Handbook of Bioceramics and Biocomposites 9783319124612

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Iulian Vasile Antoniac Editor

Handbook of Bioceramics and Biocomposites 1 3Reference

Handbook of Bioceramics and Biocomposites

Iulian Vasile Antoniac Editor

Handbook of Bioceramics and Biocomposites With 367 Figures and 80 Tables

Editor Iulian Vasile Antoniac University Politehnica of Bucharest Bucharest, Romania

ISBN 978-3-319-12459-9 ISBN 978-3-319-12460-5 (eBook) ISBN 978-3-319-12461-2 (print and electronic bundle) DOI 10.1007/978-3-319-12460-5 Library of Congress Control Number: 2016933676 # Springer International Publishing Switzerland 2016 This work is subject to copyright. All rights are reserved by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, express or implied, with respect to the material contained herein or for any errors or omissions that may have been made. Printed on acid-free paper This Springer imprint is published by SpringerNature The registered company is Springer International Publishing AG Switzerland

Dedicated to the memory of Professor Raquel LeGeros and Professor Daniel Bunea

Foreword

This 44-chapter Handbook of Bioceramics and Biocomposites edited by Prof. Iulian Antoniac of the University of Bucharest provides a comprehensive view of one of the best investigated classes of biomaterials in research and practice, namely bioceramics and related biocomposites. In his logical approach to this theme, Prof. Antoniac uses a tripartite division, beginning with the “History and Materials Fundamentals” in which 15 chapters cover terrain from the early developments via different bioceramic classes to an exciting variety of composites, for example with polymers, graphene, or natural biopolymers, such as collagen. The fact that the first two chapters, which are on the history and development of bioceramics and bioactive glasses, are authored by the two “fathers” of these fields, Profs. Guy Daculsi and Larry Hench, respectively, puts the standard of the entire handbook on elevated ground. Nevertheless, the choice of the remaining authors is ample proof that the Editor was determined to make no compromises with respect to the knowledge, experience, and standing of all those chosen to be contributors. The middle part, “Materials Engineering and Biological Interactions,” is naturally sub-divided into 6 chapters which address biomimetic strategies and process engineering of these biomaterials, followed by a set of 6 chapters with a life science emphasis. In the latter, surface engineering techniques and themes quintessential to interactions with living systems, including protein interactions as well as engineering approaches to reduce microorganism adhesion, are discussed. In addition, testing systems in vitro and in vivo as well as bioceramic strategies for tissue engineering especially of the musculoskeletal system have an important focus. The third and final part, “Clinical Performance in Bioresorbable and LoadBearing Applications,” is a state-of-the-art resumé of what bioceramics and biocomposites have so far achieved in patient care. It is fitting that in this subdivided part, the field of orthopedics should lead the way, followed by dentistry, and then completed with other applications, including carriers for medication and implants for cranioplasty or the orbital region. Professor Antoniac has succeeded in compiling a comprehensive handbook which will undoubtedly become a standard reference work on the theme of bioceramics and biocomposites. The all-embracing nature of the approach he has

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adopted has been highly effective in spanning the traditional fields of development in these biomaterial classes as well as their novel modifications, which hold promise for a wide range of clinical applications in the future. He and his truly international author team have to be congratulated on this invaluable contribution to the biomaterial literature. C. James Kirkpatrick M.D., Ph.D., D.Sc., F.R.C.Path. Emeritus Professor of Pathology Johannes Gutenberg University Mainz, Germany

Preface

Information related to different aspects about bioceramics and biocomposites science together with their accompanying technology and applications in medical practice is scattered in the relevant literature. In provided this Handbook of Bioceramics and Biocomposites, the editor believes that the latter stage has been reached in many parts of these investigated classes of biomaterials in research and practice. Also, many medical applications based on the bioceramics and composites are in clinical use for a long time and researchers have been studying their performance in order to offer potential solutions for their improvement. In approaching his task, the Editor has tried to bring together into one source book all the information that is available about bioceramics and related biocomposites in terms of material fundamentals, materials engineering, biological interactions, and clinical performance in various medical applications, from orthopedics and dentistry to carriers for medication and implants for cranioplasty or the orbital region. In order to do this, I asked for the help of many colleagues worldwide to be contributors to this handbook. Another important fact was that the contributors have different backgrounds, from materials science to biology or clinicians in different medical specializations. The topic of bioceramics and related biocomposites has attracted many researchers of other fields to make contributions, at the same time helping traditional ceramic science and technology in its transition to work at multidisciplinary research frontiers. Having a synergic effect with the rapid developments of related areas, e.g., tissue engineering, nanotechnology, drug delivery, smart materials, and structures, bioceramics and related biocomposites has become a frontier field of research leading to many technological breakthroughs. As a result, Handbook of Bioceramics and Biocomposites has expanded to three main parts. They cover the following topics: The first part, “History and Materials Fundamentals,” sub-divided into 15 chapters addresses history and development, fundamental properties, and presentation of the main bioceramics, like alumina, zirconia, calcium phosphates, hydroxyapatite, and carbonate apatite, and a large variety of related composites, from ceramicpolymer to grapheme-ceramic. “In order to move forward we must look back.” Based on this consideration and as a sign of respect for the pioneers that opened this ix

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field, the first two chapters of this handbook are dedicated to the history and development of bioceramics and bioactive glasses. Authored by two pioneers of bioceramics and related biocomposites, these chapters demand higher and elevated standards for all contributors of the handbook. Other chapters of this part present the fundamental properties and different aspects about the major bioceramics, like alumina, zirconia, calcium phosphates, hydroxyapatite, and carbonate apatite, and a large variety of related composites, from ceramic-polymer to grapheme-ceramic. The second part, “Materials Engineering and Biological Interactions,” covers two important aspects: biomimetic strategies, and surface engineering and interactions with living cells. Six chapters are dedicated to biomimetic strategies and present essential requirements and manufacturing and evaluation aspects for different bioceramics and related composites used for tissue engineering and regeneration. The last six chapters of this part are dedicated to surface engineering and interactions with living cells, and describe processing technologies, characterization, and biocompatibility evaluation of different bioceramic coatings and biocomposites. The last part, “Clinical Performance in Bioresorbable and Load-Bearing Applications,” provides a generous view on the clinical performance of various medical applications which comprise 17 chapters. The field of orthopedics and dentistry lead the way. Many clinicians have contributed to these parts describing the performance of various implants based on their clinical experience or retrieval analysis. The last chapters of this part describe other applications, including carriers for medication and implants for cranioplasty or the orbital region. This part is very important because the dedicated chapters are written mainly by clinicians, who could appreciate better the performance of various bioceramics and related biocomposites. Until the end, the clinical applications of the new bioceramics or related biocomposites must be the main target of the researchers who work in this field, and collaborative activities with clinicians are very useful and important. It is hoped that the handbook will be used and useful, not perfect but a valuable contribution to the bioceramics and related biocomposites field that I believe is evolving sufficiently to deserve such a publication. The handbook can also serve as the basis of instructional course lectures for audiences ranging from advanced undergraduate students to post-graduates in materials science and engineering and biomedical engineering. I wish readers to find the Handbook of Bioceramics and Biocomposites informative and useful in their endeavors. March 2016 Bucharest, Romania

Iulian Vasile Antoniac

Contents

Volume 1 Part I 1

2

History and Materials Fundamentals

...................

1

History of Development and Use of the Bioceramics and Biocomposites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Guy Daculsi

3

Bioactive Glass Bone Grafts: History and Clinical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Larry L. Hench

23

3

Fundamental Properties of Bioceramics and Biocomposites . . . . . Maria Grazia Raucci, Daniela Giugliano, and Luigi Ambrosio

35

4

Bioinert Ceramics: Zirconia and Alumina . . . . . . . . . . . . . . . . . . . Corrado Piconi and Alessandro Alan Porporati

59

5

Calcium Phosphates Sergey V. Dorozhkin

....................................

