arrest Milgrom et al. (2018) Arrest 888 Children (34 years old)/deciduous 30 66 Children (26 years old)/ deciduous 1421 Days CTT, Cross-tooth-brushing technique; GIC, glass ionomer cement. Months Gp 1: annual 12% SDF Gp 2: biannual 12% SDF Gp 3: annual 38% SDF Gp 4: biannual 38% SDF Gp 1: 38% SDF Gp 2: placebo Annual application of 38% SDF is more effective than 3-weekly applications of SDF or NaF varnish Caries arrest rates: Gp 1: 55.2% Gp 2: 58.6% Gp 3: 66.9% Gp 4: 75.7% SDF is more effective in cares arrest at 38% than 12% and when applied biannually rather than annually Averaged proportion of arrested caries lesions: Gp 1: 72% Gp 2: 5% No harms were observed Topical 38% SDF is effective and safe in arresting cavities in preschool children
Table 25.2 Clinical effects of silver diamine fluoride (SDF) applications on caries in roots of the elderly. References Cariostatic effects studied Subjects/dentitions Durations Groups compared Main findings Tan et al. (2010) Prevention 306 Elders (mean age 78.8 6 6.2 years old)/ permanent roots 3 Years Zhang et al. (2013) Arrest and prevention 277 Elders (6089 years old)/ permanent roots 24 Gp 1: OHI Gp 2: OHI 1 3-monthly CHX Gp 3: OHI 1 3-monthly NaF (5%) Gp 4: OHI 1 annual 38% SDF Gp 1: OHI Gp 2: OHI 1 annual 38% SDF Gp 3: OHI 1 annual 38% SDF 1 biannual OHE Li et al. (2016) Arrest 83 Elders (mean age 72.2 6 5.8 years old)/permanent roots 30 New root caries surfaces: Gp 1: 2.5 Gp 2: 1.1 Gp 3: 0.9 Gp 4: 0.7 CHX, 5% NaF and 38% SDF were all more effective than OHI alone New/arrested root caries surfaces: Gp 1: 1.33/0.04 Gp 2: 1.00/0.28 Gp 3: 0.70/0.33 Annual 38% SDF together with biannual OHE is effective in preventing and arrest caries Root caries arrest rates: Gp 1: 45% Gp 2: 90% Gp 3: 93% 1. 38% SDF application with or without KI application is effective in root caries arrest 2. KI application does not reduce black staining of SDF in the long term Months Months CHX, Chlorhexidine; KI, potassium iodide; OHE, oral health education; OHI, oral hygiene instruction. Gp 1: no treatment Gp 2: 38% SDF Gp 3: 38% SDF/KI
Applications of silver diamine fluoride in management of dental caries 683 SDF can be used in conjunction with other dental materials for the management of caries in deciduous teeth. McDonald and Sheiham (1994) used SDF together with SnF2 as a reducing agent after minimal cavity preparation of caries in deciduous teeth, followed by the restoration with composite resin (CR). It was found that this combination treatment led to less carious lesions in progression compared to SnF2 treatment or CR restoration alone. Further, a systematic review proposed that SDF can be added to the set of other noninvasive methods to treat caries lesions in deciduous teeth, such as the Hall crown technique and sealing lesions with accessible margins to optimize the cariostatic efficacy (Horst, 2018). Different application frequencies, concentrations (1238 wt.% SDF), and procedures of SDF treatments have been used in clinical practice (Elevate Oral Care, 2016; Horst et al., 2016; Crystal and Niederman, 2016; Fung et al., 2013; SDI, 2016). Several randomized clinical trials found that biannual application of SDF led to a higher caries arrest rate in deciduous teeth than annual or three times weekly SDF applications. Further, the use of 38 wt.% SDF showed higher caries arrest rates in deciduous teeth than the use of 12 wt.% SDF (Yee et al., 2009; Zhi et al., 2012; Fung et al., 2016, 2018; Duangthip et al., 2016, 2017). Therefore based on the current evidence, the biannual application of 38 wt.% SDF is the most effective protocol. Regarding the procedures of SDF treatment the consensus is that the removal of carious tissue before the SDF topical application is unnecessary (Lo et al., 2001; Dos Santos et al., 2012). In 2017 the American Academy of Pediatric Dentistry announced a guideline of using 1 minute 38 wt.% SDF application followed by gentle air-drying in arresting carious deciduous teeth. They proposed that this procedure can lead to the best cariostatic efficacy of SDF treatment. Further, they expected similar cariostatic efficacy of SDF treatment in the carious permanent teeth by using the same procedures (Crystal et al., 2017). 25.3.2 Management of coronal caries in adults SDF treatment can arrest incipient enamel caries in permanent teeth. It has been reported that the application of 10 wt.% SDF exerted a comparable caries-arrest efficacy to that of GIC sealant treatment or tooth-brushing in first permanent molars (Braga et al., 2009). Further, SDF can be used for adults who cannot tolerate conventional treatments, or for those with extreme risks of caries such as salivary dysfunction (Horst et al., 2016; Chu and Lo, 2008b; Burgess and Vaghela, 2018). A typical example demonstrating the benefits of the application of SDF to arrest rampant dental caries in permanent teeth was reported in 2014 (Chu et al., 2014). The patient suffered from chronic oral graft versus host disease following bone marrow transplantation. One side effect of the disease made it painful for him when moving his mouth, and therefore he neglected oral hygiene, finally leading to multiple dental decay. In this case SDF was used not only for arresting the caries but also for relieving the pain resulting from hypersensitivity. Also stainless steel crowns were needed to restore the chewing function. The prognosis for the patient was good, and the application of SDF for similar situations was recommended. A treatment involving the application of SDF on cavitated lesions followed by restoration with GIC,
684 Advanced Dental Biomaterials called silver modified ARTs (SMART), was introduced. In 2016 a case reported that when SMART was applied on cavitated caries of a 71-year-old female, this resulted in a positive outcome (Fa et al., 2016). SDF treatment can prevent the development of dental caries in permanent teeth. A clinical trial reported that the biannual application of SDF could not only arrest active caries in deciduous teeth and first permanent molars, but also prevent the development of new caries (preventive fraction in deciduous dentition: 79.7%; preventive fraction on permanent first molar: 65.0%) (Llodra et al., 2005). Another study reported that the annual application of 38 wt.% SDF is effective in preventing caries development in the fissures of permanent molars (Liu et al., 2012a). SDF may also be able to prevent secondary caries. An in vitro study proposed that SDF application prior to the application of GIC or CR increased the resistance of the restorations to secondary caries (Mei et al., 2016b). Another in vitro study used SDF in conjunction with casein phosphopeptideamorphous calcium phosphatemodified GIC, and the development of secondary caries was also successfully inhibited (Zhao et al., 2017b). 25.3.3 Management of root caries in the elderly Due to the advancements in medicine and dentistry, people not only live longer, but also maintain more of their teeth into their older ages. This inevitably results in more caries, especially root caries, in this elderly population (Gluzman et al., 2013). The prevalence of root caries in the elderly has been reported to range from 29% to 89%, which increases with age (Gluzman et al., 2013; Li et al., 2016). Therefore root caries is a worldwide pandemic in need of an effective preventive treatment. SDF treatment can effectively arrest and prevent root caries. It has been shown that the annual application of 38 wt.% SDF is more effective in preventing new root caries than merely giving oral hygiene instruction (OHI) alone (Tan et al., 2010). Another study found that the annual application of SDF in conjunction with biannual OHI was effective in preventing and arresting root caries (Zhang et al., 2013). In 2013 a review concluded that the annual application of 38 wt.% SDF is the best choice for preventing primary root caries, prevailing over other preventive agents like chlorhexidine (CHX) and sodium fluoride (Gluzman et al., 2013). In addition, in 2017 a review concluded that SDF is the most cost-effective preventive treatment for root caries (Schwendicke and Gostemeyer, 2017). An annual application of SDF followed by the addition of KI can also effectively arrest root caries (Li et al., 2016). It has been proposed that the application of KI following SDF treatment will remove excessive Ag1 by forming white AgI, thereby eliminating the black staining of the treated lesion (Ngo et al., 2002). Further, topical application of SDF with the addition of KI has been proposed to have a desensitizing effect which may be another benefit of SDF treatment on root caries (Craig et al., 2012; Castillo et al., 2011). However, the long-term camouflaging effect and anticaries efficacy of this combination treatment are still questionable, and further studies are required.
Applications of silver diamine fluoride in management of dental caries 25.4 685 Cariostatic mechanism of silver diamine fluoride Even though topical SDF application has been used for preventing and arresting dental caries across the age spectrum for almost 50 years, the cariostatic mechanism of SDF remains obscure (Mei et al., 2017). The effects of SDF on dental mineral, cariogenic bacteria, and dentinal organic content, as shown in Table 25.3, have been extensively investigated in order to understand the cariostatic mechanism of SDF. Further, as SDF treatment involves the effects of fluoride and silver, cariostatic effects of fluoride and silver in SDF have been proposed (Rosenblatt et al., 2009; Zhao et al., 2017a; Yamaga et al., 1972). 25.4.1 Cariostatic effects of silver diamine fluoride on dental mineral The interaction between SDF and dental mineral has been proposed to be (Rosenblatt et al., 2009; Yamaga et al., 1972): Ca10 ðPO4 Þ6 ðOHÞ2 1 AgðNH3 Þ2 F ! CaF2 1 Ag3 PO4 1 NH4 OH (25.1) As CaF2 and Ag3PO4 are formed following topical SDF treatment, both Ca21 and PO32 4 released from demineralizing dental mineral can be preserved (Yamaga et al., 1972). Studies have reported that higher percentages of calcium and phosphate with less mineral loss were detected in demineralized dentine treated with SDF than those without treatment (Mei et al., 2013d; Chu et al., 2012). An ex vivo study has also reported that a highly remineralized zone abundant in calcium and phosphate was found on the SDF-treated dentinal carious lesion (Mei et al., 2014b). Further, the preservation of calcium and phosphate can increase the mineral content of the treated lesion, leading to higher microhardness. It was found that the application of SDF increased the mineral density of treated incipient enamel caries in permanent premolars (Punyanirun et al., 2018), and the microhardness of the SDFtreated caries was higher than that of the active carious lesions (Chu and Lo, 2008a; Mei et al., 2013d). It has been reported that SDF-treated teeth could adsorb Ca21 from a remineralization solution, and inhibit the Ca21 release from enamel treated with SDF when in demineralization solution (Zhao et al., 2017a). Further, another study reported that a higher fluoride uptake was found in SDF-treated dentine (Knight et al., 2009). These findings are consistent with the formation of CaF2 following SDF treatment (Yamaga et al., 1972). CaF2 was found in hydroxyapatite (HAP) powder, enamel powder, and dentine powder mixed with SDF (Lou et al., 2011; Zhao et al., 2017a). However, another study found that the amount of CaF2 formed following SDF treatment dropped drastically after immersion in artificial saliva (Suzuki et al., 1974). This indicates that CaF2 may not be directly involved in the demineralization inhibition effect of SDF treatment.
Table 25.3 Cariostatic effects of silver diamine fluoride (SDF) on dental mineral, cariogenic bacteria, and dentinal organic content. Cariostatic effects of SDF on dental mineral Reference Methodologies Main findings Chu et al. (2008a) Carious deciduous teeth treated with either 38% SDF or 5% NaF were extracted to undergo KHN measurement Lou et al. (2011) HAP powders mixed with 38% SDF, NaF, or AgNO3 were inspected with SEM, EDX, and ED Liu et al. (2012b) Mei et al. (2013d) Sound premolars treated with AgF, KF, AgNO3 or water, were subjected to demineralization for 7 days before being inspected with micro-CT Dentine caries generated by multispecies biofilms (Streptococcus mutans, Streptococcus sobrinus, Lactobacillus acidophilus, Lactobacillus rhamnosus, and Actinomyces naeslundii) were treated with either 38% SDF or water and were incubated in the artificial mouth for 21 days Carious deciduous teeth treated with 38% SDF were inspected with micro-CT, EDX, SEM, and TEM Calcium phosphate with different SDF concentrations (0.38, 1.52, 2.66, 3.80 mg/mL) were incubated at 37 C for 24 h. The shape and organization of the crystals were examined by bright-field TEM and electron diffraction. Unit cell parameters of the obtained crystals were determined with powder XRD. The vibrational and rotational modes of phosphate groups were analyzed with Raman microscopy Incipient caries on proximal surfaces of premolars were treated with 38% SDF before undergoing bacterial pH-cycling for 5 days. microCT then was used for mineral density evaluation The median KHN of arrested lesions (range, 2046 or 196451 MPa) were greater than those of soft lesions (range, 520 or 49196 MPa) in the outer 25200 µm Compounds formed on treated surfaces: SDF: CaF2 and Ag(s) NaF: NaF globules AgNO3: Ag3PO4 (turns black immediately) Topical applications of AgF and KF inhibited enamel demineralization, while AgNO3 application did not 1. 38% SDF inhibits multispecies cariogenic biofilm formation on dentine carious lesions 2. The hardness and percentages of calcium and phosphorus of SDFtreated samples from the outermost 50 µm were higher Mei et al. (2014b) Mei et al. (2017) Punyanirun et al. (2018) A highly remineralized zone rich in calcium and phosphate was found on the 38% SDF-treated cavitated dentinal lesion The results suggested that SDF reacted with calcium and phosphate ions and produced fluorohydroxyapatite (FHH) Compared to the use of 1000 ppm fluoride toothpaste alone, the adjunctive use of 38% SDF enhances the remineralization of initial carious lesions based on mineral density, depth, and remineralization percentage (Continued)
Table 25.3 (Continued) Cariostatic effects of SDF on dental mineral Reference Methodologies Main findings Cariostatic effects of SDF on cariogenic bacteria Knight et al. (2005) Knight et al. (2007) Knight et al. (2009) de Almeida Lde et al. (2011) Chu et al. (2012) Mei et al. (2013b) Shah et al. (2013) Targino et al. (2014) Demineralized dentine discs treated with 50.9% SDF, 50.9% SDF/KI and KI were incubated in medium of S. mutans for 14 days. Afterwards, optical density of the medium chambers was measured to determine bacterial penetration and growth Dentine discs with or without demineralization, were treated with 50.9% SDF/KI. Next, they were incubated with S. mutans for 14 days. EPMA and SEM were then used to inspect the samples. Bacterial growth was monitored by taking optical density readings Demineralized dentine discs treated with either 50.9% SDF or 50.9% SDF/KI were incubated in S. mutans before being inspected by scanning microscopy (SEM) and EPMA Antibacterial effects of different concentrations of SDF were studied using an agar diffusion method Demineralized dentine blocks were incubated with either S. mutans or A. naeslundii to be treated with 38% SDF. Lesions were then assessed by MHT, EDX, and FTIR Carious lesions were created in dentine blocks by inoculating with dualspecies biofilm (S. mutans and L. acidophilus). They were then treated with either 38% SDF or water before being incubated at 37 C for 7 days. The biofilms were evaluated by CFU, SEM, and confocal laser scanning microscopy (CLSM), while the carious lesion was inspected with XRD, FTIR, and immune-labeling S. mutans counts in the mouths of children treated with 38% SDF, fluoride varnish or APF gel. All subjects were evaluated at 72 h, and 6, 12, and 18 months of follow-up Evaluate the antimicrobial and cytotoxic activity of 38% SDF. The MIC was evaluated by the spectrophotometric microdilution method and turbidity. The MBC was evaluated in brain heart infusion plates S. mutans migrated through all dentine discs. However, the samples treated with 50.9% SDF and SDF/KI had significantly lower optical densities SDF/KI can inhibit biofilm formation composed of S. mutans The inhibition of S. mutans might be due to the presence of silver and fluoride in the outer layer (up to 450 µm) of SDF and SDF/KItreated samples, which make it more resistant to demineralization Both 12% and 30% can inhibit S. mutans 1. 38% SDF possesses antimicrobial activity against cariogenic biofilm of S. mutans or A. naeslundii, which inhibit the demineralization 2. The hardness and percentages of calcium and phosphorus of SDFtreated samples SDF had antimicrobial activity against the cariogenic biofilms composed of S. mutans and L. acidophilus 38% SDF is more effective in inhibiting S. mutans than fluoride varnish and APF gel in vivo The MIC and MBC for SDF were 33.33 6 14.43 and 50.0 µg/mL, respectively (Continued)
Table 25.3 (Continued) Cariostatic effects of SDF on dental mineral Reference Methodologies Main findings Hamama et al. (2015) Dentine discs infected with S. mutans were treated with 50.9% SDF/KI. The discs were then fractured into two halves, stained with fluorescent LIVE/DEAD stain and observed using confocal laserscanning microscopy S. mutans biofilm demineralized enamels were treated with water, 38% SDF, APF, AHF, AHF 1 CPC, or 0.2% CHX before being incubated for 2 days. Numbers of viable microorganisms in the biofilms were counted The use of the SDF/KI is effective in reducing the numbers of S. mutans in dentinal tubules Savas et al. (2015) 38% SDF showed the highest antibacterial activity over other treatments Cariostatic effects of SDF on dentinal organic content Mei et al. (2012) Mei et al. (2013c) The inhibitory effects of 12%, 30%, 38% SDF, NaF, and AgNO3 on MMPs were studied using MMP assay kits Demineralized dentine blocks were treated with 38% SDF, 10% NaF, 42% AgNO3, and water. They were then subjected to pH cycling for 8 days before being inspected with SEM, XRD, micro-CT, and spectrophotometry with a hydroxyproline assay Mei et al. (2014a) The inhibitory effects of 12%, 30%, 38% SDF, NaF, and AgNO3 on cathepsins were studied using cathepsin assay kits Thanatvarakorn et al. (2016) Bovine dentine discs were treated with 38% SDF, KF, APF, ZnF2/HCl and ZF2 before being demineralized. TMR and optical microscope were used to examine the antidemineralization and anticollagen degradation abilities, respectively Dentine discs treated with 25% AgNO3 followed by 5% NaF, 38% SDF or water were subjected to pH-cycling for 8 days before being inspected with SEM, XRD, X-ray microtomography, and spectrophotometry with a hydroxyproline assay Zhao et al. (2017c) Inhibitory effect of SDF on MMPs increased with concentration. SDF had more inhibition on MMPs than solutions of NaF and AgNO3 1. Groups treated with 38% SDF and 42% AgNO3 had significantly less hydroxyproline liberated from the dentine matrix than groups F and W (P , .01) 2. 38% SDF can inhibit demineralization and preserved collagen from degradation 1. The solutions containing Ag1 have significantly higher inhibitory effect than the solutions containing F2 only (P , .01) 2. SDF solution at all 3 tested concentrations significantly inhibited the activity of cathepsin 38% SDF and ZnF2/HCl exerted best preservative effects against demineralization 25% AgNO3 followed by 5% NaF or 38% SDF application can preserve dentinal collagen and result in lower lesion depths AHF, Ammonium hexafluorosilicate; APF, acidulated phosphate fluoride; CFU, colony forming units; CHX, chlorhexidine; CPC, cetylpyridinium chloride; ED, electron diffraction; EDX, energy-dispersive X-ray analysis; EPMA, electron probe microanalysis; FTIR, Fourier transform infrared spectroscopy; HAP, hydroxyapatite; KHN, Knoop hardness number; KI, potassium iodide; MBC, minimum bactericide concentration; MHT, microhardness testing; MIC, minimum inhibition concentration; micro-CT, microcomputed tomography; MMP, matrix metalloproteinase; TEM, transmission electron microscopy; TMR, transverse microradiography; XRD, X-ray diffraction.