91

6

Hydroxyapatite: From Nanocrystals to Hybrid Nanocomposites for Regenerative Medicine . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Anna Tampieri, Michele Iafisco, Simone Sprio, Andrea Ruffini, Silvia Panseri, Monica Montesi, Alessio Adamiano, and Monica Sandri

119

7

Cationic and Anionic Substitutions in Hydroxyapatite Ilaria Cacciotti

.........

145

8

Carbonate Apatite Bone Replacement . . . . . . . . . . . . . . . . . . . . . . Kunio Ishikawa

213

9

Natural and Synthetic Polymers for Designing Composite Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bogdan C. Simionescu and Daniela Ivanov

233

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10

Ceramic-Polymer Composites for Biomedical Applications . . . . . . Toshiki Miyazaki, Masakazu Kawashita, and Chikara Ohtsuki

287

11

Collagen–Bioceramic Smart Composites . . . . . . . . . . . . . . . . . . . . Iulian Vasile Antoniac, Madalina Georgiana Albu, Aurora Antoniac, Laura Cristina Rusu, and Mihaela Violeta Ghica

301

12

Bioactive Glass-Biopolymer Composites for Applications in Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Yaping Ding, Marina T. Souza, Wei Li, Dirk W. Schubert, Aldo R. Boccaccini, and Judith A. Roether

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13

Resin-Based Dental Composite Materials . . . . . . . . . . . . . . . . . . . . Hanadi Y. Marghalani

357

14

Composite Hybrid Membrane Materials for Artificial Organs . . . Stefan Ioan Voicu and Marius Sandru

407

15

Graphene-Bioceramic Composites . . . . . . . . . . . . . . . . . . . . . . . . . Xingyi Xie and Marta Cerruti

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Part II Materials Engineering and Biological Interactions: Biomimetic Strategies . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 16

17

18

Essential Requirements for Resorbable Bioceramic Development: Research, Manufacturing, and Preclinical Studies . . . . . . . . . . . . . Guy Daculsi, Eric Aguado, and Thomas Miramond

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Biomimetic Strategies to Engineer Mineralized Human Tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Sandra Pina, Joaquim Miguel Oliveira, and Rui L. Reis

503

Biomimetics and Marine Materials in Drug Delivery and Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Andy H. Choi, Sophie Cazalbou, and Besim Ben-Nissan

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19

Silicate-Based Bioactive Composites for Tissue Regeneration . . . . Y.L. Zhou, Z.G. Huan, and J. Chang

20

Biomimetic Customized Composite Scaffolds and Translational Models for the Bone Regenerative Medicine Using CAD-CAM Technology . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Isidoro Giorgio Lesci, Leonardo Ciocca, and Norberto Roveri

21

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In Vitro and In Vivo Evaluation of Composite Scaffolds for Bone Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Svetlana Schussler, Khadidiatou Guiro, and Treena Livingston Arinzeh

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Contents

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Volume 2 Part III Materials Engineering and Biological Interactions: Surface Engineering and Interactions with Living Cells . . . . . . . . . . . . . . . . 22

Processing Technologies for Bioceramic Based Composites . . . . . . Ipek Akin and Gultekin Goller

23

Glass-Ceramics: Fundamental Aspects Regarding the Interaction with Proteins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . C. Gruian, E. Vanea, H.-J. Steinhoff, and Simion Simon

637 639

667

24

Bioceramic Coatings for Metallic Implants . . . . . . . . . . . . . . . . . . Alina Vladescu, Maria A. Surmeneva, Cosmin M. Cotrut, Roman A. Surmenev, and Iulian Vasile Antoniac

703

25

Sol-gel Nanocoatings of Bioceramics . . . . . . . . . . . . . . . . . . . . . . . B. Ben-Nissan, A.H. Choi, I.J. Macha, and S. Cazalbou

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26

Biomaterial Functionalized Surfaces for Reducing Bacterial Adhesion and Infection . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Maria G. Katsikogianni, David J. Wood, and Yannis F. Missirlis

757

Biocomposites used in Orthopedic Applications: Trends in Biocompatibility Assays . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Martin J. Stoddart and Mauro Alini

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Part IV Clinical Performance in Bioresorbable and Load-Bearing Applications: Orthopedics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . ...

28

Perspective and Trends on Bioceramics in Joint Replacement Corrado Piconi and Giulio Maccauro

29

Evolution of Cementation Techniques and Bone Cements in Hip Arthroplasty . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Marius Niculescu, Bogdan Lucian Solomon, George Viscopoleanu, and Iulian Vasile Antoniac

819 821

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30

Retrieval Analysis of Hip Prostheses . . . . . . . . . . . . . . . . . . . . . . . 901 Iulian Vasile Antoniac, Florin Miculescu, Dan Laptoiu, Aurora Antoniac, Marius Niculescu, and Dan Grecu

31

Clinical Limitations of the Biodegradable Implants Used in Arthroscopy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Rodica Marinescu and Iulian Vasile Antoniac

32

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Bioceramics and Biocomposites in Spine Surgery . . . . . . . . . . . . . 967 Gianluca Vadalà, Fabrizio Russo, Luca Ambrosio, and Vincenzo Denaro

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Part V Clinical Performance in Bioresorbable and Load-Bearing Applications: Dentistry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33

Biofilm Formation on Implants and Prosthetic Dental Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Lia Rimondini, Andrea Cochis, Elena Varoni, Barbara Azzimonti, and Antonio Carrassi

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. . . . . . . . . . . . . . 1029

34

Guided Bone Regeneration for Dental Implants Mishel Weshler and Iulian Vasile Antoniac

35

Performance of Dental Composites in Restorative Dentistry . . . . . 1075 Diana Dudea, Camelia Alb, Bogdan Culic, and Florin Alb

36

Clinical Evaluation of Disilicate and Zirconium in Dentistry Domenico Baldi, Jacopo Colombo, and Uli Hauschild

37

Ceramic Veneers in Dental Esthetic Treatments . . . . . . . . . . . . . . 1129 Dan Pătroi, Teodor Trăistaru, and Sergiu-Alexandru Rădulescu

38

Alveolar Augmentation Using Different Bone Substitutes . . . . . . . 1159 Cena Dimova, Biljana Evrosimovska, Katerina Zlatanovska, and Julija Zarkova

39

CAD-CAM Processing for All Ceramic Dental Restorations . . . . . 1201 Alexandru Eugen Petre

40

Failure Analysis of Dental Prosthesis . . . . . . . . . . . . . . . . . . . . . . . 1217 Florin Miculescu, Lucian Toma Ciocan, Marian Miculescu, Andrei Berbecaru, Josep Oliva, and Raluca Monica Comăneanu

. . . . 1115

Part VI Clinical Performance in Bioresorbable and Load-Bearing Applications: Other Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1247

41

Bioceramics and Composites for Orbital Implants: Current Trends and Clinical Performance . . . . . . . . . . . . . . . . . . . . . . . . . . 1249 Francesco Baino

42

Current Implants Used in Cranioplasty . . . . . . . . . . . . . . . . . . . . . 1275 Dumitru Mohan, Aurel Mohan, Iulian Vasile Antoniac, and Alexandru Vlad Ciurea

43

Marine Biomaterials as Drug Delivery System for Osteoporosis and Bone Tissue Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1309 Joshua Chou and Jia Hao

44

Antibacterial Potential of Nanobioceramics Used as Drug Carriers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1333 T.S. Sampath Kumar and K. Madhumathi

Index . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1375

About the Editor

Iulian Vasile Antoniac Faculty Materials Science and Engineering University Politehnica of Bucharest Bucharest, Romania Professor Dr. habil. Iulian Vasile Antoniac is a materials science engineer working in the field of biomaterials and medical devices. Professor Antoniac is the past World President of the International Society for Ceramics in Medicine (ISCM) and past President of the Romanian Society for Biomaterials (SRB). He completed his Ph.D. in Materials Science and Engineering from University Politehnica of Bucharest (UPB), Romania, in 1998. After several specializations in laboratories of Switzerland, Portugal, France, and the USA on surface analysis, composite materials, implant design, and biomaterials characterizations, his scientific interest spans from the synthesis and characterization of biomaterials and interactions with living tissues, retrieval implant analysis, biodegradable magnesium alloys, ceramic coatings on metallic biomaterials, to the new ceramic composites and scaffolds based on nanostructured and biologically inspired biomaterials for bone regeneration. He received in 2013 his post-doctoral degree in Materials Science and Engineering from UPB, completed his habilitation in 2015 (habilitation thesis was on “orthopedic biomaterials”), and now is the leader of Biomaterials Group from Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Romania. His professional and scientific activity comprises more than 200 international journal papers and conference proceedings, 7 research monographs, and over 10 patents in the areas of biomaterials and their clinical applications that received many international awards at the International Exhibition of Inventions. He was the recipient of the Daniel Bunea Award of the Romanian Society for Biomaterials in 2005 and the Excellence Award of the Romanian Society for Biomaterials in 2012, top individual awards offered by this professional organization. He also acts as a member of the international editorial board and reviewer for many journals and conferences on biomaterials and bioceramics. xv

Contributors

Alessio Adamiano Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Eric Aguado ONIRIS, National Veterinary School of Nantes, Nantes, France Ipek Akin Istanbul Technical University, Department of Metallurgical and Materials Engineering, Maslak, Istanbul, Turkey Camelia Alb Department of Prosthetic Dentistry and Dental Materials, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania Florin Alb Department of Periodontology, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania Madalina Georgiana Albu INCDTP – Division of Leather and Footwear Research Institute, Bucharest, Romania Mauro Alini AO Research Institute Davos, Davos Platz, Switzerland Luca Ambrosio Department of Orthopaedic and Trauma Surgery, Campus Bio-Medico University of Rome, Rome, Italy Luigi Ambrosio Department of Chemical Sciences and Materials Technology, National Research Council of Italy, Rome, Italy Aurora Antoniac Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Bucharest, Romania Iulian Vasile Antoniac University Politehnica of Bucharest, Bucharest, Romania Treena Livingston Arinzeh Department of Biomedical Engineering, New Jersey Institute of Technology, Newark, NJ, USA Barbara Azzimonti Dipartimento di Scienze della Salute, Università del Piemonte Orientale, Novara, Italy Francesco Baino Institute of Materials Physics and Engineering, Applied Science and Technology Department, Politecnico di Torino, Torino, Italy xvii

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Contributors

Domenico Baldi Department of Fixed Prosthodontics, University of Genova, Genova, Italy Besim Ben-Nissan School of Chemistry and Forensic Science, Faculty of Science, University of Technology Sydney, Ultimo, NSW, Australia Andrei Berbecaru Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Bucharest, Romania Aldo R. Boccaccini Institute of Biomaterials, University of Erlangen-Nuremberg, Erlangen, Germany Ilaria Cacciotti Engineering Department, University of Rome “Niccolò Cusano”, Rome, Italy Italian Interuniversity Consortium on Materials Science and Technology (INSTM), Rome, Italy Antonio Carrassi Dipartimento di Scienze Biomediche, Chirurgiche ed Odontoiatriche, Università degli Studi di Milano, Milan, Italy Sophie Cazalbou CIRIMAT Carnot Institute, CNRS-INPT-UPS, Faculty of Pharmacie, University of Toulouse, Toulouse, France Marta Cerruti Department of Mining and Materials Engineering, McGill University, Montreal, QC, Canada J. Chang State Key Laboratory of High Performance Ceramics and Superfine Microstructure, Shanghai Institute of Ceramics, Chinese Academy of Sciences, Shanghai, China Andy H. Choi School of Chemistry and Forensic Science, Faculty of Science, University of Technology Sydney, Ultimo, NSW, Australia Joshua Chou Advanced Tissue Regeneration and Drug Delivery Group, University of Technology Sydney, Sydney, NSW, Australia Lucian Toma Ciocan Dental Medicine Faculty, “Carol Davila” University of Medicine and Pharmacy from Bucharest, Bucharest, Romania Leonardo Ciocca Department of Biomedical and Neuromotor Sciences, University of Bologna, Bologna, Italy Alexandru Vlad Ciurea University of Medicine and Pharmacy “Carol Davila” Bucharest, Bucharest, Romania Andrea Cochis Dipartimento di Scienze della Salute, Università del Piemonte Orientale, Novara, Italy Dipartimento di Scienze Biomediche, Chirurgiche ed Odontoiatriche, Università degli Studi di Milano, Milan, Italy Jacopo Colombo Department of Fixed Prosthodontics, University of Genova, Genova, Italy

Contributors

xix

Raluca Monica Comăneanu Faculty of Dental Medicine, Titu Maiorescu University, Bucharest, Romania Cosmin M. Cotrut Materials Science and Engineering Faculty, University Politehnica of Bucharest, Bucharest, Romania Bogdan Culic Department of Prosthetic Dentistry and Dental Materials, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania Guy Daculsi Dental Faculty, Laboratory for Osteoarticular and Dental Tissue Engineering, INSERM U791, Nantes University, Nantes, France Vincenzo Denaro Department of Orthopaedic and Trauma Surgery, Campus Bio-Medico University of Rome, Rome, Italy Cena Dimova Faculty of Medical Sciences, Dental Medicine, Macedonia FYR, University “Goce Delcev” – Stip, Stip, FYR Macedonia Yaping Ding Institute of Polymer Materials, University of Erlangen-Nuremberg, Erlangen, Germany Sergey V. Dorozhkin Moscow, Russia Diana Dudea Department of Prosthetic Dentistry and Dental Materials, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania Biljana Evrosimovska Faculty of Dentistry, Macedonia, FYR, University “Sts. Cyril and Methody” Skopje, Skopje, FYR Macedonia Mihaela Violeta Ghica Faculty of Pharmacy, University of Medicine and Pharmacy “Carol Davila” Bucharest, Bucharest, Romania Daniela Giugliano Institute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy (CNR), Naples, Italy Gultekin Goller Istanbul Technical University, Department of Metallurgical and Materials Engineering, Maslak, Istanbul, Turkey Dan Grecu Department of Orthopaedics, University of Medicine and Pharmacy Craiova, Craiova, Romania Cristina Gruian Faculty of Physcis and Institute of Interdisciplinary Research in Bio Nano-Sciences, Babes-Bolyai University, Cluj-Napoca, Romania Khadidiatou Guiro Department of Biomedical Engineering, New Jersey Institute of Technology, Newark, NJ, USA Jia Hao Oral Implantalogy and Regenerative Dental Medicine, Tokyo Medical and Dental University, Tokyo, Japan Uli Hauschild Department of Fixed Prosthodontics, University of Genova, Genova, Italy