Applications of silver diamine fluoride in management of dental caries 689 As the CaF2 formed on a dental surface is unstable (Yamaga et al., 1972), CaF2 will gradually dissociate. It has been proposed that CaF2 can act as reservoir of F2 for further formation of fluorapatite (FAP) through the following interactions (ten Cate, 2013; Vogel, 2011; Yamaga et al., 1972): CaF2 ! Ca21 1 2F2 (25.2) Ca10 ðPO4 Þ6 ðOHÞ2 1 2F2 ! Ca10 ðPO4 Þ6 F2 1 2OH2 (25.3) 10Ca21 1 2F2 1 6PO32 4 ! Ca10 ðPO4 Þ6 F2 (25.4) After the dissociation of CaF2 [Eq. (25.2)], the free F2 adsorbs onto enamel and further interacts with dental mineral or attracts Ca21 and PO32 4 in saliva to form FAP [Eqs. (25.3) and (25.4)]. Subsequently, an acid-resistant FAP surface can protect the dental surface against acid attack (Featherstone, 2008). Mei et al. (2017) found that fluorohydroxyapatite (FHA), which is partial FAP-substituted HAP, was formed in the calcium phosphate [Ca3(PO4)2] medium incubated with SDF (0.383.8 mg/mL) (Mei et al., 2017). As FHA is more stable than CaF2 in solution, they proposed that FHA formation is the major demineralization inhibitory mechanism of SDF treatment. Ag3PO4 is a yellow, insoluble compound (solubility 5 6.5 3 1024 g/100 mL) (Lewis, 1920), which can be gradually reduced to black metallic silver after being exposed to light and heat (Lou et al., 2011). Metallic silver, reduced from photosensitive Ag3PO4, was found in HAP powder mixed with SDF (Lou et al., 2011). However, even though Ag3PO4 was found in enamel powder mixed with SDF, it was replaced by AgCl and silver thiocyanate after being immersed in artificial saliva (Suzuki et al., 1974). Another study also found AgCl rather than Ag3PO4 in demineralized dentine treated with SDF (Mei et al., 2013c). These findings might be due to the solubility of AgCl (solubility 5 8.9 3 1025 g/100 mL) being lower than that of Ag3PO4 (solubility 5 6.5 3 1024 g/100 mL), and therefore Ag3PO4 reacts with alkali chlorides in solution to form AgCl (Zhao et al., 2017c). Further studies are required to confirm that silver compounds are formed following SDF topical treatment on dental mineral in clinical practice. The demineralization inhibition effect of SDF treatment is mainly associated with the F2 in SDF (Horst et al., 2016). A study using microcomputed tomography found that, after artificial caries-like acidic challenge, the difference between the depth of an enamel lesion treated with AgNO3 and the depth of an enamel lesion treated with deionized water was negligible, whereas KF and AgF treatments showed significant inhibitory effects on enamel demineralization (Fig. 25.2) (Liu et al., 2012b).
690 Advanced Dental Biomaterials Figure 25.2 Typical micro-CT images of tooth blocks treated with (a) AgF; (b) KF; (c) AgNO3; (d) De-ionized water (Liu et al. 2012b). No demineralization lesions were shown in AgF and KF treated tooth blocks, whereas, demineralized lesions were observed in AgNO3 and de-ionized water treated tooth bocks. 25.4.2 Cariostatic effects of silver diamine fluoride on cariogenic bacteria SDF has strong antibacterial properties (Zhao et al., 2017a). The minimum inhibition concentration and the minimum bactericide concentration of SDF are 33.33 6 14.43 and 50.0 µg/mL, respectively (Targino et al., 2014). It was found that both 12 and 38 wt.% SDF have pronounced antibacterial capacities (de Almeida Lde et al., 2011). Further, 38 wt.% SDF has been proposed to be the most potent antibacterial agent against Streptococcus mutans compared to other antibacterial agents like sodium fluoride (NaF), acidulated phosphate fluoride (APF), ammonium hexafluorosilicate, and CHX (Shah et al., 2013; Savas et al., 2015). The application
Applications of silver diamine fluoride in management of dental caries 691 Figure 25.3 Confocal laser-scanning micrographs of the growth of multispecies cariogenic biofilm on dentine treated with (A) SDF and (B) deionized water after 14 days. The red-togreen ratio represents ratio of dead-to-live bacteria. Higher ratio was shown on the SDFtreated dentine than on the deionized water-treated dentine (Mei et al., 2013d). SDF, Silver diamine fluoride. of 38 wt.% SDF can reduce the amounts of bacteria in biofilms composed of cariogenic bacteria like S. mutans, Lactobacillus acidophilus, and Actinomyces naeslundii (Fig. 25.3) (Mei et al., 2013b,d; Chu et al., 2012). In addition, SDF treatment followed by KI application has also been reported to have pronounced antibacterial capacity, which can reduce the number of S. mutans in biofilm and in dentinal tubules (Knight et al., 2005, 2007, 2009; Hamama et al., 2015). It has been proposed that with high fluoride concentrations in SDF products (141660,000 ppm F2) (Fung et al., 2013; SDI, 2016), F2 can bind to bacterial cellular components and influence the enzymes in charge of glycolysis, such as enolase and proton-extruding adenosine triphosphatase (Targino et al., 2014; Rosenblatt et al., 2009), leading to the inhibition of biofilm formation (Zhao et al., 2017a). On the other hand, it has been proposed that 1040 ppm of Ag1 is sufficient to destroy most of the cariogenic bacteria (Lansdown, 2006). It was found that 20 ppm Ag1 was lethal to S. mutans after contact of 34 hours, while 200 ppm Ag1 is instantaneously lethal to S. mutans (Thibodeau et al., 1978). It should be emphasized that it is the Ag1 that exerts the antimicrobial effects, whereas the metallic silver is relatively inert (Peng et al., 2012). The antibacterial actions of Ag1 include cell wall destruction, disturbance of bacterial enzymes, and hindrance in microbial DNA replication (Rosenblatt et al., 2009; Targino et al., 2014). Further, it has been proposed that the silver-killed bacteria (termed “zombies”) can act as a reservoir of Ag1 to kill living bacteria (Wakshlak et al., 2015). The pivotal antibacterial reaction is associated with the interaction of Ag1 with a thiol group in bacteria, which can be simplified as: A=N 2 SH 1 AgX ! A=N 2 S 2 AgX 1 HX (25.5)
692 Advanced Dental Biomaterials In Eq. (25.5) A/N represents amino (A) or nucleic (N) acids, while SH represents a thiol group and X represents an anion such as diamine fluoride (Rosenblatt et al., 2009). By binding Ag1 to SH groups, the activities of bacterial enzymes will be inhibited. Further, by binding Ag1 to the S222 in cell membrane, the permeability of the membrane will be increased, leading to cell perturbation. Moreover, the attachment of Ag1 to guanine will cause the inhibition of the DNA replication process (Peng et al., 2012). It has been proposed that the oligodynamic action against bacteria enables Ag1 to consistently elicit antimicroorganism effects even after binding (Yamaga et al., 1972). 25.4.3 Cariostatic effects of silver diamine fluoride on organic content of dentine SDF can inhibit the degradation of dentine collagen (Mei et al., 2018). It has been reported that deciduous dentine treated with SDF led to smoother surfaces with fewer collagen fibers exposed than active dentine lesions (Fig. 25.4) (Mei et al., 2014b). Further, it has been reported that more dentine collagen under an acidic challenge is preserved by SDF treatment than by NaF, KF, or APF treatment (Thanatvarakorn et al., 2016; Mei et al., 2013c). Another study reported that 25 wt. % AgNO3 treatment followed by 5 wt.% NaF varnish on dentine collagen degradation is comparable to that of 38 wt.% SDF treatment (Zhao et al., 2017c). SDF can inhibit the proteolytic activities of matrix metalloproteinases (MMPs) and cysteine cathepins, the proteolytic enzymes in charge of dentine collagen degradation (Mei et al., 2012, 2014a). The inhibitory effect of 38% SDF on MMPs has been reported to be significantly greater than 10% NaF and 42% AgNO3 that have equivalent concentrations of F2 and Ag1 (Mei et al., 2012), whereas the inhibitory effect on cathepsins is mainly due to the effect of Ag1 (Mei et al., 2014a). Furthermore, Ag1 has been suggested to preserve dentine collagen by inactivating the catalytic functions of bacteria after interacting with a reactive side chain of their collagenase (Mei et al., 2013c). F2 and Ag1 in SDF have been proposed to have a synergistic effect on preserving dentine collagen (Mei et al., 2018). 25.5 Safety of silver diamine fluoride treatment There is a broad concentration range of commercial SDF products available on the market (Fung et al., 2013; SDI, 2016). The SDF product with the lowest concentration (Caristop 12%, Biodinamica Quı́mica E Farmacéutica Ltda, 1416 ppm F2) contains a fluoride concentration similar to that of a fluoride dentifrice (1001500 ppm F2) (Mohammed et al., 2014). The SDF product with the highest concentration (Riva Star, SDI Ltd, 60,000 ppm F2) contains fluoride at a concentration much higher than that of fluoride varnish (B22,000 ppm F2) (Hazelrigg et al., 2003). Therefore safety concerns about the extremely high F2 concentrations in
Applications of silver diamine fluoride in management of dental caries 693 Figure 25.4 Scanning electron microscopy images of (A) SDF-treated dentine surface and (B) active dentine caries (Mei et al., 2014b). SDF, Silver diamine fluoride. SDF products have been expressed (Gotjamanos, 1997; Gotjamanos and Afonso, 1997; Gotjamanos and Orton, 1998). The possibility of causing fluorosis in children following application with highconcentration SDF has been raised by several studies (Gotjamanos and Orton, 1998; Gotjamanos, 1997; Gotjamanos and Afonso, 1997). However, the Dental Services of the Health Department of Western Australia has asserted that there is no report indicating that any fluorosis has developed in patients (Shah et al., 2014). According to FDA guidelines, the average lethal dose (LD50) of SDF by oral and subcutaneous administrations in animal studies are 520 and 380 mg/kg, respectively (Horst et al., 2016). It has been proposed that the concentrations of SDF products used in clinical practice are hundreds of times less than the lethal dose of SDF by oral administration (520 mg/kg) (Horst et al., 2016). Further, a clinical study suggested that the use of SDF poses no toxic risk with reference to the US
694 Advanced Dental Biomaterials Environmental Protection Agency oral reference dose (Vasquez et al., 2012). However, allergy to silver might be a contraindication of treatment, which may cause gingivitis or mucositis (Horst et al., 2016). Some patients have been reported to have tooth pain, gum pain, gum swelling, and gum bleaching after being treated with SDF (Duangthip et al., 2018). Therefore a high-concentration SDF should still be used with caution. 25.6 Conclusion In conclusion, SDF is an efficient, affordable, effective, and safe cariostatic agent, which can be used in the management of dental caries in both deciduous and permanent teeth. The cariostatic mechanism of SDF treatment is associated with the formation of fluoride and silver compounds, antibacterial properties, and anticollagen degradation efficacy. References Braga, M.M., Mendes, F.M., De benedetto, M.S., Imparato, J.C., 2009. Effect of silver diammine fluoride on incipient caries lesions in erupting permanent first molars: a pilot study. J. Dent. Child. (Chic.) 76, 2833. Burgess, J.O., Vaghela, P.M., 2018. Silver diamine fluoride: a successful anticarious solution with limits. Adv. Dent. Res. 29, 131134. Castillo, J.L., Rivera, S., Aparicio, T., Lazo, R., Aw, T.C., Mancl, L.L., et al., 2011. The short-term effects of diammine silver fluoride on tooth sensitivity: a randomized controlled trial. J. Dent. Res. 90, 203208. Chen, K.J., Gao, S.S., Duangthip, D., Lo, E.C.M., Chu, C.H., 2018. Managing early childhood caries for young children in china. Healthcare (Basel) 6 (1), 11. Cheng, L.L., 2017. Limited evidence suggesting silver diamine fluoride may arrest dental caries in children. J. Am. Dent. Assoc. 148, 120122. Chibinski, A.C., Wambier, L.M., Feltrin, J., Loguercio, A.D., Wambier, D.S., Reis, A., 2017. Silver diamine fluoride has efficacy in controlling caries progression in primary teeth: a systematic review and meta-analysis. Caries Res. 51, 527541. Chu, C.H., Lo, E.C., 2008a. Microhardness of dentine in primary teeth after topical fluoride applications. J. Dent. 36, 387391. Chu, C.H., Lo, E.C., 2008b. Promoting caries arrest in children with silver diamine fluoride: a review. Oral Health Prev. Dent. 6, 315321. Chu, C.H., Lo, E.C., Lin, H.C., 2002. Effectiveness of silver diamine fluoride and sodium fluoride varnish in arresting dentin caries in Chinese pre-school children. J. Dent. Res. 81, 767770. Chu, C.H., Mei, L., Seneviratne, C.J., Lo, E.C., 2012. Effects of silver diamine fluoride on dentine carious lesions induced by Streptococcus mutans and Actinomyces naeslundii biofilms. Int. J. Paediatr. Dent. 22, 210.
Applications of silver diamine fluoride in management of dental caries 695 Chu, C.H., Lee, A.H., Zheng, L., Mei, M.L., Chan, G.C., 2014. Arresting rampant dental caries with silver diamine fluoride in a young teenager suffering from chronic oral graft versus host disease post-bone marrow transplantation: a case report. BMC Res. Notes 7, 3. Chu, C.H., Lo, E.C., 2008a. Microhardness of dentine in primary teeth after topical fluoride applications. J Dent 36, 387391. Contreras, V., Toro, M.J., Elias-Boneta, A.R., Encarnacion-Burgos, A., 2017. Effectiveness of silver diamine fluoride in caries prevention and arrest: a systematic literature review. Gen. Dent. 65, 2229. Craig, G.G., Knight, G.M., Mcintyre, J.M., 2012. Clinical evaluation of diamine silver fluoride/potassium iodide as a dentine desensitizing agent. A pilot study. Aust. Dent. J. 57, 308311. Crystal, Y.O., Niederman, R., 2016. Silver diamine fluoride treatment considerations in children’s caries management. Pediatr. Dent. 38, 466471. Crystal, Y.O., Marghalani, A.A., Ureles, S.D., Wright, J.T., Sulyanto, R., Divaris, K., et al., 2017. Use of silver diamine fluoride for dental caries management in children and adolescents, including those with special health care needs. Pediatr. Dent. 39, 135145. de Almeida Lde, F., Cavalcanti, Y.W., Valenca, A.M., 2011. In vitro antibacterial activity of silver diamine fluoride in different concentrations. Acta Odontol. Latinoam. 24, 127131. Dos Santos JR., V.E., de Vasconcelos, F.M., Ribeiro, A.G., Rosenblatt, A., 2012. Paradigm shift in the effective treatment of caries in schoolchildren at risk. Int. Dent. J. 62, 4751. Duangthip, D., Chu, C.H., Lo, E.C., 2016. A randomized clinical trial on arresting dentine caries in preschool children by topical fluorides—18 month results. J. Dent. 44, 5763. Duangthip, D., Wong, M.C.M., Chu, C.H., Lo, E.C.M., 2017. Caries arrest by topical fluorides in preschool children: 30-month results. J. Dent. 70, 7479. Duangthip, D., Fung, M.H.T., Wong, M.C.M., Chu, C.H., Lo, E.C.M., 2018. Adverse effects of silver diamine fluoride treatment among preschool children. J. Dent. Res. 97, 395401. Elevate Oral Care, 2016. Safety Data Sheet Advantage Arrest Silver Diamine Fluoride 38% [Online]. Elevate Oral Care (accessed 06.04.17.). Fa, B.A., Jew, J.A., Wong, A., Young, D., 2016. Silver modified atraumatic restorative technique (SMART): an alternative caries prevention tool. EDU J. 3, 1824. Featherstone, J.D., 2008. Dental caries: a dynamic disease process. Aust. Dent. J. 53, 286291. Frencken, J.E., Peters, M.C., Manton, D.J., Leal, S.C., Gordan, V.V., Eden, E., 2012. Minimal intervention dentistry for managing dental caries—a review: report of a FDI task group. Int. Dent. J. 62, 223243. Fung, M.H., Wong, M.C., Lo, E.C., Chu, C., 2013. Arresting early childhood caries with silver diamine fluoride—a literature review. J. Oral Hyg. Health 1, 117. Fung, M.H.T., Duangthip, D., Wong, M.C.M., Lo, E.C.M., Chu, C.H., 2016. Arresting dentine caries with different concentration and periodicity of silver diamine fluoride. JDR Clin. Transl. Res. 1, 143152. Fung, M.H.T., Duangthip, D., Wong, M.C.M., Lo, E.C.M., Chu, C.H., 2018. Randomized clinical trial of 12% and 38% silver diamine fluoride treatment. J. Dent. Res. 97, 171178. Gao, S.S., Zhao, I.S., Hiraishi, N., Duangthip, D., Mei, M.L., Lo, E.C.M., et al., 2016. Clinical trials of silver diamine fluoride in arresting caries among children: a systematic review. JDR Clin. Transl. Res. 1, 201210.