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Contributors

Larry L. Hench Department of Biomedical Engineering, Florida Institute of Technology, Melbourne, FL, USA Z. G. Huan State Key Laboratory of High Performance Ceramics and Superfine Microstructure, Shanghai Institute of Ceramics, Chinese Academy of Sciences, Shanghai, China Michele Iafisco Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Kunio Ishikawa Department of Biomaterials, Faculty of Dental Science, Kyushu University, Fukuoka, Japan Daniela Ivanov “Petru Poni” Institute of Macromolecular Chemistry, Iasi, Romania Maria G. Katsikogianni Laboratory of Biomechanics and Biomedical Engineering, Department of Mechanical Engineering and Aeronautics, University of Patras, Rion, Patras, Greece Biomaterials and Tissue Engineering Group, School of Dentistry, University of Leeds, Leeds, UK Advanced Materials Engineering, Faculty of Engineering and Informatics, University of Bradford, Bradford, UK Masakazu Kawashita Graduate School of Biomedical Engineering, Tohoku University, Sendai, Japan Dan Laptoiu Department of Orthopaedics and Trauma I, Colentina Clinical Hospital, Bucharest, Romania Isidoro Giorgio Lesci Department of Chemistry “G. Ciamician”, University of Bologna, Bologna, Italy Wei Li Institute of Biomaterials, University of Erlangen-Nuremberg, Erlangen, Germany Giulio Maccauro Medicine and Surgery Department, Clinical Orthopedics and Traumatology Institute, Catholic University, Rome, Italy I. J. Macha School of Chemistry and Forensic Science, Faculty of Science, University of Technology Sydney, Ultimo, NSW, Australia K. Madhumathi Department of Metallurgical and Materials Engineering, Indian Institute of Technology Madras, Chennai, Tamil Nadu, India Hanadi Y. Marghalani Operative Dentistry Department, King Abdulaziz University, Jeddah, Saudi arabia Rodica Marinescu Carol Davila University of Medicine, Bucharest, Romania

Contributors

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Florin Miculescu Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Bucharest, Romania Marian Miculescu Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Bucharest, Romania Thomas Miramond Laboratory for Osteoarticular and Dental Tissue Engineering, INSERM U791, Nantes University, Nantes, France Yannis F. Missirlis Laboratory of Biomechanics and Biomedical Engineering, Department of Mechanical Engineering and Aeronautics, University of Patras, Rion, Patras, Greece Toshiki Miyazaki Graduate School of Life Science and Systems Engineering, Kyushu Institute of Technology, Kitakyushu-shi, Fukuoka, Japan Aurel Mohan University of Oradea, Oradea, Romania Dumitru Mohan University of Oradea, Oradea, Romania Monica Montesi Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Marius Niculescu Department of Orthopaedics and Trauma I, Colentina Clinical Hospital, Bucharest, Romania Faculty of Medicine, Titu Maiorescu University, Bucharest, Romania Chikara Ohtsuki Graduate School of Engineering, Nagoya University, Nagoya, Aichi Prefecture, Japan Josep Oliva Clinica Oliva Dental, Barcelona, Spain Joaquim Miguel Oliveira 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Taipas/Guimarães, Portugal ICVS/3B’s – PT Government Associate Laboratory, Braga/Guimarães, Portugal Dan Pătroi UMF Carol Davila, Department of Fixed prosthodontics and Occlusology, Bucharest, Romania Silvia Panseri Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Alexandru Eugen Petre Department of Prosthodontics, Discipline of Fixed Prosthodontics and Dental Occlusion, University of Medicine and Pharmacy “Carol Davila”, Bucharest, Romania Corrado Piconi Medicine and Surgery Department, Clinical Orthopedics and Traumatology Institute, Catholic University, Rome, Italy

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Contributors

Sandra Pina 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Taipas/Guimarães, Portugal ICVS/3B’s – PT Government Associate Laboratory, Braga/Guimarães, Portugal Alessandro Alan Porporati Medical Products Division, CeramTec GmbH, Plochingen, Germany Sergiu-Alexandru Rădulescu UMF Carol Davila, Department of Fixed prosthodontics and Occlusology, Bucharest, Romania Maria Grazia Raucci Institute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy (CNR), Naples, Italy Rui L. Reis 3B’s Research Group – Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, Taipas/Guimarães, Portugal ICVS/3B’s – PT Government Associate Laboratory, Braga/Guimarães, Portugal Lia Rimondini Dipartimento di Scienze della Salute, Università del Piemonte Orientale, Novara, Italy Judith A. Roether Institute of Polymer Materials, University of ErlangenNuremberg, Erlangen, Germany Norberto Roveri Department of Chemistry “G. Ciamician”, University of Bologna, Bologna, Italy Andrea Ruffini Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Fabrizio Russo Department of Orthopaedic and Trauma Surgery, Campus Bio-Medico University of Rome, Rome, Italy Laura Cristina Rusu University of Medicine and Pharmacy “Victor Babes” Timisoara, Timisoara, Romania T. S. Sampath Kumar Department of Metallurgical and Materials Engineering, Indian Institute of Technology Madras, Chennai, Tamil Nadu, India Monica Sandri Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Marius Sandru Department of Polymer Particles and Surface Chemistry, Sector Biotechnology and Nanomedicine, SINTEF Materials and Chemistry, Trondheim, Norway Dirk W. Schubert Institute of Polymer Materials, University of ErlangenNuremberg, Erlangen, Germany Svetlana Schussler Department of Biomedical Engineering, New Jersey Institute of Technology, Newark, NJ, USA

Contributors

xxiii

Bogdan C. Simionescu Department of Natural and Synthetic Polymers, “Gheorghe Asachi” Technical University, Iaşi, Romania “Petru Poni” Institute of Macromolecular Chemistry, Iasi, Romania Simion Simon Faculty of Physcis and Institute of Interdisciplinary Research in Bio Nano-Sciences, Babes-Bolyai University, Cluj-Napoca, Romania Bogdan Lucian Solomon Royal Adelaide Hospital, Department of Orthopaedics and Trauma, Centre for Orthopaedic and Trauma Research and Discipline of Orthopaedics and Trauma, The University of Adelaide, Adelaide, SA, Australia Marina T. Souza Department of Materials Engineering, Federal University of São Carlos, São Carlos, SP, Brazil Simone Sprio Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Heinz-J€ urgen Steinhoff Department of Physics, University of Osnabruk, Osnabruk, Germany Martin J. Stoddart AO Research Institute Davos, Davos Platz, Switzerland Roman A. Surmenev Department of Experimental Physics, Tomsk Polytechnic University, Tomsk, Russia Maria A. Surmeneva Department of Experimental Physics, Tomsk Polytechnic University, Tomsk, Russia Anna Tampieri Institute of Science and Technology for Ceramics (ISTEC), National Research Council (CNR), Faenza, Italy Teodor Trăistaru UMF Carol Davila, Department of Fixed prosthodontics and Occlusology, Bucharest, Romania Gianluca Vadalà Department of Orthopaedic and Trauma Surgery, Campus Bio-Medico University of Rome, Rome, Italy Emilia Vanea Faculty of Physcis and Institute of Interdisciplinary Research in Bio Nano-Sciences, Babes-Bolyai University, Cluj-Napoca, Romania Elena Varoni Dipartimento di Scienze Biomediche, Chirurgiche ed Odontoiatriche, Università degli Studi di Milano, Milan, Italy George Viscopoleanu Department of Orthopaedic Surgery, “Foisor” Orthopaedic Hospital, Bucharest, Romania Alina Vladescu National Institute for Optoelectronics, Magurele, Romania Stefan Ioan Voicu Faculty of Applied Chemistry and Materials Sciences, University Politehnica of Bucharest, Bucharest, Romania Mishel Weshler Laniado Hospital, Netanya, Israel

xxiv

Contributors

David J. Wood Biomaterials and Tissue Engineering Group, School of Dentistry, University of Leeds, Leeds, UK Xingyi Xie College of Polymer Science and Engineering, Sichuan University, Chengdu, Sichuan, China Julija Zarkova Faculty of Medical Sciences, Dental Medicine, Macedonia FYR, University “Goce Delcev” – Stip, Stip, FYR Macedonia Y. L. Zhou State Key Laboratory of High Performance Ceramics and Superfine Microstructure, Shanghai Institute of Ceramics, Chinese Academy of Sciences, Shanghai, China Katerina Zlatanovska Faculty of Medical Sciences, Dental Medicine, Macedonia FYR, University “Goce Delcev” – Stip, Stip, FYR Macedonia

Part I History and Materials Fundamentals

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History of Development and Use of the Bioceramics and Biocomposites Guy Daculsi

Contents Fabrication and Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Introduction of Macroporosity and Microporosity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Physicochemical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Mechanical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactivity and Osteogenic Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Challenge of Bioactive Bioceramics in Bone Regenerative Medicine . . . . . . . . . . . . . . . . . . . . . CaP Bioceramic Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Bioceramics is a relatively new field; it did not exist until the beginning of 1970, when these materials were shown to restore osteoarticular and dental functions, as well as act as a replacement material for autografts and allograft bone reconstructions. Bioceramics used to replace, repair, or reconstruct human body parts or complex living tissues have differences in their chemical nature, properties, and applications, such as the use of alumina for hip prosthesis versus CaP bioceramics for bone regeneration. Alumina is classified as an inert bioceramic, while CaP bioceramics are considered bioactive biomaterials, able to be absorbed or bond directly with bone. This review concentrates on the development and use of bioceramics and biocomposites and is limited to CaP bioceramics. Bioactive bioceramics are recommended for use as an alternative or additive to autogenous bone for various procedures: orthopedic and dental applications, scaffolds for tissue engineering, vectors for gene therapy, and as a drug delivery system. G. Daculsi (*) Dental Faculty, Laboratory for Osteoarticular and Dental Tissue Engineering, INSERM U791, Nantes University, Nantes, France e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_2

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There are two physical properties of bioceramics that are considered important for optimal biological performance, which includes bioceramic-cell interactions, bioceramic resorption, the bioceramic-tissue interface, and new bone formation. These fundamental properties are interconnecting macroporosity and appropriate microporosity. CaP bioceramics are a recent development in bone surgery that act as a replacement for auto- and allografts, which have been engineered less than 100 years from the first medical applications and less than 30 years from the initial manufacturing of medical devices and experiments with bone regeneration. Bioactive bioceramics have largely contributed to this revolution in medicine. Numerous innovations in this field are now appearing; it is the beginning of bioceramics and not the “has-been medical device.” Keywords

Bioceramics • Calcium phosphate • Physicochemical properties • Biological properties Bioceramics is a relatively new field; it did not exist until the beginning of 1970, when these materials were shown to restore osteoarticular and dental functions, as well as act as a replacement material for autografts and allograft bone reconstructions. Since the turn of the millennium, numerous synthetic bone graft materials have become available as alternatives to autogenous bone for the purposes of repair, substitution, and augmentation. Included in these synthetic biomaterials are special glass ceramics that are described as bioactive glasses and calcium phosphates (CaPs, i.e., calcium hydroxyapatite; HA; tricalcium phosphate, TCP; and biphasic calcium phosphate, BCP). A review [1] has provided an amalgamation of the important historical information about CaP bioceramics. Albee, in 1920, reported the first successful application of a CaP reagent for the repair of a bone defect in a human patient. More than 50 years later, both the clinical use of a TCP preparation in surgically created periodontal defects in animals and the use of dense HA as an immediate replacement for tooth roots were reported. In the early 1980s, synthetic HA and β-TCP became commercially available as substitute bone materials for dental and medical applications [2]. Despite the abundance of literature, books, marketed products, and clinical data, there are still many myths, errors, and critical analyses of results concerning the ability to regenerate or rebuild bone tissue in humans. To address these issues, it is first necessary to recall some definitions: The first definition of “biomaterials” that specifically mentioned “bioceramics” was provided in 1983 by the National Institutes for Health during a conference on the clinical applications of biomaterials in the USA [3]: “A biomaterial is any substance, other than a drug, or combination of substances, synthetic or natural origin, which can be used for any period of time, as a whole or as a part of a system which treats, augments, or replaces any tissue, organ, or function of the body.” In 1986, David Williams and the European Society for Biomaterials [4] significantly simplified the definition of a biomaterial to “a non-viable material used in a medical device intended to interact with [a] biological system.”

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In addition to these general definitions, it was necessary to differentiate between the different types of bioceramics. Bioceramics used to replace, repair, or reconstruct human body parts, or complex living tissues have differences in their chemical nature, properties, and applications, such as the use of alumina for hip prosthesis versus CaP bioceramics for bone regeneration. Alumina is classified as an inert bioceramic, while CaP bioceramics are considered bioactive biomaterials, able to be absorbed or bond directly with bone. This review concentrates on the development and use of bioceramics and biocomposites and is limited to CaP bioceramics. The history from 1770 to 1950 of CaP was recently published [5]. Furthermore, there is already a sizable literature about non-resorbable, bio-inert bioceramics such as alumina or zirconia [6]. There are two pioneers that largely contributed to the use of apatites for biological procedures: Gérard Montel from France who received his PhD in Paris in 1956 [7] and Racquel LeGeros who received hers in 1967 in New York [8]. A symposium of the CNRS “Centre National de La Recherche Scientifique” on biological apatites established the basis of the physical chemistry and crystallography of biological apatites [9]. Elliott [10] confirmed the specificity of biological apatites. These important contributions, in addition to the knowledge about calcium orthophosphate from mineralogy, are the basis for biomineralization during the 16–18 and the subsequent development of CaP bioceramics. In 1988, the first International Bioceramic Symposium was initiated by H. Oonishi in Osaka, and the first volume of Bioceramics was published [11]. This was the origin of the International Society for Ceramics in Medicine, with whom numerous pioneers and past presidents are associated (Besim Ben Nissam, Mario Barbosa, Williams Bonfields, Ian Clarke, Guy Daculsi, Paul Ducheyne, Larry Hench, Kunio Ishikawa, Sabri Kayali, Sukyoung Kim, Tadashi Kokubo, Racquel LeGeros, Panjian Li, Antonio Moroni, Takashi Nakamura, Hajime Ohgushi, Hironobu Oonishi, Marcello Prado, Laurent Sedel, Yli Urpo, Xingdong Zhang, Iulian Antoniac, etc.). The society organizes annual meetings for bioceramics around the world. Twenty-six symposia have been organized; the last one was Bioceramics 26 and took place in Barcelona (Maria Pau Ginebra). The focus of these symposia has evolved over the years, from classic bio-inert bioceramics to resorbable bioceramics and the introduction of scaffolds, composites, and biotechnologies applied to CaP bioceramics. The abbreviations for various CaPs are presented in Table 1. Interestingly, the use of calcium orthophosphate for clinical applications (e.g., bone grafting) largely preceded all of the previous studies on CaP synthesis and characterization, the biological crystals, and the analysis of their specificity. In 1920, Albee reported the first successful use of a CaP reagent for the repair of a bone defect in human [12]. Then, largely through the separate efforts of Jarcho, de Groot, Aoki, and LeGeros [13–17], synthetic HA and β-TCP became commercially available as bone substitution materials for dental and medical applications. From these initial studies on bioceramics, a controversy developed concerning the resorption and clinical efficiency of these materials from the preclinical and first clinical reports. This controversy stemmed from the initial characterization of deficiency; often only the Ca:P ratio was considered, and a complete crystallographic analysis was not