696 Advanced Dental Biomaterials Gluzman, R., Katz, R.V., Frey, B.J., Mcgowan, R., 2013. Prevention of root caries: a literature review of primary and secondary preventive agents. Spec. Care Dentist. 33, 133140. Gold, J., 2017. Limited evidence links silver diamine fluoride and caries arrest in children. J. Evid. Based Dent. Pract. 17, 265267. Gotjamanos, T., 1997. Safety issues related to the use of silver fluoride in paediatric dentistry. Aust. Dent. J. 42, 166168. Gotjamanos, T., Afonso, F., 1997. Unacceptably high levels of fluoride in commercial preparations of silver fluoride. Aust. Dent. J. 42, 5253. Gotjamanos, T., Orton, V., 1998. Abnormally high fluoride levels in commercial preparations of 40 per cent silver fluoride solution: contraindications for use in children. Aust. Dent. J. 43, 422427. Hamama, H.H., Yiu, C.K., Burrow, M.F., 2015. Effect of silver diamine fluoride and potassium iodide on residual bacteria in dentinal tubules. Aust. Dent. J. 60, 8087. Hazelrigg, C.O., Dean, J.A., Fontana, M., 2003. Fluoride varnish concentration gradient and its effect on enamel demineralization. Pediatr. Dent. 25, 119126. Hellwig, E., Lennon, A.M., 2004. Systemic versus topical fluoride. Caries Res. 38, 258262. Horst, J.A., 2018. Silver fluoride as a treatment for dental caries. Adv. Dent. Res. 29, 135140. Horst, J.A., Ellenikiotis, H., Milgrom, P.L., 2016. UCSF protocol for caries arrest using silver diamine fluoride: rationale, indications and consent. J. Calif. Dent. Assoc. 44, 1628. Knight, G.M., Mcintyre, J.M., Craig, G.G., Mulyani, Zilm, P.S., Gully, N.J., 2005. An in vitro model to measure the effect of a silver fluoride and potassium iodide treatment on the permeability of demineralized dentine to Streptococcus mutans. Aust. Dent. J. 50, 242245. Knight, G.M., Mcintyre, J.M., Craig, G.G., Mulyani, Zilm, P.S., Gully, N.J., 2007. Differences between normal and demineralized dentine pretreated with silver fluoride and potassium iodide after an in vitro challenge by Streptococcus mutans. Aust. Dent. J. 52, 1621. Knight, G.M., Mcintyre, J.M., Craig, G.G., Mulyani, Zilm, P.S., Gully, N.J., 2009. Inability to form a biofilm of Streptococcus mutans on silver fluoride- and potassium iodidetreated demineralized dentin. Quintessence Int. 40, 155161. Lansdown, A.B., 2006. Silver in health care: antimicrobial effects and safety in use. Curr. Probl. Dermatol. 33, 1734. Lewis, E.I., 1920. Inorganic Chemistry. Cambridge University Press. Li, R., Lo, E.C., Liu, B.Y., Wong, M.C., Chu, C.H., 2016. Randomized clinical trial on arresting dental root caries through silver diammine fluoride applications in communitydwelling elders. J. Dent. 51, 1520. Liu, B.Y., Lo, E.C., Chu, C.H., Lin, H.C., 2012a. Randomized trial on fluorides and sealants for fissure caries prevention. J. Dent. Res. 91, 753758. Liu, B.Y., Lo, E.C., Li, C.M., 2012b. Effect of silver and fluoride ions on enamel demineralization: a quantitative study using micro-computed tomography. Aust. Dent. J. 57, 6570. Llodra, J.C., Rodriguez, A., Ferrer, B., Menardia, V., Ramos, T., Morato, M., 2005. Efficacy of silver diamine fluoride for caries reduction in primary teeth and first permanent molars of schoolchildren: 36-month clinical trial. J. Dent. Res. 84, 721724. Lo, E.C., Chu, C.H., Lin, H.C., 2001. A community-based caries control program for preschool children using topical fluorides: 18-month results. J. Dent. Res. 80, 20712074.
Applications of silver diamine fluoride in management of dental caries 697 Lou, Y.L., Botelho, M.G., Darvell, B.W., 2011. Reaction of silver diamine [corrected] fluoride with hydroxyapatite and protein. J. Dent. 39, 612618. McDonald, S.P., Sheiham, A., 1994. A clinical comparison of non-traumatic methods of treating dental caries. Int. Dent. J. 44, 465470. Mei, M.L., Li, Q.L., Chu, C.H., Yiu, C.K., Lo, E.C., 2012. The inhibitory effects of silver diamine fluoride at different concentrations on matrix metalloproteinases. Dent. Mater. 28, 903908. Mei, M.L., Chu, C.H., Lo, E.C., Samaranayake, L.P., 2013a. Fluoride and silver concentrations of silver diammine fluoride solutions for dental use. Int. J. Paediatr. Dent. 23, 279285. Mei, M.L., Chu, C.H., Low, K.H., Che, C.M., Lo, E.C., 2013b. Caries arresting effect of silver diamine fluoride on dentine carious lesion with S. mutans and L. acidophilus dualspecies cariogenic biofilm. Med. Oral Patol. Oral Cir. Bucal. 18, e824e831. Mei, M.L., Ito, L., Cao, Y., Li, Q.L., Lo, E.C., Chu, C.H., 2013c. Inhibitory effect of silver diamine fluoride on dentine demineralisation and collagen degradation. J. Dent. 41, 809817. Mei, M.L., Li, Q.L., Chu, C.H., Lo, E.C., Samaranayake, L.P., 2013d. Antibacterial effects of silver diamine fluoride on multi-species cariogenic biofilm on caries. Ann. Clin. Microbiol. Antimicrob. 12, 4. Mei, M.L., Ito, L., Cao, Y., Li, Q.L., Chu, C.H., Lo, E.C., 2014a. The inhibitory effects of silver diamine fluorides on cysteine cathepsins. J. Dent. 42, 329335. Mei, M.L., Ito, L., Cao, Y., Lo, E.C., Li, Q.L., Chu, C.H., 2014b. An ex vivo study of arrested primary teeth caries with silver diamine fluoride therapy. J. Dent. 42, 395402. Mei, M.L., Lo, E.C., Chu, C.H., 2016a. Clinical use of silver diamine fluoride in dental treatment. Compend. Contin. Educ. Dent. 37, 9398. quiz100. Mei, M.L., Zhao, I.S., Ito, L., Lo, E.C., Chu, C.H., 2016b. Prevention of secondary caries by silver diamine fluoride. Int. Dent. J. 66, 7177. Mei, M.L., Nudelman, F., Marzec, B., Walker, J.M., Lo, E.C.M., Walls, A.W., et al., 2017. Formation of fluorohydroxyapatite with silver diamine fluoride. J. Dent. Res. 96, 11221128. Mei, M.L., Lo, E.C.M., Chu, C.H., 2018. Arresting dentine caries with silver diamine fluoride: what’s behind it? J. Dent. Res. 97, 751758. Milgrom, P., Horst, J.A., Ludwig, S., Rothen, M., Chaffee, B.W., Lyalina, S., et al., 2018. Topical silver diamine fluoride for dental caries arrest in preschool children: a randomized controlled trial and microbiological analysis of caries associated microbes and resistance gene expression. J. Dent. 68, 7278. Mohammed, N.R., Lynch, R.J., Anderson, P., 2014. Effects of fluoride concentration on enamel demineralization kinetics in vitro. J. Dent. 42, 613618. Monse, B., Heinrich-Weltzien, R., Mulder, J., Holmgren, C., van Palenstein Helderman, W. H., 2012. Caries preventive efficacy of silver diammine fluoride (SDF) and ART sealants in a school-based daily fluoride toothbrushing program in the Philippines. BMC Oral Health. 12, 52. Available from: https://doi.org/10.1186/1472-6831-12-52. Murdoch-Kinch, C.A., McLean, M.E., 2003. Minimally invasive dentistry. J. Am. Dent. Assoc. 134, 8795. Ngo, H., Knight, M., George, G., Sekiguchi, T., 2002. United States Patent. US 6,461,161 B1. Peng, J.J., Botelho, M.G., Matinlinna, J.P., 2012. Silver compounds used in dentistry for caries management: a review. J. Dent. 40, 531541.
698 Advanced Dental Biomaterials Punyanirun, K., Yospiboonwong, T., Kunapinun, T., Thanyasrisung, P., Trairatvorakul, C., 2018. Silver diamine fluoride remineralized artificial incipient caries in permanent teeth after bacterial pH-cycling in-vitro. J. Dent. 69, 5559. Rosenblatt, A., Stamford, T.C., Niederman, R., 2009. Silver diamine fluoride: a caries “silver-fluoride bullet”. J. Dent. Res. 88, 116125. Savas, S., Kucukyilmaz, E., Celik, E.U., Ates, M., 2015. Effects of different antibacterial agents on enamel in a biofilm caries model. J. Oral Sci. 57, 367372. Schwendicke, F., Gostemeyer, G., 2017. Cost-effectiveness of root caries preventive treatments. J. Dent. 56, 5864. SDI, 2016. Riva Star SDS. Version No: 4.1.1.1 ed. SDI Limited. Shah, S., Bhaskar, V., Venkataraghavan, K., Choudhary, P., Ganesh, M., Trivedi, K., 2013. Efficacy of silver diamine fluoride as an antibacterial as well as antiplaque agent compared to fluoride varnish and acidulated phosphate fluoride gel: an in vivo study. Indian J. Dent. Res. 24, 575581. Shah, S., Bhaskar, V., Venkatraghavan, K., Choudhary, P., Ganesh, M., Trivedi, K., 2014. Silver diamine fluoride: a review and current applications. J. Adv. Oral Res. 5, 2535. Suzuki, T., Nishida, M., Sobue, S., Moriwaki, Y., 1974. Effects of diammine silver fluoride on tooth enamel. J. Osaka. Univ. Dent. Sch. 14, 6172. Tan, H.P., Lo, E.C., Dyson, J.E., Luo, Y., Corbet, E.F., 2010. A randomized trial on root caries prevention in elders. J. Dent. Res. 89, 10861090. Targino, A.G., Flores, M.A., dos Santos Junior, V.E., de Godoy Bene Bezerra, F., de Luna freire, H., Galembeck, A., et al., 2014. An innovative approach to treating dental decay in children. A new anti-caries agent. J. Mater. Sci. Mater. Med. 25, 20412047. ten Cate, J.M., 2013. Contemporary perspective on the use of fluoride products in caries prevention. Br. Dent. J. 214, 161167. Tenuta, L.M., Cury, J.A., 2010. Fluoride: its role in dentistry. Braz. Oral Res. 24 (Suppl 1), 917. Thanatvarakorn, O., Islam, M.S., Nakashima, S., Sadr, A., Nikaido, T., Tagami, J., 2016. Effects of zinc fluoride on inhibiting dentin demineralization and collagen degradation in vitro: a comparison of various topical fluoride agents. Dent. Mater. J. 35, 769775. Thibodeau, E.A., Handelman, S.L., Marquis, R.E., 1978. Inhibition and killing of oral bacteria by silver ions generated with low intensity direct current. J. Dent. Res. 57, 922926. Vasquez, E., Zegarra, G., Chirinos, E., Castillo, J.L., Taves, D.R., Watson, G.E., et al., 2012. Short term serum pharmacokinetics of diammine silver fluoride after oral application. BMC Oral Health 12, 60. Vogel, G.L., 2011. Oral fluoride reservoirs and the prevention of dental caries. Monogr. Oral Sci. 22, 146157. Wakshlak, R.B., Pedahzur, R., Avnir, D., 2015. Antibacterial activity of silver-killed bacteria: the “zombies” effect. Sci. Rep. 5, 9555. Wilson, N.H.F., 2007. Minimally Invasive Dentistry: The Management of Caries. Quintessence Pub., Chicago, IL; London. Wright, J.T., White, A., 2017. Silver diamine fluoride: changing the caries management paradigm and potential societal impact. N. C. Med. J. 78, 394397. Yamaga, R., Nishino, M., Yoshida, S., Yokomizo, I., 1972. Diammine silver fluoride and its clinical application. J. Osaka. Univ. Dent. Sch. 12, 120. Yee, R., Holmgren, C., Mulder, J., Lama, D., Walker, D., van Palenstein Helderman, W., 2009. Efficacy of silver diamine fluoride for arresting caries treatment. J. Dent. Res. 88, 644647.
Applications of silver diamine fluoride in management of dental caries 699 Zhang, W., Mcgrath, C., Lo, E.C., Li, J.Y., 2013. Silver diamine fluoride and education to prevent and arrest root caries among community-dwelling elders. Caries Res. 47, 284290. Zhao, I.S., Gao, S.S., Hiraishi, N., Burrow, M.F., Duangthip, D., Mei, M.L., et al., 2017a. Mechanisms of silver diamine fluoride on arresting caries: a literature review. Int. Dent. J. 68, 6776. Zhao, I.S., Mei, M.L., Burrow, M.F., Lo, E.C., Chu, C.H., 2017b. Effect of silver diamine fluoride and potassium iodide treatment on secondary caries prevention and tooth discolouration in cervical glass ionomer cement restoration. Int. J. Mol. Sci. 18 (2), 340. Zhao, I.S., Mei, M.L., Li, Q.L., Lo, E.C.M., Chu, C.H., 2017c. Arresting simulated dentine caries with adjunctive application of silver nitrate solution and sodium fluoride varnish: an in vitro study. Int. Dent. J. 67, 206214. Zhi, Q.H., Lo, E.C., Lin, H.C., 2012. Randomized clinical trial on effectiveness of silver diamine fluoride and glass ionomer in arresting dentine caries in preschool children. J. Dent. 40, 962967.