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Table 1 Various calcium phosphates Abbreviation CaP MCPM MCPA OCP DCPD DCPA TetCP β-TCP Mg-whitlockite ACP ACa,Mg,CO3P OHap, or HA

Explanation Any calcium orthophosphate Monocalcium phosphate monohydrate Ca(H2PO4)2 H2O Monocalcium phosphate anhydrous Ca(H2PO4)2 Octacalcium phosphate Ca8H2(PO4)2 5H2O Dicalcium phosphate dihydrate, CaHPO4 2H2O, named also brushite Dicalcium phosphate anhydrous, CaHPO4, named also monetite Tetracalcium phosphate, Ca4(PO4)2O Beta tricalcium phosphate Ca3(PO4)2 β-TCP like with structural HPO42 and Mg2+ ions Amorphous calcium phosphate that gives no X-rays diffraction peak pattern As above but with Mg2+, CO32 (and HCO3) ions contain Hydroxyapatite Ca10(HPO4)6(OH)2

performed. Crystallographic analyses are essential to identify the ionic content of the materials, such as carbonates, magnesium, and other substitutions, which can strongly influence their resorption, interaction with cells, and biocompatibility. More than 50 years after Albee’s report, the clinical use of a TCP preparation in surgically created periodontal defects in animals was reported by Nery et al. [18], and the use of dense HA as immediate root replacements for teeth was reported by Denissen [19]. Characterization of the deficiency was also reported for mixtures of HA and TCP. The term BCP was first used by Ellinger et al. [20] to describe the bioceramic previously known as TCP by Nery et al. in 1975 [21]; in 1986, it was shown by LeGeros to consist of a mixture of 80 % HA and 20 % β-TCP using X-ray diffraction [17]. The first basic studies performed on BCP that varied the HA:β-TCP ratios were reported by LeGeros et al. [22–24]. These studies demonstrated that the bioactivity of these ceramics could be controlled by manipulating the HA:β-TCP ratios. Subsequent studies on BCP [25–29] led to a significant increase in the manufacture and use of commercial BCP bioceramics as bone substitution materials for dental and orthopedic applications [30–38], providing various shapes for blocks and various particles sizes for granules (Fig. 1). The development of CaP bioceramics and other related biomaterials for bone grafts necessitated control of the processes of biomaterial resorption and bone formation at the expense of the biomaterial. Synthetic bone graft materials are now available as alternatives to autogenous bone for repair, substitution, or augmentation. The first patent on HA ceramics was made by Jarcho in 1978 [39] and will be one of the first commercial products used in human “Calcitite™.” The commercial CaP bioceramic products that are currently available are listed on Table 2. The main attractive feature of bioactive bioceramics such as HA, β-TCP, BCP, and substituted, non-stoichiometric HA is their ability to form a strong direct bond with the host bone, resulting in a strong interface compared to bio-inert or biotolerant materials that form a fibrous interface [38, 40–44]. The dynamic interface formed

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Fig. 1 Various shapes of CaP bioceramics

between the bioactive material and host bone is believed to result from a sequence of events involving interactions with the cells, i.e., formation of carbonate HA (CHA; similar to bone mineral) through dissolution/precipitation processes [23, 42, 43, 45, 46]. Commercial CaP bioceramics are sold in Europe, the USA, Brazil, Japan, Korea, Taiwan, and China as bone graft or bone substitute materials for orthopedic and dental applications under various trademarks (Table 2). Bioactive bioceramics are recommended for use as an alternative or additive to autogenous bone for various procedures: orthopedic and dental applications, scaffolds for tissue engineering, vectors for gene therapy, and as a drug delivery system. These commercial bioceramics are available as blocks, particulates (granules), and custom-designed shapes, such as wedges for tibial opening osteotomies, cones for spine and knee applications, and as an insert for vertebral cage fusions. Additionally, they are also available in a particulate form for injectable moldable materials that can be combined with polymers (natural or synthetic) [47].

Fabrication and Properties There are numerous studies describing the production of CaP bioceramics [48–52]. Apatite is considered calcium deficient when the Ca:P ratio is lower than the stoichiometric value of 1.67 for pure calcium hydroxyapatite, Ca10(PO4)6(OH)2. Calcium-deficient apatites (CDAs) may be represented by the formula Ca10-xMx(PO4)6-y(HPO4)y (OH)2, where M represents other ions (e.g., sodium or magnesium) that can be substituted for the calcium (Ca) ions. Calcium deficiency in apatites depends on the synthesis conditions (precipitation or hydrolysis methods), reaction pH, and temperature [38, 42, 53, 54].

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Table 2 Commercially available products Calcium orthophosphate HA

Bovine bone apatite (unsintered)

Bovine bone apatite (sintered)

β-TCP

Trade name and producer Cementek (Teknimed, France) Osteogen (Impladent, NY, USA) Actifuse (ApaTech, UK) Apaceram (Pentax, Japan) ApaPore (ApaTech, UK) Bioroc (Depuy-Bioland, France) Bonefil (Pentax, Japan) BoneMedik (MetaBiomed, China) Bonetite (Pentax, Japan) Boneceram (Sumitomo Osaka Cement, Japan) BoneSource (Stryker Orthopaedics, NJ, USA) Calcitite (Zimmer, IN, USA) Cerapatite (Ceraver, France) HAP91 (JHS Biomaterials, Brazil) Neobone (Toshiba Ceramics, Japan) Ostegraf (Ceramed, CO, USA) Ostim (Heraeus Kulzer, Germany) Synatite (SBM, France) BioOss (Geitslich, Switzerland) Laddec (Ost-Developpement, France) Lubboc (Ost-Developpement, France) Oxbone (Bioland biomateriaux, France) Tutoplast (IOP, CA, USA) BonAP Cerabone (aap Implantate, Germany) Endobon (Merck, Germany) Osteograf (Ceramed, CO, USA) PepGen P-15 (Dentsply Friadent, Germany) Bioresorb (Sybron Implant Solutions, Germany) Biosorb (SBM S.A., France) Calciresorb (Ceraver, France) Cerasorb (Curasan, Germany) Ceros (Thommen Medical, Switzerland) ChronOS (Synthes, PA, USA) Conduit (DePuy Spine, USA) Granulado (Keramat, Spain) JAX (Smith and Nephew Orthopaedics, USA) Osferion (Olympus Terumo Biomaterials, Japan) OsSatura TCP (Integra Orthobiologics, CA, USA) Syncera (Oscotec, Korea) Vitoss (Orthovita, PA, USA) (continued)

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Table 2 (continued) Calcium orthophosphate BCP (HA + β-TCP)

BCP (HA + α-TCP) FA + BCP (HA + β-TCP) Carbonateapatite Coralline HA

Trade name and producer 4Bone (MIS, Israel) Atlantik (MedicalBiomat, France) BCP Bicalphos (Medtronic, MN, USA) Biosel (Depuy Bioland, France) BonePlus (Mega’Gen, Korea) BoneSave (Stryker Orthopaedics, NJ, USA) BonitMatrix (Dot, Germany) Calciresorb (Ceraver, France) CellCeram (Scaffdex, Finland) Ceraform (Teknimed, France) Ceratite (NGK Spark Plug, Japan) Cross-Bone (Biotech International, France) Eliz (Kyeron, The Netherlands) Eurocer (FH Orthopedics, France) Graftys BCP (Graftys, France) Hatric (Arthrex, Naples, FL, USA) Indost (Polystom, Russia) Kainos (Signus, Germany) MBCP, MBCP+ (Biomatlante, France) OptiMX (Exactech, USA) OsSatura BCP (Integra Orthobiologics, CA, USA) OssGen (OssGen Inc, Korea) Osteosynt (Einco, Brazil) SBS (Expanscience, France) Sinbone HT (Purzer, Taïwan) TCH (Kasios, France) Triosite (Zimmer, IN, USA) Tribone (Stryker, Europe) XPand (XPand Bioetch, The Netherlands) Skelite (Millennium Biologix, ON, Canada) FtAP (Polystom, Russia) Healos (Orquest, CA, USA) Interpore (Interpore, CA, USA) ProOsteon (Interpore, CA, USA)

Currently, the precipitation of HA, β-TCP, and formation of HA/TCP mixtures (BCP) are performed using CDA. CDA is obtained by precipitation under various pH and temperature conditions: the lower the pH, the higher the temperature required for the precipitation of apatite [38, 42]. For example, CDA can be produced by precipitation at 80–100  C at low pH (pH 4–6). At lower temperatures and lower pH, non-apatitic CaPs, e.g., dicalcium phosphate dihydrate (DCPD; CaHPO4  2H2O) or octacalcium phosphate (OCP; Ca8H2(PO4)6  5H2O), are obtained [42].