Index Note: Page numbers followed by “f,” “t,” and “b” refer to figures, tables, and boxes, respectively. A Ablation, 532 Abrasion resistance, 50 51 Abrasives, 589 Academy of Dental Materials, 166 Accelerator, 45 46 Acetone, 330 Acetoxy silastic adhesives, 636 638, 638f Acetoxy silicone adhesives, 632, 638t Achromatic color, 8 9 Acid Blue 9. See Erioglaucine disodium Acid decomposable glasses, 178 Acid etching of dentin, 238 technique, 232 235 citric acid, 233 234 EDTA, 235 ferric chloride solution, 234 NaOCl, 235 phosphoric acid, 233 of titanium surfaces, 590 Acid orthophosphate, 378 Acid-functionalized monomers, 130 Acid-resistant ceramic, 67 Acidic functional monomers, 156 157 Acidulated phosphate fluoride (APF), 690 691 ACPS. See 3Acryloxypropyltrimethoxysilane (ACPS) ACPs. See Amorphous calcium phosphates (ACPs) Acrylic artificial teeth, 99 100 Acrylic denture base materials, 80 86 acrylic artificial teeth, 99 100 chemical structure and mechanism of polymerization, 82 85 commercial forms and composition gel form, 86 powder liquid form, 85 86 denture base material, 80 development, 80 86 and manufacturing technologies, 86 94 requirements, 81t denture lining materials, 94 99 Acrylic polymers, 79 80 Acrylic resins, 2, 82, 128, 617 618 3-Acryloxypropyltrimethoxysilane (ACPS), 76 Actinomyces naeslundii, 690 691 Activ GP, 219 Active surface, 569 570 ADA. See American Dental Association (ADA) Addition reaction, cross-linking by, 625 Additive manufacturing (AM), 519 520 Adhesion, 230 of fibers to polymer matrix, 310 tozirconia, 329 333 Adhesives adhesively bonded amalgam restoration, 114 advancement antibacterial properties, 242 245 bioactive properties, 245 247 classifications, 237 241 based on clinical steps, 240 241 based on interaction with smear layer, 241 first-generation adhesives, 237 composition, 231 232 bonding, 231 232 etchant, 231 primer, 231
702 Adhesives (Continued) dentin bonding, 241 242 dentistry, 319 320 development, 229 eighth-generation, 240 fifth-generation, 239 fourth-generation, 238 239 indications, 230 231 in restorations, 230 second-generation, 237 238 seventh-generation, 240 sixth-generation, 239 240 third-generation, 238 Admixed composition alloy, 112 113 Ag3PO4, 689 Agar, 202 Agarose, 653 AH-26 resin sealers, 280 Air abrasion, 330 AISI 316L steel, 545 Al2O3/SiO2 ratio, 179 180 Alcohol, 330 Alginate, 201 202, 649 650 Alkali glass. See Glass A Alkaline earth ions. See Alkaline ions Alkaline glass fillers, 130 131 Alkaline ions, 178 179 Alkyl silicate, 203 “All-in-one” adhesive system. See Seventhgeneration adhesives Allergic reaction, 25 26 Allografts, 353 354, 353t DFDBA, 354 fresh or frozen, 353 mineralized freeze-dried bone, 354 Alloplasts, 355 362 bioactive glasses, 360 361 calcium phosphates, 357 360, 358t calcium polyphosphate, 361 calcium sulfate, 362 glass ionomers, 361 362 magnesium-based biodegradable materials, 362 polymers, 357 Alloyed gold, 107 Alloys, 217 218 silver, 217 218 α β titanium (α β Ti), 549 α form, 39 Index α-hemihydrate, 39 α-titanium (αTi), 549 α-tricalcium phosphate (α-TCP), 358t, 359, 398 399 Alternative restorative treatment (ART), 189, 219 Alumina (Al2O3), 89 90, 117 118, 320, 322, 527 528, 589 airborne-particle abrasion, 330 ceramic brackets, 552 553 ceramics, 55 56 materials, 485 Alumina-based ceramics, 323 Alumina-toughened zirconia (ATZ), 320 Aluminosilicate polyacrylate (ASPA). See Glass ionomers Aluminum, 218 219 Aluminum oxide. See Alumina (Al2O3) Alveolar bone properties, 587 AM. See Additive manufacturing (AM) Amalgam. See Dental amalgam Amalgamation process, 109 Amalgambond, 234 Amalgomer, 186 American Dental Association (ADA), 238 Ammonium hexafluorosilicate, 690 691 Ammonium persulfate, 180 Amorphous, 55 Amorphous calcium phosphates (ACPs), 358t, 387, 388t, 396 398, 483 Anatomical dental duplicate capture, 518 Animal-derived biomaterials, 647 Anterior composite restoration, 148 Antheraea pernyi SF (APSF), 650 Antibacterial activity, 465 composites, 152 153, 152t nanoparticle modified endodontic sealer, 489 properties of adhesives, 242 245 Antimicrobial peptides, 3 APAFILL-G, 412 Apatites, 399 403 apatite-containing formulations, 401 402 formation, 456 APSF. See Antheraea pernyi SF (APSF) Arabic Gum, 39 40 Aramid fibers, 88, 133 Arestin, 483
Index Arginine, 245 Aromatic hydrocarbons, 82 Aromatic polyamide fibers, 88 ART. See Alternative restorative treatment (ART) Atlantis abutment, 529 Atraumatic restorative treatment. See Alternative restorative treatment (ART) Atraumatic restorative technique (ART), 677 Attachments in clear aligners, 569 571, 570f, 570t ATZ. See Alumina-toughened zirconia (ATZ) Austenitic NiTi wires (A-NiTi wires). See Superelastic NiTi Austenitic SS, 545 547 Australian stainless steel wires, 557 Autoclave, 77 Autografts, 352 353, 352t Autopolymerizing resin, 86 B BAG. See Bioactive glass (BG) Bakelite, 79 81 Barium, 144 Barium oxide (BaO), 130 131 Barium titanate (BaTiO3), 90 bcc lattice. See Body-centered cubic lattice (bcc lattice) BCP. See Biphasic calcium phosphate (BCP) BD. See Biodentine (BD) Beautifil from Shofu, 152 Begg light archwire technique, 557 Benzalconium chloride, 153, 242 243 Benzoylperoxide (BPO), 135 β-hemihydrate, 39 β-titanium (β-Ti), 549 alloys, 561 β-tricalcium phosphate (β-TCP), 358t, 359, 360f, 398 399, 415, 657 BFS. See Biaxial FS (BFS) BG. See Bioactive glass (BG) BHT. See Butylated hydroxytoluene (BHT) Biaxial FS (BFS), 187 188 BIC. See Bone implant contact (BIC) Bidirectional fibers, 308 309 Bio-glass. See Bioactive glass (BG) Bio-inks, 663 664 703 Bio-Oss, 354 355 Bioactive cement, 600 Bioactive glass (BG), 131, 246 247, 274, 360 361, 454, 657 658 clinical applications, 460 469 antibacterial activity, 465 bioactive glass scaffolds, 465 468 bone graft substitute, 461 bone regeneration, 461 462 coating of implants, 463 drug delivery system, 462 particle size of bioactive glasses, 468 469 role in minimal invasive dentistry, 465 use in toothpastes, 463 464 composition, 456 460 factors affecting apatite formation, 456 fibers, 29 future, 469 methods of synthesis, 460 scaffolds, 465 468 Bioactive/bioactivity, 28 29 of glasses, 454 456 mechanism of action, 455 solubility, 455 456 molecules, 654 nanoscale glass for root canal disinfection, 490 properties of adhesives, 245 247 remineralizing composites, 150 151 restorative materials, 661 662 BD, 662 MTA, 661 662 Bioaggregate, 262 Bioceramic(s), 380 381, 603, 655 658 BCP, 657 bioactive glasses, 657 658 bioceramic-based sealers, 284, 284t HA, 656 657 materials, 256, 262 263 onlays or crowns, 380 381 restorations, 380 381 tricalcium phosphate, 655 656 Biochemical binding to bone, 660 Biocompatibility, 2, 22 23, 23t, 25f, 118 119 of resin composites, 145 147 toxicology of mercury, 118 119 Biodegradability, 2
704 Biodegradable/biodegradation, 28 of implanted graft materials and bone formation, 363 365 macrophages respond to small fragments and particles, 365f materials, 644 metal scaffolds, 658 659 polymers, 469 Biodentine (BD), 262, 600, 662 Biofilm, 74 BioForce wires, 560 Bioglasses, 657 658, 661 Biological apatite, 378 Biological nanoceramics, 488 489 Biomaterials, 1, 106, 217 218, 644. See also Nanomaterials classification, 646 647 properties and implant surface characteristics of dental implants, 584 in tissue engineering and regeneration, 647 used in dentistry, 305 306 ceramics, 306 composites, 306 metals, 305 Biomedical engineering, 2 Biomedical science, 2 BiominF. See Fluoride-containing BG (BiominF) BioRoot RCS, 600 Biphasic calcium phosphate (BCP), 406, 657 Biphasic CaPO4 formulations, 404 Bis-GMA. See Bisphenol-A glycidyldimethacrylate (Bis-GMA) Bis-GMA/TEGDMA resin matrix, 130 131 Bis(acryloyloxymethyl)tricyclo[5. 2. 1. 02, 6]decane. See TCD-DI-HEA Bismethacryloyloxy ethyl phosphate (BMEP), 156 157 Bismuth oxide, 257 261 Bisphenol-A glycidyldimethacrylate (BisGMA), 128 130, 147, 277 Bitewing imaging technique, 510t Black staining of SDF-treated teeth, 677f Blackboard, 533 534 Blended learning, 533 534 Blue Elgiloy, 557 BMEP. See Bismethacryloyloxy ethyl phosphate (BMEP) Index BMPs. See Bone morphogenetic proteins (BMPs) BMSF. See Bombyx mori SF (BMSF) Body-centered cubic lattice (bcc lattice), 543 544 Bombyx mori SF (BMSF), 650 Bond strength evaluation, 71 73 Bonded amalgam restorations, 114 Bonding, 231 232 Bone augmentation procedures, 353 formation, 347 348 regeneration, 405, 461 462 replacement, 347 348 biodegradation of implanted graft materials and bone formation, 363 365 future of bone tissue graft materials, 365 366 graft materials biocompatibility and degradation products, 363 natural tissues and synthetic biomaterials, 350 362 rationale behind use of bone replacement graft materials, 348 349, 351t scaffolds, 458 supported guides, 526 Bone grafts, 347 348, 645 natural tissues and synthetic biomaterials, 350 362 allografts, 353 354, 353t alloplasts, 355 362 autografts, 352 353, 352t, 353t xenografts, 354 355 substitute, 461 Bone implant contact (BIC), 659 Bone morphogenetic proteins (BMPs), 349 BMP-2, 649 651 Bone tissue engineering, scaffolds for, 645 646 Bone tissue graft materials, future of, 365 366 Borate glass, 454 Borate-based BGs, 462 Borate-based bioactive glasses, 458 460 Borate-based bioglasses, 657 658 Borax, 205 206 Boron, 144
Index BOs. See Bridging oxygens (BOs) “Bottom-up” nanomanufacturing, 219 220 BPO. See Benzoylperoxide (BPO) Braided NiTi wires, 560 Braided wires, 557 Branched polymers, 620 Bridging oxygens (BOs), 178 180 Brilliant Blue FCF. See Erioglaucine disodium Brittleness, 19 Brushite. See Dicalcium phosphate dihydrate (DCPD) Bulk-fill composites, 157 165, 160t Bulk-fill RBCs, 138 Butylated hydroxytoluene (BHT), 290 292 C C-factor. See Cavity configuration factor (Cfactor) CAD CAM system. See Computer-aided design computer-aided manufacturing system (CAD CAM system) Calcination method, 39, 39t Calcium, 378, 390 391 Calcium fluoride (CaF2), 222 223 Calcium fluoroaluminosilicate glass powder, 190 Calcium hydroxide (Ca(OH)2), 271 272, 275f sealers, 278 279, 280t, 286t Calcium hydroxyapatite (CHAp), 590 Calcium orthophosphate (CaPO4), 378 379, 381, 382t, 383 385 CaPO4-containing endodontic and/or root canal sealers, 406, 406t ceramic brackets, 408f chewing gums, 386 389 classification according to dental specialties endodontics, 404 406 oral and maxillofacial surgery, 407 oral applications, 413 orthodontics, 407 409 periodontics, 411 413, 411f prosthodontics, 409 411 tissue integration of dental implant, 410f classification according to existing CaPO4 705 ACPs, 396 398 α-TCP and β-TCP, 398 399 apatites, 399 403 biphasic and multiphasic CaPO4 formulations, 404 DCPD and DCPA, 395 396 MCPM and MCPA, 395 OCP, 396 TTCP, 403 404 clinical applications in dentistry, 394 413, 394t for dental caries prevention and in dentifrices, 385 394 dentin hypersensitivity treatments, 392 394 information on current biomedical applications, 385 teeth remineralization, 390 392 tissue engineering approaches, 414 415 toothpastes, 386 389, 388t Calcium oxide (CaO), 63, 657 658 Calcium phosphate (CaP), 150, 357 360 CaP based bioceramics, 363 CaP coated titanium surfaces, 590 dicalcium phosphates, 359 360 hydroxyapatite, 357 359, 358t TCP, 359 Calcium phosphorous nanoparticles (Ca P nanoparticles), 485 Calcium polyphosphate (CPP), 361, 396 397, 650 Calcium silicate, 262 263 Calcium sulfate, 362 Calcium sulfate dehydrate, 44 Calcium sulfate hemihydrate, 38 40, 44, 44f Calcium-deficient hydroxyapatite (CDHA), 358t, 393 394, 399 403 Calcium-enriched mixture cement, 600 Calset, 155, 156f Calthane, 619 Camphorquinone (CQ), 136 free radical formation by light activation, 135f Candida albicans, 27 CaP. See Calcium phosphate (CaP) Ca P nanoparticles. See Calcium phosphorous nanoparticles (Ca P nanoparticles) Caphosol, 413
706 Capseal, 284t, 406, 406t Carbides, 133 134 Carbon, 543 544, 544t Carbon dioxide (CO2), 331 laser, 531 Carbon fiber, 307 308 carbon fiber reinforced carbon, 586 587 Carbon nanotubes (CNTs), 133 134, 602 Carbon-carbon double bonds (C 5 C), 139 Carbon-carbon single bonds (C C), 139 Carbonate hydroxyl apatite nanoparticles, 483 Carbonization, 532 Carbothermic smelting, 621 Carcinogenicity, 27 28 CARES. See Computer-Aided Restoration Service (CARES) Caries management, clinical effects of silver diamine fluoride applications on in adults, 683 684 in children, 677 683 in the elderly, 684 Cariogenic bacteria cariostatic effects of silver diamine fluoride on, 690 692 Carnauba, 9 11 Casearia sylvestris, 601 602 Casein phosphopeptide (CPP), 483 Casein phosphopeptide ACP (CPP ACP), 390 CAT. See Clear aligner therapy (CAT) Cavalite, 402 403 Cavity. See Dental caries Cavity configuration factor (C-factor), 142, 154 CBCT. See Cone beam computed tomography (CBCT) CCD. See Charged-coupled device (CCD) CCTP. See Collagen/CS/β-TCP (CCTP) CDHA. See Calcium-deficient hydroxyapatite (CDHA) Celluloid, 80 Cellulose nitrates, 79 80 Cellulose phthalate, 483 Cementitious materials, graphene to improve, 600 609 Cements, 2 Cementum, 304 Cephalometric projections, 511t Index Ceramic bonding, 65 74 bond strength evaluation, 71 73 fatigue, 73 74 mechanism, 65 71 chemical surface conditioning, 65 70 mechanical surface conditioning, 70 71 Ceramic brackets, 551 553, 551b. See also Metal brackets alumina, 552 553 polycrystalline zirconia, 552 Ceramic(s), 133 134, 306, 318, 586 587, 646 biological interaction, 74 77 physical parameters, 76 77 sterilization methods, 77 surface chemistry, 74 76 ceramic-based biomaterials, 647 ceramic resin ceramic system, 71 73 materials in dentistry glass ceramics, 56 61 oxide ceramics, 61 64 polymer-infiltrated ceramics, 64 65 scaffolds, 646 CEREC 3D, 529 Cerec system, 518 519 Cerium trioxide (Ce2O3), 63 Cetylpyridinium chloride, 153 CGIS. See Computer-guided implant surgery (CGIS) Chairside milling, 518 CHAp. See Calcium hydroxyapatite (CHAp) Charged-coupled device (CCD), 514 515 Chemical adhesion with tooth, 188 189 Chemical composition, 55 56 Chemical degradation reaction process, 458 Chemical structure and mechanism of polymerization, 82 85 Chemical surface conditioning hydrofluoric acid etching, 65 67 primer, 67 70 Chemical vapor deposition (CVD), 599 600 CVD-grown graphene, 607f Chemical-resistant glass. See Glass C Chemically activated resins, 97 98 Chewing gums, 386 389 Chitosan (CS), 278, 483 484, 647, 651 653 CS alginate hybrid scaffolds, 649
Index Chlor-XTRA, 270t Chlorhexidine (CHX), 153, 201 202, 242 243, 266 267, 272 273, 684, 690 691 Chlorinated polyethylene, 619 Chloropercha, 286t Chroma. See Saturation of color Chromium, 543 544, 544t CHX. See Chlorhexidine (CHX) CHX-Plus, 270t Citric acid, 233 234 CLD. See Cross-link density (CLD) Cleaning agents, 330 Clear aligner therapy (CAT), 562 Clear aligners, 562 572 attachments, 569 571, 570f, 570t cytotoxicity, 571 572 forces of thermoplastic aligners, 565 567 material composition, 562 565 mechanical properties, 567 569 thermoforming process, 565 Clearfil Bond System F, 237 238 Clearfil Protect Bond, 244 Clinical durability of resin composites, 147 148 Clinpro white varnish, 399 Closed systems, 518 519 CMO SAPS. See Complementary metal oxide semiconductor active pixel sensor (CMO SAPS) CNIS. See Computer-navigated implant surgery (CNIS) CNTs. See Carbon nanotubes (CNTs) Coating of implants, 463 Cobalt-based wear-resistant alloy, 548, 548t Cobalt chromium brackets (CoCr brackets), 548 Cobalt chromium-based archwires, 557 558, 558t coDiagnostiX software, 523 Coe-Comfort materials, 633 635, 634f Coe-Soft materials, 633 635, 634f Coefficients of thermal expansion (CTE), 14, 57 Cold-cured resins, 85 86, 95, 96t acrylic resins, 84 Collagen, 490, 606, 647 648 collagen-bioactive ceramic composites, 648 sponge-form scaffolds, 606 607 707 Collagen/CS/β-TCP (CCTP), 648 Colloidal dispersion of rGO-coated hydroxyapatite, 607 608 Color, 8 9 Coloring agents, 635 638 acetoxy silastic adhesives, 636 638 colored flocking, 635 636 extrinsic colors, 636 intrinsic stains, 636 Commercially pure titanium (cpTi), 585, 659 grade 2 CP-Ti, 549 grade 4 CP-Ti, 549, 585 587 Complementary metal oxide semiconductor active pixel sensor (CMO SAPS), 514 Compomers, 130, 153 154 Composite resin (CR), 683 Composites, 306 CS scaffolds, 653 material, 55, 137 138, 220 preheating, 155 resins, 290 292, 306 Compressive load, 19 Compressive strength (CS), 16, 49, 114 115, 180 181, 185 186 Computed tomography (CT), 510, 511t, 515, 520, 663 664 Computer-aided design computer-aided manufacturing system (CAD CAM system), 93 94, 517 521 abutments in implant dentistry, 528 block, 64 65 CAD/CAM blocks, 319 Chairside milling, 518 composites, 159, 165 custom implant abutments, 529 530 in implant restorative dentistry, 527 528 laboratory and industrial milling, 518 519 machining of restorations, 519 three-dimensional printing, 519 521 versions, 58 60, 60f Computer-Aided Restoration Service (CARES), 529 Computer-assisted technologies, 51 Computer-guided implant surgery (CGIS), 522, 524 Computer-navigated implant surgery (CNIS), 522, 524 527
708 Computer-supported implant dentistry, 521 530 CAD CAM abutments in implant dentistry, 528 custom implant abutments, 529 530 system in implant restorative dentistry, 527 528 CGIS, 524 CNIS, 524 527 materials, 529 prosthetic abutments, 528 recent advances in implant technologies, 522 523 three-dimensional printing in implant dentistry, 521 522 Condensation polymerization, 201 reaction, cross-linking by, 624 625 silicone, 203 Cone beam computed tomography (CBCT), 509 510, 511t, 512, 515, 520 Conventional cements, 329 Conventional endodontic treatment, 490 Conventional glass ionomers, 218 219 Conventional plaster casting materials, 205 206 Conventional silicate-based glasses, 460 461 Conventional waxes, 204 205 Cooling process, 57 Copper, 543 544, 544t NiTi, 560 Core obturation materials Gutta-percha, 275 277 resilon, 277 root canal sealers, 277 289 silver points, 274 275, 275t Coronal caries management in adults, 683 684 in children, 677 683 Corrosion, 116 117 marginal sealing, 116 117 resistance, 545 Coupling agents, 332 333 CPP. See Calcium polyphosphate (CPP); Casein phosphopeptide (CPP) CPP ACP. See Casein phosphopeptide ACP (CPP ACP) cpTi. See Commercially pure titanium (cpTi) Index CQ. See Camphorquinone (CQ) CQ/TA, 136 Creep, 115 116, 568 diffusion, 61 62 resistance, 568 Cross-link density (CLD), 201 202 Cross-linking by addition reaction, 625 collagen scaffolds, 647 648 by condensation reaction, 624 625 materials, 86 polymers, 620 of polymers, 624 Crystalline HA, 488 phase, 57 structure, 63 theory, 44 CS. See Chitosan (CS); Compressive strength (CS) CT. See Computed tomography (CT) CTE. See Coefficients of thermal expansion (CTE) Cumulative fluoride [F]c release, 184 Curing, 175 176 mechanisms of acrylic denture base materials cold-cured acrylic resins, 84 heat-cured acrylic resins, 83 microwave-cured acrylic resins, 84 85 pour-type denture resins, 85 visible light cured resins, 84 resin composites classification according to, 138 Cushioning effect, 94, 96 Custom-made trays, 198 199 CVD. See Chemical vapor deposition (CVD) Cysteine cathepins, 692 Cytotoxicity, 24 25 tests, 26 D Daro water-based skin adhesive, 638, 639f DC. See Degree of conversion (DC) DCPA. See Dicalcium phosphate anhydrous (DCPA) DCPD. See Dicalcium phosphate dehydrate (DCPD); Dicyclopentadiene (DCPD) Deformation resistance, 167t
Index Degree of conversion (DC), 136, 139 140 cavity location effect, 140 light-curing duration effect, 140 system effect, 139 140 tip distance effect from RBC surface, 140 resin increment thickness effect, 139 resin shade effect, 139 Demineralization process, 378, 378f Demineralized freeze-dried bone allograft (DFDBA), 354 Denatured collagen, 647 Densely sintered aluminum-oxide ceramic, 61 62 Dental applications of nanobiomaterials, 481 492 attrition process, 380 CaPO4 classification according to dental specialties, 404 413 cements, 176 zinc oxide eugenol, 15 ceramics, 321 322 composites, 108 109 enamel, 378 erosion, 379 380 filling biomaterials amalgam, 109 111 gold fillings, 107 108 fluorosis, 27 FRCs, 311 gypsum and investments dies and models produced from digital data, 51 52 production of calcium sulfate hemihydrate, 38 40 properties of gypsum products, 38 setting and manipulation characteristics of gypsum products, 42 47 setting expansion hygroscopic setting expansion, 47 51 types of gypsum products, 40 41 implantology, 410 411 impression materials, 3 IMs for dental prosthesis characteristics, 207 211 elastic, 200 204 inelastic, 204 206 709 materials, 12, 217 218 prosthetics. See Prosthodontics public health, 383 385 pulp, 304 resin composite. See Resin composites restorations, 8 silane products, 134 stone. See α form stone Type V, 41, 42f Dental abrasion process, 380 Dental adhesives, 27 28, 237, 240 Dental amalgam, 2, 106, 109 111, 217 218, 290, 305, 492 493 amalgam bonding, 113 114 biocompatibility, 118 119 bonding bonded amalgam restorations, 114 nonbonded amalgam restorations, 113 114 nonbonded vs. adhesively bonded amalgam restorations, 114 composition, 112 113 high-copper dental amalgam, 112 113 low-copper dental amalgam, 112 dental filling biomaterials, 107 111 dimensional change, 117 failure mode, 118 hardness, 117 118 history, 111 112 material properties compressive and tensile strength, 114 115 creep, 115 116 tarnish and corrosion, 116 117 Young’s modulus, 118 Dental biomaterials, 1, 8 bioactivity, 28 29 biodegradation, 28 biological properties biocompatibility, 22 23 usage tests, 24 in vitro testing, 23 in vivo testing, 24 carcinogenicity, 27 28 cytotoxicity tests, 26 electrical conductivity and resistivity, 15 fluoride and caries, 26 27 foreign body reaction, 29 30 limitation of mechanical testing methods, 22
710 Dental biomaterials (Continued) mechanical properties and characterization methods, 16 22, 17f optical properties, 8 9 osseointegration, 29 osteoinduction, 29 thermal properties, 9 14 coefficient of thermal expansion, 14 heat of fusion, 11 specific heat, 13 temperature, 9 thermal conductivity, 12 thermal diffusivity, 13 14 transition temperatures, 9 11 toxicity and cytotoxicity, 24 25 viscosity, 14 Dental caries, 26 27, 107, 377 378, 379f, 400 401 CaPO4 in, 385 394 detection, 516 location effect, 140 preparation, 229 Dental implants, 581 582 biomaterial properties and implant surface characteristics, 584 biomechanical properties, 584 585 comparison of different hard tissue with, 584t functionally graded/hierarchical dental implant surfaces, 590 materials, 586 588 alveolar bone properties, 587 influence of oral health and systemic disease on implant survival, 587 588 modification, 588 590 of PEEK dental implants, 590 592 of zirconia implants, 592 nanomaterials for, 484 486 osseointegration, 583 584 surface properties, 585 586 titanium screw in dog mandible, 582f Dental mineral, cariostatic effects of silver diamine fluoride on, 685 689 DentalSlice software, 523 Dentifrices, CaPO4 in, 385 394 Dentin, 230, 303 304 bonding, 241 242 hypersensitivity treatments, 392 394 Index Dentine collagen, cariostatic effects of silver diamine fluoride on, 692 Dentistry, 2, 106, 217 219, 383, 455 456. See also Computer-supported implant dentistry; Digital dentistry biomaterials used in, 305 306 clinical applications of CaPO4 in, 394 413 GICs in, 219 zirconia in, 222, 324 325, 325t Denture base material, 79 80 development, 80 86 and manufacturing technologies, 86 94 fiber-reinforced resins, 87 89 hybrid reinforcement, 91 hypoallergenic resins, 91 novel technologies in manufacturing removable denture base, 93 94 particulate-reinforced resins, 89 91 rubber-reinforced resins, 87 thermoplastic resins, 92 93 requirements, 81t Denture lining materials, 94 99 clinical indication, 94 95 hard relining, 95 soft relining, 96 98 tissue conditioners, 98 99 DFDBA. See Demineralized freeze-dried bone allograft (DFDBA) Di-organo-polysiloxane, 625 Diaket, 292 Diametral tensile strength, 20 DiaRoot. See Bioaggregate Dibutyl phthalate, 86 Dicalcium phosphate anhydrous (DCPA), 358t, 387, 388t, 395 396 Dicalcium phosphate dehydrate (DCPD), 150 151, 358t, 359 360, 360f, 387, 388t, 395 396 Dicalcium phosphates, 359 360 Dicyclopentadiene (DCPD), 154 Die stone. See Modified α-hemihydrate Dies, 37 38 and models produced from digital data, 51 52 Differential thermal analysis, 9 11 Diffusion creep, 61 62 Digital camera and imaging systems, 517 Digital data, dies and models produced from, 51 52
Index Digital dentistry, 508. See also Dentistry CAD CAM system, 517 521 caries detection, 516 computer-supported implant dentistry, 521 530 digital radiography and magnetic resonance imaging, 508 515 lasers and dental applications, 530 533 photography and shade selection, 516 517 technology and dental education, 533 534 Digital imaging, 511t, 513t Digital radiography, 508 515 clinical applications, 509 512 intraoral, extraoral, including cone beam CT, 508 509 limitations, 512 515 Digital simulators, 534 Digital X-ray systems, 514 Dihydrate calcium sulfate, 38 39 Dilatant fluid, 14 Diluents monomers, 141 Dimer acid, 150 dimer acid based dimethacrylate monomer, 150 Dimethacrylates, 128 Dimethylaminododecyl methacrylate (DMADDM), 243 244 Dimethyldichlorosiloxane, 622 Diode laser, 532 Dipentaerythritol Penta acrylate phosphate (PENTA), 130 Diphenyl (2,4,6-trimethylbenzoyl) phosphine oxide (TPO), 136 Direct composite resins, 138 restoration, 154 Direct contact tests, 23 Direct dental materials, 217 218 restorations, 110 Direct metal layer sintering (DMLS), 520 522 Direct restoration, 138, 166 171 Direct restorative dental materials, 3 Dismetral compression test, 20 DMADDM. See Dimethylaminododecyl methacrylate (DMADDM) 711 DMAE-CB. See Methacryloxylethylcetyldimethyl ammonium chloride (DMAE-CB) DMAHDM. See 2-Methacryloyloxyethyl phosphorylcholine, dimethylaminohexadecyl methacrylate (DMAHDM) DMFT (number of decayed, missing, and filled teeth due to caries), 387 DMLS. See Direct metal layer sintering (DMLS) DMP-1, 606 Dominant wavelengths, 8 9 Double-walled halloysite nanotubes, 133 134 Drug delivery system, 462 Dry skulls, 512 Drynaria fortunei, 355 DSPP, 606 Dual-cured composite, 136 Ductility, 19 Duplex stainless steel, 547 DX-511, 143, 149 E E-beam deposition. See Electron-beam deposition (E-beam deposition) E-glass fibers, 133, 586 587 E-learning tools, 533 534 Early childhood caries (ECC), 677 Easy Guide software, 523 Edge adhesives, 639 EDTA. See Ethylenediamine tetra-acetic acid (EDTA) Education materials, 533 534 Eicosapentaenoic acid, 586 Eighth-generation adhesives, 240 Elastic cavity wall concept, 143, 155 Elastic impression materials. See Inelastic impression materials agar, 202 alginate, 201 202 polyethers, 200 polysulfide, 201 silicones, 202 204 Elastic limit, 18 Elastic modulus, 18 19, 567 Elastic region, 17 19 Elasticity, 305
712 Elastin, 653 Elastomeric materials, 14, 198 Electric glass. See Glass E Electrical conductivity and resistivity, 15 Electrical fibers, 133 Electrically assisted enhanced remineralization, 392 Electroforming, 107 Electron-beam deposition (E-beam deposition), 590 592 Electrospinning process, 1 2, 491 492 Electrospun nanomaterials, 491 492, 493t Elemental magnesium, 362 Elgiloy. See Cobalt chromium-based archwires Enamel, 230, 303, 393 demineralization, 379 etching technique, 231 236 acid etching, 232 235 laser etching, 235 236 self-etching, 236 Endo-CPM, 600 Sealer, 284t EndoBinder, 600 Endocem, 600 Endodontic(s), 256, 312, 383 385, 404 406 imaging, 512 materials intracanal medicaments, 271 274 perforation repair materials, 292 293, 293b as root canal irrigants, 263 270 root canal obturation materials, 274 289 root-end filling materials, 290 292 in vital pulp therapy, 256 263 nanobiomaterials, 488 490 regeneration, 490 EndoSeal MTA, 600 EndoSequence, 262 263, 600 601 EndoSequence BC Sealer, 395, 488 489 Endosequence root repair material (ERRM), 262 Endosequence root repair putty (ERRP), 262 Energy dispersive spectroscopy, 28 29 Enolase, 691 Epigallocatechin-3-gallate. See Green tea extract Index Epithane-3, 619 Epoxy resin, 50, 128 sealers, 280 Equia system, 487 488 Erbium, chromium: yttrium scandium gallium garnet radiation (Er, Cr:YSGG radiation), 235 236 Erbium laser, 531 532 Erosive mechanism, 184 ERRM. See Endosequence root repair material (ERRM) ERRP. See Endosequence root repair putty (ERRP) Essix C 1 , 569 Esters, 82 Esthet-X (Dentsply Caulk, United States), 130 131 Esthetic(s), 188 orthodontic brackets, 325 restoration, 144 Etch-and-rinse technique, 233 Etchant, 231 Etching, 231 232 Ethyl alcohol, 203 Ethylene glycol dimethacrylate, 83 Ethylenediamine tetra-acetic acid (EDTA), 235, 265 266 Etkon, 529 “Evaporable” water. See “Loosely bound” water Everest, 529 Exceed-30 aligner material, 564 565 Exceed-40 aligner material, 564 565 Excitation purity, 8 9 Extraoral digital imaging, 509 510, 511t MRI, 509 prostheses, 616 radiography, 508 509 Extrinsic colors, 636 F F-BAG. See Fluoride-based BAG (F-BAG) Fabrication of dental implants, 521 522 Face-centered cubic lattice (fcc lattice), 543 544 Failure analysis of zirconia restorations, 333 335
Index Failure mode, 118 Fatigue, 73 74 fcc lattice. See Face-centered cubic lattice (fcc lattice) FDA. See US Food and Drug Administration (FDA) FDBA. See Freeze-dried bone allograft (FDBA) Fe-phosphorous alloys, 658 Feldspathic ceramic, 318, 322 323 Feldspathic porcelain, 56 58 Ferric chloride solution, 234 Ferrite, 545 steel, 547 FGMs. See Functionally graded materials (FGMs) FHA. See Fluorohydroxylapatite (FHA) Fiber distribution, 308 length, 309 orientation, 136, 308 309, 309f position and placement, 89 quantity, 308 types, 306 Fiber-reinforced composites (FRCs), 3, 133 134, 136, 302, 302f, 306 311 carbon, 307 308 clinical applications, 311 312 endodontics, 312 implants, 311 orthodontic, 312 periodontal, 312 prosthodontics, 312 tooth restoration, 311 glass, 307 influencing factors on mechanical properties adhesion of fibers to polymer matrix, 310 impregnation of fibers with polymer matrix, 310 polymerization shrinkage, 311 water sorption, 310 polyethlene, 308 Fiber-reinforced resins, 87 89 position and placement of fibers, 89 Fibrin, 648 649 Fifth-generation adhesives, 239 Fillers, 130 134, 306 713 fiber-reinforced composites, 133 134 nanofilled resin composite, 132 133 prepolymerized filler particles, 132 size and loading, 131 132 Filtek Supreme (3M ESPE), 132 Fineness, 46 Finite element stress analysis, 155 First Class, 533 534 First generation adhesives, 237 M-NiTi. See Stabilized martensitic NiTi of Y-TZP, 63 64 Fixed dental prosthesis (FPD), 318 Flat-panel detectors, 515 Flexural strength (FS), 16, 186 188 Flowable bulk-fill materials, 158 Flowable composites, 137 139 Fluorapatite, 220 222, 385 386, 399 403, 454, 478 480, 689 Fluorescent agents, 136 Fluoride, 26 27 anticariogenic potential, 152 concentrations, in SDF products, 691 fluxes, 179 180 toxicity, 27 Fluoride release from fillers, 152 from GICs, 183 185 affecting factors, 184 185 chemical adhesion with tooth, 188 189 compressive strength, 185 186 esthetics, 188 flexural strength, 186 188 GICs in ART, 189 mechanism, 184 moisture sensitivity of GICs, 189 nanoapatite-filled glass ionomers, 189 190 RMGICs, 190 191 source of fluoride, 183 184 thermo-cured glass ionomers, 190 and recharge of compomer, 153 154 restorative materials, 218 Fluoride-based BAG (F-BAG), 247 Fluoride-containing BG (BiominF), 464, 465t Fluoride-modified dental implants, 590 Fluorine ions, 242 243 Fluorine-containing glass ceramics, 58