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Another method of CDA preparation is hydrolysis of non-apatitic CaPs, including amorphous CaP, Cax(PO4,HPO4)y; DCPD; dicalcium phosphate anhydrous, CaHPO4; OCP; and β-TCP, Ca3(PO4)2 [38, 42, 54–56]. During hydrolysis of DCPD in NaOH solutions, the calcium deficiency of the unsintered apatite and the subsequent HA:β-TCP ratio of the BCP produced after sintering can be controlled by two variables: the concentration of the NaOH solution and the ratio of the weight of the DCPD to the volume of the NaOH solution [54]. After creation of the raw powder, all of the processes used to manufacture bioceramics require high temperatures, in excess of 1000 . Under this condition, there is thermal decomposition of CDA, HA, and other CaP compounds that can affect their chemical nature, biological properties (dissolution), mechanical properties (grain boundaries and lattice defects), and biocompatibility (e.g., CaO content) [57].

Introduction of Macroporosity and Microporosity In 1971, Klawitter and Hulbert [58] demonstrated the importance of porosity for bone graft substitutions. Porous ceramics are potentially attractive because of their large interconnecting pore structure, which facilitates bone ingrowth, cell colonization, and new vascularization. There are numerous publications that describe the processes for creation of porous bioceramics and how to determine the optimal pore size. Indeed, from these studies, we now know that the ideal pore size for tissue growth is 300–600 μm [59]. Other studies have demonstrated that with the same porous structure, β-TCP generally degrades more rapidly than ceramics made of HA [60]. To understand the macroporosity of the coral skeleton that was used, the aragonite or calcite were thermally decomposed into CaPs in the presence of an alkaline phosphate solution. Other methods to elucidate the macroporosity include release with hydrogen peroxide, the release of oxygen during heating involving a macropore, and, more classically, the use of a porogen (naphthalene, polyethylene, camphor, poly(methyl methacrylate), polyurethane, sugar, starch, etc.). After mixing the porogen with the powder, or a template with a slurry, during heating, the organic phase is volatilized or oxidized, leaving calibrated spaces behind. During the first 10 years of bioceramic development, studies focused on the chemical nature (HA or TCP) and macroporosity of the materials. Indeed, the biological properties of the bioceramic – degradation-dissolution, absorption, and biological interactions (with the fluid, cells, and tissues) – depend on the micro(microscale) and macrostructure (macropores, particle size, and blocks) of the material. There are two physical properties of bioceramics that are considered important for optimal biological performance, which includes bioceramic-cell interactions, bioceramic resorption, the bioceramic-tissue interface, and new bone formation. These fundamental properties are interconnecting macroporosity and appropriate microporosity [38, 61, 62]. Macroporosity is introduced into the CaP ceramic through incorporation of volatile materials (e.g., naphthalene, hydrogen peroxide,

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History of Development and Use of the Bioceramics and Biocomposites

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Fig. 2 Macroporosity in CaP bioceramics (biphasic calcium phosphate MBCP)

Fig. 3 BCP sintered at 1050  C (a) and 1200  C (b)

or other porogens [polymers, sugar, starch, etc.]), heating at temperatures that induce sublimation or calcination (80–500  C), and subsequent sintering at higher temperatures [14, 61] (Fig. 2). Microporosity is a consequence of the temperature and duration of sintering: the higher the temperature, the lower the microporosity and the lower the specific surface area (Fig. 3a, b). From our own experience, it appears that micropores enhance the diffusion and permeability of the bioceramic, increasing the interconnections of the macropores. Only a few publications have investigated the importance of microporosity; however, some of the previous studies appear to show that the osteoinductive

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properties of bioactive ceramics are due to their micropore content [63]. Other reports demonstrated that the higher osteogenic properties in bony sites with equivalent macroporosity have a higher content of micropores [64, 65]. Analysis of the currently available commercial CaP bioceramics revealed that most materials have a porosity of 70 %, although large differences can be observed among the registered bioceramics. The total porosity (macroporosity plus microporosity) of these products is reported to be approximately 70 % of the bioceramic volume. However, the ratios of macroporosity and microporosity are often different, with the percent microporosity varying from 3 % to 25 %. Low percent microporosity and low surface area can result in lower bioactivity and lower dissolution properties.

Physicochemical Properties Because β-TCP has a higher solubility than HA [66], the extent of dissolution of bioceramics with comparable macroporosity and particle size will depend on the HA:β-TCP content: increased HA leads to less dissolution. Dissolution is also affected by the method of synthesis and sintering of the material. The specific surface area, microporosity, mechanical properties, and method of particle release can influence the degradation-dissolution and absorption properties of the bioceramic by the surrounding cells and tissue. This phenomenon may be caused by processing variables (sintering time and temperature) that could affect the total macroporosity and microporosity: the greater the macroporosity and microporosity, the greater the extent of dissolution. In vivo, dissolution of bioceramics is accompanied by a decrease in crystal size and increase in macroporosity and microporosity. While the influences of crystal size, microporosity, and pore content on bioceramic properties have been studied, few papers have focused on the atomic and nanoscale structures of these materials. Crystal defects, grain boundaries, and lattice defects have largely only been studied in minerals, metals, and ceramics [6]. Biomineralization has been investigated using high-resolution transmission electron microscopy (TEM) [67, 68]. In 1991 [69], one report revealed the different kinds of lattice defects in CaP ceramics at the unit cell level (Fig. 4). The authors found that the ultrastructural properties depended on the sintering temperature. Numerous lattice defects, including atomic vacancies, dislocations, and two types of grain boundaries, were described. In another study, a new type of threedimensional lattice defect was observed at the nanoscale level in an HA bioceramic sintered at low temperature (900  C). This defect involved intrinsic nanopores within the material (Fig. 5). The defects in the nanostructure of bioceramic crystals have an important role in the dissolution process and demonstrate that all bioceramics are soluble (both HA and TCP), in contrast to claims about the non-resorbability of HA often reported in the literature. The extent of bioceramic dissolution depends on the number of defects [70] and cell mediation (Fig. 6). Based on these data, the notion that HA bioceramics are “non-resorbable” should be revisited and revised.

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History of Development and Use of the Bioceramics and Biocomposites

Fig. 4 High-resolution transmission electron microscopy of HA(*) and TCP phase ( ), with grain boundaries (arrow) associating the two phases in BCP at the molecular level

Fig. 5 Unit cell and 3-D lattice defects involving nanoporosity in low-temperature sintered HA bioceramic

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Fig. 6 Macrophages biological dissolution of HA bioceramic crystal (Calcitite™), phagolysosome containing HA crystals with dissolution areas (arrow)

Mechanical Properties The pore size and percent macroporosity of BCP ceramics are believed to affect their mechanical properties [70]. The preparation method was also found to have a significant influence on the compressive strength of the material. The compressive strength reported for bioactive bioceramics varies from 0.5 to 12 MPa. However, the mechanical properties by themselves are not the best criteria for predicting the efficacy of bone ingrowth. For example, BCP with high mechanical properties and low microporosity (resulting from high sintering temperatures) can exhibit reduced bioresorption and bioactivity. However, the initial mechanical properties can increase 2–3  (2–6 MPa) a few weeks after implantation due to the physicochemical events of dissolution and biological precipitation into the micropores [29, 71].