714 Fluormica, 58 Fluoroaluminosilicates, 178, 179f components, 179t glass, 218 Fluorocore composite, 114 115 Fluorohydroxyapatite (FHA), 400, 689 Fluorophlogopite, 58, 59t Food grade silicones, 623 Force, 16 17 Foreign body reaction, 29 30 Formability, 549, 554 555 Formaldehyde, 27 28 Four-point flexural bond strength, 73 Fourth-generation adhesives, 238 239 FPD. See Fixed dental prosthesis (FPD) Fractographic analysis of zirconia restorations, 333 335 Fracture resistance, 167t stress, 19 toughness, 467 468 FRCs. See Fiber-reinforced composites (FRCs) Free-radical addition polymerization, 82 83 Freeze-dried bone allograft (FDBA), 354 Fresh allografts, 353 FRIACOM (computer-assisted analysis device), 514 Frozen allografts, 353 FS. See Flexural strength (FS) FTIR analysis, 180 182 Functional tissue engineered products, 1 Functionally graded materials (FGMs), 590 Functionally graded/hierarchical dental implant surfaces, 590 G GAG. See Glycosaminoglycan (GAG) Gamma waves, 77 Gel form, 86 Gel point, 142 Gelation phase in GICs, 182 Generex B ProRoot Endo Sealer, 284t Gentamicin sulfate, 462 Geometric Studio, 518 GFRCs, 308 Gibbs free energy, 75 GICs. See Glass-ionomer cements (GICs) Gilmore test for final setting time, 47 Index Glass, 318 fibers, 133, 307 particles, 218 219 polyalkenoate cements, 175 176 powder decomposition, 182 transition temperature, 11, 455 Glass A, 307 Glass C, 307 Glass Carbomer by GCP Dental, Netherlands, 189 190 Glass ceramics, 55 56 feldspathic porcelain, 56 58 fluorine-containing glass ceramics, 58 leucite-reinforced porcelain, 58 lithium disilicate, 58 61 Glass E, 307 Glass ionomers, 177, 177f, 189 190, 361 362 conventional, 218 219 nanoapatite-filled, 189 190 as “nondental” cement, 191 sealers, 280, 286t thermo-cured, 190 Glass ionomer cement (GIC), 677 Glass R, 307 Glass S, 307 Glass-ceramics, 306, 319 321 blocks, 527 528 Glass-infiltrated aluminum-oxide ceramic, 61 Glass-infiltrated ceramics, 321 322 Glass-infiltrated zirconia development, 337 338 Glass-ionomer cements (GICs), 2, 175 176, 217 219, 290 292, 306, 487 applications, 219 in ART, 189 ASPA, 177, 177f chemistry of setting reaction, 181 183 components, 178 181 composition and nature of acid component, 180 181 of glass component, 178 180 decomposition of glass powder, 182 development, 176 178 fluoride release from, 183 185 gelation phase, 182 maturation phase, 183 moisture sensitivity of, 189
Index nanomodifications of GIC powders, 219 224 setting reaction in, 181 183 water, 181 Glossary of Prosthodontics Treatment (GPT), 616 GLUMA, 290 292 Glycerol phosphate-dimethacrylate (GPDM), 130, 156 157, 158f Glycosaminoglycan (GAG), 648 GNS. See Graphene nanosheets (GNS) GO. See Graphene oxide (GO) GO chitosan hydroxyapatite (GO CS HA), 603 Gold, 107, 305 advantages and disadvantages of gold wires, 556t alloys, 555 556 fillings, 107 108, 108f gold copper alloy system, 15 plated brackets, 550 GP. See Gutta-percha (GP) GPDM. See Glycerol phosphatedimethacrylate (GPDM) GPT. See Glossary of Prosthodontics Treatment (GPT) Grain boundaries sliding, 61 62 Grandio (Voco, Germany), 130 131 Graphene and derivatives, 599 600 films, 599 to improve cementitious materials, 600 609 nanofillers, 609 Graphene nanosheets (GNS), 603, 604f Graphene oxide (GO), 599 600, 605f GO-modified scaffolds, 606 607 GO-modified β-tricalcium phosphate scaffold, 607 608 Grassmann’s law, 8 9 Gray unit (Gy unit), 513 Green Elgiloy, 558 Green machining, 519 Green tea extract, 153 Grit-blasting, 70 71, 337, 589 GTR membranes. See Guided tissue regeneration membranes (GTR membranes) Guided surgery, 526 715 Guided tissue regeneration membranes (GTR membranes), 483 484 Gusuibu, 355 Gutta-percha (GP), 219, 275 277, 276t GuttaFlow Bioseal, 600 GuttaFlow sealer, 489 Gy unit. See Gray unit (Gy unit) Gypsum products, 2, 37 38 H HA. See Hyaluronic acid (HA); Hydroxyapatite (HA) Hafnia (HfO2), 75 76 Hap. See Hydroxyapatite (HA) Hard relining, 95 Hardness, 117 118, 305 HCA. See Hydroxy-carbonate apatite (HCA) HCl. See Hydrochloric acid (HCl) Heat of fusion, 11, 12t heat-activated resins, 97 98 heat-cured acrylic resins, 83 heat-cured acrylics, 95 HEBP. See 1-Hydroxyethylidene-1,1bisphosphonate (HEBP) HEMA. See Hydroxyethyl methacrylate (HEMA) Hemihydrate crystalline species, 39 Herculite XRV (Kerr, United States), 130 131 Hexafluoroisopropanol-based silk (HFIPbased silk), 651 HF. See Hydrofluoric acid (HF) HFIP-based silk. See Hexafluoroisopropanolbased silk (HFIP-based silk) High molecular-weight resin, 149 High-copper amalgam, 116 117 dental amalgam, 112 113 High-strength core ceramics, 55 56 High-temperature vulcanization silicone (HTV silicone), 624 625 Highest occupied molecular orbitals (HOMO), 75, 76f HL. See Hybrid layer (HL) HOMO. See Highest occupied molecular orbitals (HOMO) Homogeneous conductor’s resistance, 15 Hot isostatic pressing, 61 62
716 Hot pressing, 61 62 HTR Synthetic Bone, 357 HTV silicone. See High-temperature vulcanization silicone (HTV silicone) “Hue”, 8 9 Human tissue replacements, 1 Human-derived biomaterials, 647 Hyaline, 407 408 Hyaluronate. See Hyaluronic acid (HA) Hyaluronic acid (HA), 651 Hybrid acrylic, 83 ceramic, 159 165 reinforcement, 91 Hybrid layer (HL), 231 232 Hydration, 47 Hydrochloric acid (HCl), 590, 621 622 Hydrocolloids, 198 Hydrofluoric acid (HF), 65 66, 332, 590 etching, 65 67, 66f, 67t Hydrogen peroxide (H2O2), 269 Hydrolyzable alkyl group, 134 Hydrophilic elastomeric IMs, 202 203 IMs, 208 PEG hydrogels, 654 Hydrophilicity, 585 586 Hydroxy-carbonate apatite (HCA), 454 Hydroxyapatite (HA), 28, 220 222, 230, 347 348, 357 359, 358t, 360f, 386 387, 388t, 399 403, 454 455, 478 480, 646, 654 657 fillers, 91 HA-containing toothpaste, 393 394 HA/GNS composites, 607 608 nanocrystals, 485 structure, 394 Hydroxyethyl methacrylate (HEMA), 156 157, 190 1-Hydroxyethylidene-1,1-bisphosphonate (HEBP), 269, 270t Hydroxyl-free radicals (HO2), 269 Hygroscopic setting expansion, 47 51, 47t, 48f compressive strength, 49 dimensional stability, 51 mixing technique, 42 43 pouring the impression, 43 properties, 38 Index reproduction, 49 setting processes, 44 46 fineness, 46 modifying agents, 45 46 effect of pH, 46 rate of setting reaction, 44 spatulation, 45 stages, 44 temperature, 45 W/P ratio, 45 setting time final setting time, 46 47 gilmore test for final setting time, 47 initial setting time, 46 surface hardness and abrasion resistance, 50 51 tensile strength, 50 types, 40 41, 41f Hypoallergenic resins, 91 Hypochlorous acid (HOCl), 231 232 Hypoclean A, 270t Hypoclean B, 270t I IA. See Itaconic acid (IA) ICADS. See International Caries Assessment and Detection System (ICADS) ICP, 183 ICs. See Impression compounds (ICs) IGI. See Image-guided implantology (IGI) IIT. See Image intensifier tubes (IIT) IKI. See Iodine potassium iodide (IKI) Image intensifier tubes (IIT), 515 Image-guided implantology (IGI), 525 Imidazole-modified CS, 652 Immobilization, 485 486 Immunomodulatory strategies, 365 366 Implant3D Media Lab Software, 523 ImplantMaster TM software, 523 Implants coating, 463 failure, 583 FRCs in, 311 grade silicones, 622 technologies, recent advances in, 522 523 Impregnation of fibers with polymer matrix, 310 Impression compounds (ICs), 205
Index Impression materials (IMs), 198 comparison of various types of dental impression properties, 210t dimensional accuracy/dimensional stability, 207 208 elastic, 200 204 recovery/flexibility, 208 209 inelastic, 204 206 mechanical properties, 209 miscellaneous, 209 211 wettability, 208 Impression plasters (IPs), 205 206 Impression wax, 204 205 IMs. See Impression materials (IMs) In vitro cytotoxicity, 26 In vitro testing, 23 In vivo study, 660 In vivo testing, 24 Incremental layering technique, 143 Indirect composites, 138, 159 165, 163t Indirect contact tests, 23 Indirect dental materials, 217 218 Indirect tensile test, 20 Industrial grade silicones, 623 Industrial milling, 518 519 Inelastic impression materials. See Elastic impression materials ICs, 205 impression wax, 204 205 IPs, 205 206 metallic oxide pastes, 206 Infiltrated ceramics, 306 Inflammatory cells, 29 30 Infrared (IR) analysis, 183 spectroscopy, 457 458 Initiator accelerator system, 134 136 dual-cured composite, 136 light-cured composites, 135 136 self-cured composites, 135 Inkjet 3D printers, 663 664 Inorganic fillers, 130 particles, 486 Inorganic oxides, 136 Inorganic polyP, 361 Inox steel. See Stainless steel brackets (SS brackets) Intermediate restorative material (IRM), 290 292, 291t 717 International Association for Dental Research, 189, 386 International Caries Assessment and Detection System (ICADS), 516 International Standards Organization (ISO), 276 Intracanal medicaments, 271 274, 272t bioactive glass, 274 calcium hydroxide, 271 272, 275f chlorhexidine, 272 273 Ledermix, 273 rationale of usage, 271b triple antibiotics pastes, 273 Intraoral capture, 518 Intraoral digital imaging, 509 510, 510t Intraoral digital radiography, 514 515 Intraoral prostheses, 616 Intraoral radiography, 508 509 Intrinsic stains, 636, 637f Invisalign appliance by Align Technology, 562, 564 565 trays, 569 Iodine potassium iodide (IKI), 265, 269 Ion-delivering agents, 385 386 Ionomers, 220 222 salt hydrogel, 190 191 IPs. See Impression plasters (IPs) IRM. See Intermediate restorative material (IRM) Iron, 543 544 Iron oxide, 136 iRoot, 600 601 iRoot SP, 284t ISO. See International Standards Organization (ISO) Isocyanates, 571 chemical reaction for synthesis of polyurethanes from, 571f Itaconic acid (IA), 218 K Ketac, 487 Ketones, 82 Kevlar fibers, 88 L Laboratory milling, 518 519 Lactobacillus, 378
718 Lactobacillus acidophilus, 690 691 Lactose-modified CS, 649 LASER. See Light amplification by the stimulated emission of radiation (LASER) Laser etching technique, 235 236. See also Acid etching technique Lasers, 71, 331 irradiations, 235 236, 331 Lava, 529 LCUs. See Light control units (LCUs) LED. See Light-emitting diode (LED) Ledermix, 273 Leucite (K2O Al2O3 4SiO2), 322 leucite-based ceramics, 323 leucite-reinforced ceramics, 319 leucite-reinforced porcelain, 58 Light amplification, 71 fluorescence technology, 516 Light amplification by the stimulated emission of radiation (LASER), 531 532 and dental applications, 530 533 in dentistry, 531 light, 532 mechanism of laser action, 532 533 Light control units (LCUs), 136 Light-curing composites, 135 136 photoinitiators, 135 136 duration effect, 140 modes, 142 143 resins, 146 system effect, 139 140 tip distance effect from RBC surface, 140 Light-emitting diode (LED), 140 Linear polymers, 620 Lining material use, 155 156 Lipopolysaccharides (LPS), 266 267 Liquid, 86 Liquid silicone rubber (LSR), 627, 627t Lithium disilicate (Li2O 2SiO2), 58 61, 319, 322 lithium disilicate based ceramics, 323 lithium disilicate ceramic blocks, 527 528 reinforced ceramics, 319 Lithography, 219 220 Index LMSRT (computer-assisted analysis device), 514 Load deflection rate, 554 Local toxicity, 24 Long-term fluoride release, 184 “Loosely bound” water, 181 Low elastic modulus liners, 143 Low molecular-weight compounds, 128 130 Low-and medium-fusing dental porcelain frits, 58 Low-copper dental amalgam, 112 Low-shrinkage composites, 148 150 Dimer acid based dimethacrylate monomer, 150 high molecular-weight resin, 149 low-shrinkage TCD urethane based monomer, 150 ring-opening epoxy siloxane, 148 149 Low-shrinkage TCD urethane based monomer, 150 Low-temperature degradation (LTD), 63, 332 “Lower contraction stress” monomers, 143 Lower shrinkage stress monomer chemistry, 143 Lowest unoccupied molecular orbitals (LUMO), 75, 76f LPS. See Lipopolysaccharides (LPS) LSR. See Liquid silicone rubber (LSR) LTD. See Low-temperature degradation (LTD) Lucirin, 136 Luminous reflectance, 8 9 LUMO. See Lowest unoccupied molecular orbitals (LUMO) Luting/lining cement, 176 Luvit, 533 534 M M-510 gel catalyst, 628 M-511 platinum silicone rubber, 628 631 properties, 629 silicone fluids, 629 631 M-513 softening agent, 628 M-514 antislump agent, 628 Machining, 219 220 of restorations, 519 Macrolevel surface roughness, 585
Index Magnesia (MgO), 63 Magnesium (Mg), 362 biodegradable materials, 362 Magnetic resonance imaging (MRI), 508 515, 511t, 663 664 clinical applications, 509 512 intraoral, extraoral, including cone beam CT, 508 509 limitations, 512 515 Maleic acid, 218 Malleability, 19 Mandibular autografts, 352 353 Manganese (Mn), 543 544, 544t Marginal sealing, 116 117 Martensitic active NiTi, 560 Martensitic steels, 547 Matrix metalloproteinases (MMPs), 692 Maturation, 218 219 phase in GICs, 183 Maxillofacial pathology, 383 385 Maxillofacial prosthetic rehabilitation, 3 coloring agents, 635 638 historical background, 616, 617t ideal properties of maxillofacial material, 617, 618t M-511 platinum silicone rubber, 628 631 mechanical and physical properties, 618t primers, 632 search for ideal materials for, 617 620 silicones, 620 622 skin adhesives, 638 639 soft liners and tissue conditioners, 633 635 Maxillofacial prosthodontics, 616 Maxillofacial silicones, 622 628. See also Silicones A-304, 632 A-306, 632 A-330G, 632 A-2186, 626, 626f common room temperature vulcanizing silicones, 625 626 comparative properties, 627t food grade, 623 G-611, 632 HTV silicone elastomer, 625 implant grade, 622 industrial grade, 623 MDX4 4210, 626 719 medical grade, 623 liquid silicone elastomers, 627 VerSiTal silicone elastomers, 628 recommendations, 627 research studies about, 630t RTV silicone elastomer, 624 625 Maxillofacial surgery, 407 Maximum elastic deflection, 554 Maximum flexibility, 554 MCPA. See Monocalcium phosphate anhydrous (MCPA) MCPM. See Monocalcium phosphate monohydrate (MCPM) MDP. See Methacryloyloxydecyl dihydrogen phosphate (MDP) MDPB. See Methacryloyloxydodecyl pyridinium bromide (MDPB) Mechanical surface conditioning grit-blasting, 70 71 laser, 71 polishing/grinding, 71 Medical grade liquid silicone elastomers, 627 silicones, 623, 623f VerSiTal silicone elastomers, 628 Melt-blending process, 590 592, 591f Mercury toxicology, 118 119, 119t Mesenchymal stem cells (MSCs), 469, 583 584, 606 607 4-MET. See Trimelitic acid methacrylate (4MET) Metal brackets, 542 550. See also Ceramic brackets CoCr brackets, 548 precious, 550 SS brackets, 543 547 titanium brackets, 548 549 Metal(lic), 217 218, 305, 658 661 biodegradable metal scaffolds, 658 659 ceramics, 306 core of PFMs, 318 319 metal-free restorations, 380 metal-reinforced cements, 178 nanoparticles, 659 oxide pastes, 206 oxides, 320 titanium, 659 660 zirconia, 660 661 Methacrylate resin
720 Methacrylate resin (Continued) liners, 97 sealers, 280, 281t Methacrylate-based SEP, 236 Methacrylic acid, 128 Methacryloxylethyl-cetyldimethyl ammonium chloride (DMAE-CB), 153 Methacryloyl fluoride methyl acrylate, 152 Methacryloyloxy propyltrimethoxysilane (MPS), 68 Methacryloyloxydecyl dihydrogen phosphate (MDP), 68 70, 130, 330 Methacryloyloxydodecyl pyridinium bromide (MDPB), 153, 242 243 2-Methacryloyloxyethyl phosphorylcholine, dimethylaminohexadecyl methacrylate (DMAHDM), 243 244 Methyl groups (CH3 groups), 621 Methyl methacrylate monomers, 82 83 Methyl methacrylate structure, 83 Methylchloride, 621 622 Methylpyrrolidinone CS, 652 MFP. See Monofluorophosphate (MFP) MFP 1 DCPD, 387 MI varnish, 396 397 μ-ONE HA implant, 400 MicroMega MTA, 600 Microspheres of Mg-doped TCP, 398 399 Microwave-cured acrylic resins, 84 85 Milling, 219 220 burs, 519 Mineral trioxide aggregate (MTA), 257 261, 261t, 284, 600, 661 662 MTA Angelus, 262 MTA Bio, 600 MTA Fillapex, 284t, 600 601 MTA HP Repair, 600 MTA Plus, 284t, 600 601 Mineralization, 600 601 Mineralized freeze-dried bone allografts, 354 Mineralized tooth structure, 230 Minimally invasive dentistry, 319 320, 465, 675 676 Mixing technique adding powder and water, 43 measuring powder, 42 measuring water, 42 Modified glass bead sterilizer, 155 Modified stone. See Modified α-hemihydrate Index Modified α-hemihydrate, 39 40 Modulus of elasticity, 553 Modulus of resilience, 554 “Moist bonding” technique, 242 Moisture sensitivity of GICs, 189 Molybdenum, 543 544, 544t “Monoblock”, 277 Monocalcium phosphate anhydrous (MCPA), 358t, 395 Monocalcium phosphate monohydrate (MCPM), 358t, 389 390, 395 Monochromatic light, 8 9 Monocrystalline alumina brackets, 552 Monofluorophosphate (MFP), 387 Monolithic restoration, 517 Monolithic zirconia, 63 64 Monomer chemistry, 141 Monomer phosphate, 333 Moodle platforms, 533 534 MPS. See Methacryloyloxy propyltrimethoxysilane (MPS) MRI. See Magnetic resonance imaging (MRI) MSCs. See Mesenchymal stem cells (MSCs) MTA. See Mineral trioxide aggregate (MTA) MTAD, 267 269, 268t MTAS, 284t Mucocompressive techniques, 198 199 Mucosa supported guides, 526 Mucositis, 413 Mucostatic techniques, 198 199 Multiphasic CaPO4 formulations, 404 Multistranded NiTi wires, 560 Multistranded stainless steel wires, 557 N N,N-dimethyl-p-toluidine, 135 N-vinylpyrrolidone (NVP), 222 N’Durance composite, 150 NACP. See Nanoparticles of amorphous calcium phosphate (NACP) NaF. See Sodium fluoride (NaF) NAg. See Nanoparticles of silver (NAg) Nano-FAp (nFAp), 220 222 Nano-Hap. See Nanohydroxyapatite (nHAp) Nano-RMGICs. See Nanomodified resinmodified glass ionomer cements (Nano-RMGICs)
Index Nano-TiO2 (nTiO2), 222 223 Nanoapatite-filled glass ionomers, 189 190 Nanobiomaterials, 478, 481. See also Restorative nanobiomaterials dental applications, 481 492 endodontic, 488 490 for preventive dentistry, 482 483 Nanocapsules, 245 Nanoclusters, 133, 133f Nanocomplexes in CPP ACP, 483 Nanocomposites, 486 487 Nanodentistry, 477 478 Nanodiamond coatings, 485 486 Nanodiamond-loaded poly(lactide-coglycolide), 483 484 Nanofibrous scaffolds, 490 Nanofilled composites, 138 139 Nanofilled resin composite, 132 133 Nanofillers, 133, 133f Nanoglass ionomer cement applications of GICs, 219 nanomodifications of GIC powders, 219 224 Nanoglass ionomers, 487 488 Nanohydroxyapatite (nHAp), 220 222, 652 gutta percha, 489 Nanohydroxyl apatite toothpaste, 483 Nanoionomers, 487 488 Nanomaterials, 477 478, 481t, 490, 602. See also Biomaterials applications in disciplines of dentistry, 482f characterization, 478, 480f for dental implants, 484 486 general properties, 480 481 for periodontics, 483 484 potential, 492 493 and tissue engineering, 490 491 Nanomodifications of GIC powders, 219 224 GIC modified with nanoparticles, 222 223 mechanical properties of, 221t nanohydroxyapatite and ionomers, 220 222 nanomodified RMGICs, 223 224 powder-based, 220 of GIC powders, 219 224 glass ionomer cements modified with other nanoparticles, 222 223 721 mechanical properties, 221t nano-RMGICs, 223 224 nanohydroxyapatite and ionomers, 220 222 powder-based nanomodification of GICs, 220 Nanomodified resin-modified glass ionomer cements (Nano-RMGICs), 223 224 Nanoparticles (NPs), 82 GIC modified with, 222 223 nanoparticles-based endodontic sealer, 488 489 Nanoparticles of amorphous calcium phosphate (NACP), 242 243 Nanoparticles of silver (NAg), 242 243 Nanoscience, 478 480 Nanosized HAp, 222 Nanosized zirconia (nZrO2), 222 Nanostructured titanium implant surfaces, 589 Nanotechnology, 219, 477 478 Nanozirconia, 485 Native collagen, 647 Natural biomaterials, 217 218, 478 480, 647 653. See also Synthetic biomaterials agarose, 653 alginate, 649 650 collagen, 647 648 CS, 651 653 elastin, 653 fibrin, 648 649 hyaluronate, 651 silk, 650 651 Natural materials at metric scales, 478, 479f Natural polymers, 646 Natural tissues for bone grafting, 350 362 Naturally derived scaffold materials, 644 645 Naval Ordinance Laboratory (nol), 558 559 NBOs. See Nonbridging oxygens (NBOs) NC. See Network connectivity (NC) Neodymium (Nd), 531 Neodymium:yttrium aluminium garnet laser (Nd:YAG laser), 531 NeoMTA Plus, 600 Network connectivity (NC), 455 Newtonian fluids, 14 nFAp. See Nano-FAp (nFAp)
722 Index nHAp. See Nanohydroxyapatite (nHAp) Nickel, 543 544, 544t Nickel-containing alloys, 27 28 Nickel titanium wires (NiTi wires), 558 560 advantages and disadvantages, 559t generations, 559 560 Niobium (Nb2O5), 543 544, 544t, 602 Nisin, 244 245 NiTi wires. See Nickel titanium wires (NiTi wires) Nitinol, 558 559 Nitric acid (HNO3), 590 Nitrogen, 543 544, 544t NMR analysis, 183 Nonbiodegradable materials, 644 Nonbonded amalgam restorations, 113 114 Nonbonded bonded amalgam restorations, 114 Nonbridging oxygens (NBOs), 178 180, 457 Noneugenol sealers, 286t “Nonevaporable” water. See “Tightly bound” water Noninvasive radiography, 508 Nonreleased antibacterial agents, 152t Nonthermal processes, 590 592 Normal setting expansion, 47 Novabone product, 461 Novamin, 464, 465t NPs. See Nanoparticles (NPs) NVP. See N-vinylpyrrolidone (NVP) Nylon fibers, 88 89 Organofunctional (methoxy) group, 134 Organofunctional silanes, 134 Ormocers, 109 Orthodontic(s), 312, 383 385, 407 409, 541 archwires, 553 562 classification, 555 562, 555f properties and ideal features, 553 555, 554f brackets, 542 553, 543t classification of orthodontic materials, 542b clear aligners, 562 572 tooth movement, 565 566 OrthoMTA, 600 Orthophosphate ions, 390 391 Osseoconductive materials, 586 Osseointegration, 29, 581 582, 659 cellular and biomaterial aspects, 583 584 Osteoblasts, 347 348 Osteocalcin, 606 Osteoconduction, 349 Osteoconductive grafts, 645 Osteogenesis, 349, 606 607 Osteoinduction, 29, 348 349 Osteopontin, 606 Oxide ceramics, 55 56 densely sintered aluminum-oxide ceramic, 61 62 glass-infiltrated aluminum-oxide ceramic, 61 zirconia, 62 64 Oxiranes, 148 149 Oxyapatite, 358t O Occlusal imaging technique, 510t Octacalcium phosphate (OCP), 358t, 396 Odontoblasts, 303 304 “One-bottle” systems, 239 “One-step self-etch” adhesive system. See Seventh-generation adhesives Open systems, 518 519 Oral cavity, 71 pathology, 383 385 surgery, 407 tissues, 37 38 Oral hygiene instruction (OHI), 684 Organic acids, 378 P PAA. See Polyacrylic acid (PAA) Packable composites, 138 139 Packing techniques, 158f Panoramic imaging technique, 511t para-chloroaniline (PCA), 267 Paraffin, 9 11, 10f Partial fiber reinforcement, 87 Particle size of bioactive glasses, 468 469, 468t Particulate filler composites, 136 Particulate-reinforced resins alumina, 89 90 titanium, 90 91 zirconia, 90
Index Pattern wax, 204 205 PCA. See para-chloroaniline (PCA) PCL. See Poly-e-caprolactone (PCL) PDGF. See Platelet-derived growth factor (PDGF) Pediatric dentistry, 383 385 PEEK. See Polyetheretherketone (PEEK) PEG. See Polyethyleneglycol (PEG) PENTA. See Dipentaerythritol Penta acrylate phosphate (PENTA) Peptide-based biomaterials, 647 Perforation repair materials, 292 293, 293b Periapical imaging technique, 510t Periodontal disease, FRCs in, 312 Periodontics, 383 385, 411 413, 411f. See also Orthodontic(s) nanomaterials for, 483 484 Periodontology. See Periodontics PerioSim, 534 Periradicular surgery, 290 PET-G. See Polyethylene terephthalate-G (PET-G) PET-G polymer, 569 PFM. See Porcelain-fused metal (PFM) PFZ. See Porcelain fused to zirconia (PFZ) PGA. See Polyglycolic acid (PGA) pH effect, 46 17 4 PH SS alloys, 547 Phagocytic mechanisms, 364 Pharmacology, 2 Pharmacy, 2 Phenol-formaldehyde, 79 80 Phenyl-propanedione (PPD), 135 136 photoinitiator systems, 136 Phosphate glass, 454 phosphate-based bioglasses, 657 658 Phosphoric acid, 233 gel, 330 Phosphorous, 543 544, 544t Phosphorous pentoxide (P2O5), 657 658 Photography, 516 517 Photoinitiators, 135 136 Photopolymerization, 108 109 Photostimulable phosphor plates (PSP), 514 PICN. See Polymer-infiltrated ceramic network (PICN) Pigments, 136 Planmeca Planmill system, 518 519 723 Plasma gas etching treatment, 590 592 proteins, 583 technology, 331 Plaster of Paris, 217 218 Plastic brackets, 550 551, 551t plastic-coated NiTi, 560 region, 17 19 Plasticizers, 86, 619 Platelet-derived growth factor (PDGF), 349, 364 365 Platelets, 583 Platinum primers, 632, 633t PMMA. See Polymethyl methacrylate (PMMA) Polishability, 138 139 Polishing/grinding, 71 Poly-e-caprolactone (PCL), 654 655 Poly-Works, 518 Poly(1,4-b-D-glucopyranosamine). See Chitosan Poly(acrylic acid) (PAA), 176 177, 180 181, 190, 218 219 Poly(alkenoate) cement, 188 Poly(alkenoics), 180 Poly(D,L-lactide-co-glycolide), 483 Poly(D,L-lactide), 483 Poly(L-lactide)-co-(e-caprolactone) scaffolds, 483 484 Poly(methylmethacrylate)-based denture materials, 590 592 Polyacid(s), 180 181, 218 modified composite, 153 154 Polyaminocarboxylic acid, 266f Polybiquanide family, 273f Polycarbonate, 93 Polycrystalline alumina brackets, 552 ceramics, 306, 320, 328 zirconia, 552 Polydimethylsiloxane. See Silicones Polyelectrolyte cement, 176 177 Polyelectrolytes, 180 Polyethene, 89 Polyetheretherketone (PEEK), 586 587 dental implant modification, 590 592 Polyethers, 200, 200f polyether ether ketones, 3
724 Polyethylene, 308 fibers, 133 Polyethylene oxide. See Polyethyleneglycol (PEG) Polyethylene terephthalate-G (PET-G), 562 Polyethyleneglycol (PEG), 647, 654 Polyglycolic acid (PGA), 415, 655 PolyHEMA matrix, 190 191 Polymer-infiltrated ceramic network (PICN), 73, 74t blocks, 159 165 materials, 165 Polymeric/polymer(s), 357 fibers, 133 matrix fibers adhesion to, 310 impregnation of fibers with, 310 polymer-based biomaterials, 647 polymer-infiltrated ceramics, 58, 64 65 polymer monomer ratio, 82 reinforced composites, 586 587 Polymerization chemical structure and mechanism of, 82 85 curing mechanisms of acrylic denture base materials, 83 85 initiators, 190 kinetics, 142 shrinkage, 138, 140 143, 311 Polymerization shrinkage stress, 141, 155 affecting factors in dental composites C-factor, 142 polymerization kinetics, 142 substrate compliance, 142 viscoelastic behaviour, 142 volumetric shrinkage, 141 142 water sorption, 142 consequences, 141f management incremental layering technique, 143 light-curing modes, 142 143 low elastic modulus liners, 143 lower shrinkage stress monomer chemistry, 143 preheating, 143 Polymethyl methacrylate (PMMA), 79 80, 82, 99, 357 Polyoxyethylene. See Polyethyleneglycol (PEG) Index Polyphosphates (polyP), 361 Polypropylene fibers, 89 Polysaccharide-based biomaterials, 647 Polysiloxane (PVS), 202 203, 203f, 207 208 Polysorbate-80 detergent (Tween 80), 267 269 Polysulfide, 201 Polyurethane (PC), 564 brackets, 550 551 elastomers, 619 Polyvinyl chloride (PVC), 81 and copolymer, 619 Porcelain, 56, 71 Porcelain fused to zirconia (PFZ), 320 Porcelain-fused metal (PFM), 58, 318 Porcine bone graft tissue, 355 Pore-free high-fusing product, 57 Porous Fe alloys, 658 Portland cement, 257 261 Positive air pressure method, 565 Postextraction bone resorption, 412 Potassium iodide (KI), 675 676, 684 Potassium sulfate (K2SO4), 205 206 Pour-type denture resins, 85 Powder-based nanomodification of GIC, 220 Powder liquid ratio, 49, 85 86 liquid, 86 powder, 85 86 PPD. See Phenyl-propanedione (PPD) Pre-gel stage, 142 Precipitated hydroxyapatite, 358t Precipitation-hardenable steels, 547 Preheating, 143 Prepolymerized filler(s) fabrication, 132, 132f particles, 132 Pressure, 16 17 Presurgical 2D radiograph imaging, 522 Preventive dentistry, nanobiomaterials for, 482 483 Preventive nanocomposites surface coatings, 482 483 Primers, 67 70, 231, 632 Pristine graphene films, 599 600 Procera-software, 523, 529 Processing wax, 204 205 Proportional limit, 18 ProRoot MTA Gray, 257 261
Index Prosthetic abutments, 528 Prosthetic dentistry. See Prosthodontics Prosthodontics, 312, 383 385, 409 411 Protein-based biomaterials, 647 Proton-extruding adenosine triphosphatase, 691 PSP. See Photostimulable phosphor plates (PSP) “Pulse-delay”, 143 Punch-test method, 21 Pure titanium wires, 561 562 PVC. See Polyvinyl chloride (PVC) PVS. See Polysiloxane (PVS) Pyrolysis, 458 459 Pyroplastic flow, 57 Q QADM. See Quaternary ammonium dimethacrylate (QADM) QDs. See Quantum dots (QDs) QMiX, 270, 270t QPEI. See Quaternary ammonium polyethyleneimine (QPEI) QTH, 139 QTH LCU, 139 140 Quantum dots (QDs), 245 Quaternary ammonium compounds, 153 Quaternary ammonium dimethacrylate (QADM), 153, 242 243 Quaternary ammonium polyethyleneimine (QPEI), 153, 489 Quick-Set, 600 R Radiograph operators, 514 Raman analysis, 180 181 Range of activation, 554 Range of deflection, 554 RANKL. See Receptor activator of nuclear factor jB ligand (RANKL) Rapid heat-polymerized resin, 83 Rapid Injection System, 92 Rapid prototyping (RP), 93 94, 519 520 RapidForm, 518 Rayon flocking for silicone elastomers, 635 636, 636f Recaldent, 396 397 Receptor activator of nuclear factor jB ligand (RANKL), 364 365 725 Recombinant human bone morphogenetic protein-2 (rhBMP-2), 413 Recombinant human growth and differentiation factor-5 (rhGDF-5), 413 Recombinant human platelet-derived growth factor BB (rhPDGF-BB), 413 Red Elgiloy, 558 Reduced GO (rGO), 600 Regenerative tissue engineering, 644 Reinforcement glass. See Glass R Released antibacterial agents, 152t Relief ACP, 393 394 Remineralization, 378f, 391 392, 465 Resilience, 19, 19f Resilon, 277 Resin composites, 128 attempts for improvement regarding both material formulation and manipulation, 156 165 regarding manipulation, 154 156 regarding material formulation, 148 154 chemical structure of resin systems, 131f classification according to composite consistency, 137 138 according to curing techniques, 138 according to fillers size and distribution, 136, 137f, 137t according to packing (placement) technique, 138 clinical indications, 138 139 composition fillers, 130 134 initiator accelerator system, 134 136 pigments, 136 resin matrix, 128 130 silane coupling agent, 134 degradation, 147 evolution, 129f guidelines for laboratory evaluation, 166 methods for mechanical evaluation, 167t properties and limitations biocompatibility, 145 147 clinical durability, 147 148 degradation, 147 degree of conversion, 139 140 esthetic properties, 144 145
726 Resin composites (Continued) mechanical properties, 145, 146t polymerization shrinkage, 140 143 polymerization shrinkage stress, 141 solubility, 144 thermal properties, 144 recommendations for future clinical studies, 166 171 strain, 141 Resin-modified glass-ionomer cements (RMGICs), 190 191, 223 224, 487 nanomodified, 223 224 Resin(s), 620 cements, 130 increment thickness effect, 139 matrix, 128 130 resin-based composites, 136 restorations, 138 resin-based dental composites, 486 resin-based materials, 319 resin-modified cements, 178 resin ceramic bond strength, 71 resin ionomer suspension and compomer, 292 sealers, 280 283, 286t shade effect, 139 Resistivity values of dental hard tissues, 15, 15t Resorbable-GTR membranes, 483 484 Restoration, 317 Restorative dental biomaterials, 380 Restorative dental materials, 15, 318 Restorative dentistry, 107, 317 Restorative materials, 218, 469 Restorative nanobiomaterials, 486 488. See also Nanobiomaterials nanocomposites, 486 487 nanoglass ionomers, 487 488 RetroMTA, 600 Retroplast, 290 292 Retroplast-GLUMA system, 290 292 RGD peptides, 651 rGO. See Reduced GO (rGO) rhBMP-2. See Recombinant human bone morphogenetic protein-2 (rhBMP-2) rhGDF-5. See Recombinant human growth and differentiation factor-5 (rhGDF-5) rhPDGF-BB. See Recombinant human platelet-derived growth factor BB (rhPDGF-BB) Index Rigid IMs, 198 199 Ring polymers, 620 Ring-opening epoxy siloxane, 148 149 RMGICs. See Resin-modified glass-ionomer cements (RMGICs) Room-temperature vulcanization silicone (RTV silicone), 98, 624 625 cross-linking by addition reaction, 625 by condensation reaction, 624 625 Root canal chlorhexidine, 266 267 citric acid, 267 EDTA, 265 266 filling materials, 404 405 HEBP, 269 hydrogen peroxide, 269 IKI, 269 irrigants interactions between different, 268t properties and effects of, 264t materials used as root canal irrigants, 263 270 MTAD, 267 269, 268t obturation, 404 405 materials, 274 289 properties, 263b QMiX, 270, 270t sealers, 277 289, 278b, 286t sodium hypochlorite, 263 265 tetraclean, 269 Root caries management in the elderly, 684 Root-end filling materials, 290 292 Amalgam, 290 composite resins, 290 292 Diaket, 292 GIC, 292 other types of cement, 292 requirements of, 290b resin ionomer suspension and compomer, 292 zinc oxide eugenol cements, 290, 291t RP. See Rapid prototyping (RP) RTV silicone. See Room-temperature vulcanization silicone (RTV silicone) Rubber-reinforced resins, 87 Rubbery consistency silicones, 622 RUNX2, 606 Rutile, 549
Index S S174000 PH alloys, 547 S2G. See Scan2Guide (S2G) Saliva, 106 Salivary functions, 379 Sand-blasting. See Grit-blasting Sandwich technique, 155 Sankin apatite root canal sealer, 284t, 406, 406t Saturation of color, 9 Scaffolds, 465 466, 644 for bone tissue engineering, 645 646 functions and features, 646 Scan2Guide (S2G), 523 Scanning electron microscopy (SEM), 460, 464f, 466f “Sclerotic dentin”, 235 Scopus database, 383 385, 384t Scotchbond 2, 238 Sculpturing clays and waxes, 635, 635f Second-generation adhesives, 237 238 of dental zirconia, 63 64 Secondary caries, 242 243 Selective infiltration etching, 332 Selective laser sintering, 520 521 Selective pressure techniques, 198 199 Selenium, 543 544, 544t Self-adhesive composites, 156 157, 157t resin cements, 130 Self-cured composites, 135 Self-etching primer (SEP), 236 Self-etching technique, 236 Self-healing composites, 154 Self-priming etchant. See Two-step self-etch systems Self-setting apatite-forming, 401 402 SEM. See Scanning electron microscopy (SEM) Sensitization, 543 544 SEP. See Self-etching primer (SEP) Seventh-generation adhesives, 240 Seventh-generation adhesives, 241 SF. See Silk fibroin (SF) Shade selection, 516 517 shade-matching technique, 516 517 Shear strength, 21 Shore A hardness of VST silicones, 628, 629t 727 Short-term fluoride burst release, 184 resilient liners, 98 99 Sialogram, 511t SICAT software, 523 Sieverts unit (Sv unit), 513 Silane(s), 319, 332 333 coupling agent, 134 film thickness, 134 silane-treated aluminum particles, 89 90 Silanization, 485 486 Silastic MDX4 4210 biomedical grade elastomer, 626 Silica (SiO2), 130 131, 455 456, 657 658 nanocomposites, 485 silica-based filler, 91 vapor nanofilm, 331 Silicate cement, 176 silicate-based bioactive glasses, 457 458 Silicon, 657 658 Silicon carbide, 133 134 Silicon dioxide (SiO2), 91, 130 Silicon nitride, 133 134 Silicones, 620 622 A-221-05, 627 A-223-30, 627 A-223-40, 627 basic silicone unit, 621f basic steps in silicone polymer synthesis, 621f composition of silicone elastomer, 622t condensation, 203 fluids, 629 631 polymer structures, 620 622, 620f PVS, 202 203 silicone-based liners, 97 98 silicone-based sealers, 286t VPES, 204 Silk, 650 651 Silk fibroin (SF), 650 Silorane, 130, 143, 148 149 Silver (Ag), 90 91, 153 points, 274 275, 275t Silver diamine fluoride (SDF), 675 700 cariostatic effects on caries in roots of the elderly, 682t on cariogenic bacteria, 690 692
728 Silver diamine fluoride (SDF) (Continued) on coronal caries in deciduous and permanent teeth, 678t on dental mineral, 685 689, 686t on organic content of dentine, 692 clinical effects on caries management, 676 684 in adults, 683 684 in children, 677 683 in the elderly, 684 history, 676 safety of SDF treatment, 692 694 Silver fillings. See Dental amalgam Silver modified ARTs (SMART), 683 684 Silver NPs (Ag NPs), 90 91 SIMODONT dental trainer, 534 Simplant, SurgiCase software, 523 Simvastatin-loaded porous implant surfaces, 660 Single-composition alloy, 112 113 Single-crystal alumina, 552 Single-edge-notched test method, 21 Sintering process, 56 57, 318, 552 Sixth-generation adhesives, 239 240 Sizing process, 310 Skin adhesives, 638 639 Slim, 518 Slip casting process, 306, 320 Smart composites, 150 154, 151f antibacterial composites, 152 153, 152t bioactive remineralizing composites, 150 151 polyacid modified composite, 153 154 self-healing composites, 154 SmartTrack, 564 565 Smear Clear, 270t Smear layer, 231 232 adhesives classification based on interaction with, 241 dissolving systems, 241 removing systems, 241 Sodium chloride, 45 46 Sodium fluoride (NaF), 152 Sodium hypochlorite (NaOCl), 231 232, 235, 263 265, 330 Sodium oxide (Na2O), 455 456, 657 658 Sodium phosphate, 201 202 Sodium-containing glass G338, 184 185 Sodium fluoride (NaF), 684, 690 691 Index Sodium-free glass LG26, 184 185 Soft liners, 633 635, 634f Soft relining, 96 98 methacrylate resin liners, 97 silicone-based liners, 97 98 “Soft-start” protocol, 143 Solid liquid transition endotherms, 9 11 Solid solid endotherms, 9 11 Solubility, 455 456 of resin composites, 144 Spatulation, 45 Specialty dentistry, 383 385 Specific heat, 13 Spin-coating, 590 592 Springback, 554 SrF2. See Strontium fluoride (SrF2) SrO. See Strontium oxide (SrO) SS. See Stainless steel (SS) Stabilized martensitic NiTi, 559 Stainless steel (SS), 463, 542 543 alloys, 542 543, 546t archwires, 556 557 advantages and disadvantages, 556t austenitic SS, 545 547 brackets, 543 547 classifications, 544 545 effect of different elements on properties, 544t materials, 545b Stannous octoate, 203 Steel, 305 18 8 steel. See 304L steel 304L steel, 545 Stereolithography, 527 Sterilization methods, 77 Stiffness, 554 Strain, 17, 18f Strength, 554 Streptococcus mutans, 243 244, 378, 690 691 Stress, 17 relaxation, 568 stress-shielding, 584 585 stress strain cure region, 18 relationship, 16 Strontium, 144 Strontium fluoride (SrF2), 152 Strontium oxide (SrO), 130 131
Index Substrate compliance, 142 Subtractive manufacturing, 517 Sugar-free gums, 390 Sulfuric acid (H2SO4), 590 Super austenitic SS, 547 Super ethoxybenzoic acid (Super EBA), 290 292, 291t Supercable NiTi, 560 Superelastic NiTi, 560 Surface chemistry, 74 76 grinding methods, 330 331 hardness, 50 51 roughness, 585 smoothness, 145 topography of zirconia surface, 329 333 Synthetic biomaterials, 217 218, 647, 654 655. See also Natural biomaterials for bone grafting, 350 362 PCL, 654 655 PEG, 654 PGA, 655 Synthetic calcium phosphate materials, 246 Synthetic HA, 357 359 Synthetic materials at metric scales, 478, 479f Synthetic polymers, 646 Synthetic scaffold materials, 644 645 Syringe, 14 Systemic toxicity, 24 T TA. See Tertiary amine (TA) Tantalum, 543 544, 544t Tarnish, 116 117 Tartaric acid ( 1 ), 180 181 TCD-DI-HEA, 143 TCP. See Tricalcium phosphate (TCP) Tech Biosealer, 600 Teeth anatomy and physiology, 302 304 cementum, 304 dental pulp, 304 dentin, 303 304 enamel, 303 cross section, 303f development, 304 mechanical properties, 305 729 remineralization, 385 386, 390 392 TEETHMATE DESENSITIZER, 393 394 TEGDMA. See Triethylene glycol dimethacrylate (TEGDMA) Temperature, 9, 10f, 45 Temperature transition range (TTR), 559 Temporomandibular joint disorders, 512 Tensile bond strength test, 71 73 loading, 19 strength, 20, 50, 114 115 Tertiary amine (TA), 135 Tetracalcium phosphate (TTCP), 389 390, 403 404 Tetraclean, 269 Tetracycline-HCl, 660 TetricEvoCeram (IvoclarVivadent, Liechtenstein), 130 131 TFS. See Three-point FS (TFS) Thallium amalgam, 109 TheraCal LC, 600 Thermaflo, 155 Thermal conductivity, 12, 13t of fine composites, 144 Thermal diffusivity, 13 14 Thermal shock, 58 Thermal stress, 57 58, 144 Thermo-cured glass ionomers, 190 Thermocouples, 9 Thermoelasticity of NiTi alloys, 559 560 Thermoforming process, 565 Thermogram, 9 11, 10f Thermoplastic aligners forces, 565 567 differences in force generation, 565 566 material factors affecting force delivery, 566 567 materials used for, 562, 564t Thermoplastic resins, 92 93 thermoplastic acetal, 92 thermoplastic acrylic, 93 thermoplastic nylon, 92 thermoplastic polycarbonate, 93 Thermoset urethane elastomers, 619 Thiol-ene siloxane, 202 203 Third generation adhesives, 238 of zirconia, 64
730 Thixotropic silicone extrinsic stains, 636, 637f Thixotropy, 14 Three-dimension (3D) CT imaging, 522 523 images, 509 510 printed scaffolds, 663 664 printers, 522 printing, 519 521, 530, 663 664 in implant dentistry, 521 522 systems, 37 38 structure, 645 Three-point flexural bend test, 186, 187f Three-point FS (TFS), 187 188 “Three-step etch-and-rinse” technique. See “Total-etch” technique Ti 6Al 4V alloy, 549, 585 587, 659 “Tightly bound” water, 181 Timolium wires, 562 Tissue conditioners, 98 99, 633 635 engineering, 1, 465 466, 490 491 approaches in calcium orthophosphate, 414 415 scaffolds, 1 2 triad, 644 regeneration, 490 491 Titania (TiO2), 75 76, 222 223, 589 Titanium, 90 91, 305, 543 544, 544t, 585 587, 659 660 brackets, 548 549 foams, 660 implants, 325 surface modifications and respective properties, 586t modification of titanium implants, 588 590 nanocomposites, 485 titanium-based wires, 561 562 titanium niobium wires, 561 titanium vanadium wires, 562 Titanium dioxide. See Titania (TiO2) Titanium plasma spraying (TPS), 588 589, 588f Titanium molybdenum alloys (TMA), 561 properties, 561b Tomography, 511t Tooth decay. See Dental caries Index restoration, 311 tooth-supported guides, 526 Toothpastes, 386 389, 388t, 463 464 Top-down manufacturing, 219 220 Top-down nanomanufacturing, 219 220 “Total-etch” technique, 238 240 Toughness, 19 Toxic reaction, 25 26 Toxicity, 24 25 Toxicology of mercury, 118 119 TPO. See Diphenyl (2,4,6-trimethylbenzoyl) phosphine oxide (TPO) TPS. See Titanium plasma spraying (TPS) Transbond XT, 407 408 Transition temperatures, 9 11 Transmission electron microscopy, 112 Tribochemical silica, 331 Tricalcium phosphate (TCP), 355 356, 359, 388t, 394, 646, 655 656 Triclosan, 153 Triethylene glycol dimethacrylate (TEGDMA), 128 130 TEGDMA/Bis-GMA ratios, 141 Trimelitic acid methacrylate (4-MET), 156 157 Triple antibiotics pastes, 273 Tris-buffer solution, 454 455 Tropoelastin, 653 TTCP. See Tetracalcium phosphate (TTCP) TTR. See Temperature transition range (TTR) Twisted wires, 557 Two-dimensional image (2D image), 514 515 Two-step etch-and-rinse system, 240 self-etch systems, 241 systems, 240 241 Tx Studio software, 523 U UDMA. See Urethane dimethacrylate (UDMA) UF. See Urea formaldehyde (UF) Ultimate tensile strength, 20 Ultraviolet (UV), 77 Unalloyed zirconia, 326 Uncontrolled periodontal lesions, 483 Undisclosed CaPO4, 388t
Index Unidirectional fibers, 308 309 United Nations Environment Programme (UNEP), 106 Urea formaldehyde (UF), 154 Urethane dimethacrylate (UDMA), 84, 128 130, 153 154 Urethane methacrylate resin sealers, 280 US Food and Drug Administration (FDA), 454 UV. See Ultraviolet (UV) V Vacuum-form method, 565 Vascular endothelial growth factor (VEGF), 364 365 VDO. See Vertical dimension of occlusion (VDO) Venus Diamond Nanohybrid Composite (Heraeus Kulzer), 143 Venus Diamond resin composite, 150 Vertical dimension of occlusion (VDO), 94 Vibrant human oral atmosphere, 16 Vinyl polyether siloxane (VPES), 204 Vinyl resins, 81 Vinylpolysiloxane (VPS), 204 Virtual Implant Planning software, 523 Viscoelastic behaviour, 142 Viscoelasticity, 305 Viscosity, 14, 207 208 Visible light cured resins, 84 Vita Enamic, 159 165 Vital pulp therapy bioaggregate, 262 biodentine, 262 endosequence, 262 263 materials used in, 256 263, 258t MTA, 257 261 MTA Angelus, 262 properties, 257b Vitapan Chart system, 516 517 Vitrification, 56 Volumetric shrinkage, 141 142 Voxel-Man, 534 VPES. See Vinyl polyether siloxane (VPES) VPS. See Vinylpolysiloxane (VPS) VST silicones, 628 elastomers for maxillofacial prosthetics, 628f properties, 629t 731 Vulcanite, 79 80 Vulcanization reaction of polymers, 624 W W/P ratio. See Water/powder ratio (W/P ratio) Water absorption, 568 569 wear, abrasion, and intraoral aging, 569 in GICs, 181 sorption, 142, 310 storage of resin composites, 144 water-based adhesives, 639 Water/powder ratio (W/P ratio), 44 45, 50f Watt’s method, 149 Wax(es), 204 205 warmer, 155 Wear resistance, 167t Weibull moduli, 73 Wettability of IMs, 208 Whiskers, 133 134 White spot lesions (WSL), 246 247 Wollastonite (CaSiO3), 602 Working range, 554 World Health Organization, 189 WSL. See White spot lesions (WSL) X X-ray diffraction (XRD), 28 29, 458 459 X-ray(s), 663 664 photoelectron spectroscopy, 457 458 Xenoestrogenic effect, 147 Xenografts, 354 355. See also Allografts XRD. See X-ray diffraction (XRD) Y Y-TZP. See Yttria-stabilized tetragonal zirconia crystal (Y-TZP) YAG. See Yttrium aluminium garnet (YAG) YbF3. See Ytterbium trifluoride (YbF3) Yellow Elgiloy, 558 Yield strength, 18 Young’s modulus, 108 109, 118, 484 485, 553 Ytterbium, 130 131 Ytterbium trifluoride (YbF3), 152 Yttria (Y2O3), 63, 326
732 Yttria-stabilized tetragonal zirconia crystal (Y-TZP), 63 Yttrium, 660 661 Yttrium fluoride, 130 131 Yttrium oxide. See Yttria (Y2O3) Yttrium-stabilized tetragonal zirconia, 326 328 physicochemical, electrical, and thermal properties of, 327t tetragonal monoclinic transformation phase during crack propagation, 327f Yttrium-stabilized zirconium oxide (ZrO2, Y-TZP), 527 528 Yttrium aluminium garnet (YAG), 531 Z Zinc oxide (ZnO), 117 118, 133 134, 242 243, 245 nanoparticles, 153 Zinc oxide eugenol (ZOE), 271, 278, 279t cements, 176, 290, 291t sealers, 278, 279t, 286t Zinc oxide eugenol impression paste. See Metallic oxide pastes Zinc phosphate cement, 176 Zirconia (ZrO2), 61 64, 62f, 68f, 72t, 75f, 90, 222, 320, 322, 328, 485, 586 587, 660 661 adhesion to, 329 333 Index classification, 321 324, 321t, 322f in dentistry, 222, 324 325, 325t failure and fractographic analysis of zirconia restorations, 333 335 fully crystallized glass-ceramic milling block, 321f limitations and challenges, 337 338 mechanical testing of zirconia ceramics, 335 337 modification of zirconia implants, 592 surface topography, clinical treatments of zirconia surface, 329 333 yttrium-stabilized tetragonal zirconia, 326 328 zirconia-based ceramics, 324 zirconia-lithium silicate, 60 61 zirconia/silica, 133 134 ZTA, 328 329 Zirconia-toughened alumina (ZTA), 320, 328 329 Zirconium (Zr), 324, 661 Zirconium dioxide, 324, 660 661 Zirconium-oxide/zirconia ceramics, 55 56 ZnO QDs (ZnOQDs), 245 ZOE. See Zinc oxide eugenol (ZOE) Zombies, 691 692 ZTA. See Zirconia-toughened alumina (ZTA)