Bioactivity and Osteogenic Properties Bioactivity is the ability of a material to form CHA on its surface in vitro [40, 72] or in vivo [5, 73–75]. Osteoinductivity is the ability of a material to induce bone formation de novo or ectopically (in non-bone-forming sites). CaP bioceramics do not usually display osteoinductive properties. However, several reports have shown osteoinductive properties in some CaP bioceramics, such as coralline HA (derived from coral), BCP [76–78], and TCP [79, 80]. Reddi [81] hypothesized that these apparent osteoinductive properties are the ability of particular ceramics to concentrate bone growth factors circulating in the biological fluids; these growth factors can then induce bone formation. Ripamonti [76] and Kuboki [77] independently postulated that the geometry of the material is a critical parameter in bone formation. Others have

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Fig. 7 Biological apatite precipitation (arrow) on residual calcium phosphate crystals after implantation. (high-resolution transmission microscopy)

speculated that low oxygen tension in the central region of the implants might provoke dedifferentiation of pericytes from blood microvessels into osteoblasts [82]. It has been also suggested that the nanostructured rough surface or the surface charge of implants might cause asymmetrical division of stem cells into osteoblasts [79]. The surface microstructure appears to be a common property of biomaterials that can induce ectopic bone formation. Recent studies have indicated the critical role of micropores during ceramic-induced osteogenesis. For example, it was reported that bone formation occurred in the muscle of dogs inside porous CaP ceramics that displayed surface microporosity; however, osteogenesis was not observed inside macroporous ceramics with a dense surface [80]. It was also reported that metal implants coated with a microporous layer of OCP could induce ectopic bone formation in the muscle of goats, while a smooth layer of carbonated apatite on porous metal implants was not able to induce osteogenesis [83]. In all of the previous experiments, ectopic bone formation occurred inside macroporous ceramic blocks. Therefore, ceramic properties such as composition, geometry, porosity, size, and microstructure are critical parameters for bone formation. The main explanations for the osteoinductive properties of CaP bioceramics appear to be the formation of microcrystals with Ca:P ratios similar to those of bone apatite crystals observed after implantation, interaction with the biological fluids and cells, and the combined effects of integration of non-collagenic proteins involved in cell adhesion, stem cell differentiation in an osteogenic line, and biomineralization. Using high-resolution TEM, it was demonstrated that the formation of these bone apatite-like microcrystals (Fig. 7) that appeared after implantation of CaP (HA, TCP, and BCP) were nonspecific, i.e., not related to the implantation site (osseous or non-osseous sites), subject of implantation, or type of CaP bioceramic [24, 37].

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The Challenge of Bioactive Bioceramics in Bone Regenerative Medicine Tissue engineering for bone regeneration involves the seeding of osteogenic cells (e.g., mesenchymal stem cells, MSCs) onto the appropriate scaffolds and subsequent implantation of the seeded scaffolds into the bone defect. Bone marrow-derived MSCs are multipotent cells that are capable of differentiating into, at a minimum, osteoblasts, chondrocytes, adipocytes, tenocytes, and myoblasts [84–89]. MSCs can be isolated from a small volume of bone marrow and cultured because of their proliferative capacity; they maintain their functionality after expansion in culture and cryopreservation [87]. Therefore, MSCs are a readily available and abundant source of cells for tissue engineering applications. Bioactive CaP ceramics, in combination with cells and tissue engineering derivatives, are the new challenge in bone regeneration.

CaP Bioceramic Composites There are numerous clinical situations where materials are needed to restore and regenerate bone, and the best material for this process is an autologous bone graft. However, there are numerous limitations with autologous bone grafts that have described in the literature [90–92]. The ideal synthetic biomaterial should be injectable and moldable, set in the defect, and favor bone apposition and growth while being degraded by body fluids and cells. Ultimately, the material should be replaced by mature bone tissue within a healing period of weeks [7]. Composite biomaterials for bone repair have been developed in the last 30 years. A composite material is composed of at least two or more phases: a continuous phase (the matrix consisting of polymers) and a dispersed phase (consisting of granules, fibers, etc.). There should be an interaction between the two phases, with a specific interface that plays an important role in the mechanical performance, environmental stability, handling, rheological properties, and biological performance of the material [93]. In an effort to make generic CaP bioceramics more biomimetic to bone, as well as support new minimally invasive surgical techniques, multiphasic biomaterials were developed that associated inorganic solid state chemistry (CaP bioceramics) with inorganic polymers (hydrogel and polymers of synthetic or natural origin) to replace and regenerate bone tissue in osseous or dental defects, or for resorbable osteosynthesis; some of these association of bioceramic and polymers are presented in Table 3. This multidisciplinary research involves solid state chemistry with CaP bioceramics, organic chemistry with synthetic polymers, and biochemistry with polymers of natural origin. Several composites have been created using collagen, fibrin, silk, or various synthetic polymers (polysaccharides, polycaprolactone, poloxamer, poly(ethylene glycol), etc.) [94–97]. In a composite biomaterial, the putty associates the various bioactive properties of the bioceramic and provides numerous advantages for the handling, surgical indication, and kinetics of bone regeneration. For resorbable osteosynthesis, essentially the polyesthers (PLA and GLA, and derivatives, copolymers) were used in

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Table 3 Example of some composites on the European or US market HA/collagen

HA/sodium alginate HA/Poly-L-lactic acid TCP/PLLA, PLA BCP/PLDLLA HA/polyethylene HA/CaSO4 Algae-derived HA BCP/collagen BCP/fibrin BCP/silicon HA Si/poloxamer BCP/polysacharide

Bioimplant (Connectbiopharm, Russia) Bonject (Koken, Japan) CollapAn (Intermedapatite, Russia) HAPCOL (Polystom, Russia) LitAr (LitbicAr, Russia) Bialgin (Biomed, Russia) SuperFIXSORB30 (Takiron, Japan) Bilock (Biocomposite, USA) Bioscrew (Conmed, USA) Osteotwin (Biomatlante, France) HAPEX (Gyrus, TN, USA) Hapset (LifeCore, MIN, USA) Algipore (Dentsply Friadent, Germany) Collagraft (Zimmer, IN, USA) Matribone (Biom’Up, France) TricOS (Baxter BioScience, France) FlexHA (Xomed, FL, USA) Actifuse Apatech-Baxter Inn’Os (Biomatlante, France)

combination with a mineral content. Numerous resorbable osteosynthesis, plaque, nails, screws, spine cages, etc. were used other the world through the pioneering efforts of Michel Vert particularly [98].

Conclusion CaP bioceramics are a recent development in bone surgery that act as a replacement for auto- and allografts, which have been engineered less than 100 years from the first medical applications and less than 30 years from the initial manufacturing of medical devices and experiments with bone regeneration. In 2014 during an ESB meeting in Liverpool, Larry Hench reported four points in the history of healthcare revolutions: 1. Prevention of loss of life (1850–1950; clean water, antiseptics, antibiotics, immunization, and sterilization) 2. Replacement of tissues (1950–present; antiseptics, antibiotics, sterilization, and biomaterials) 3. Regeneration of tissues (1980–present; tissue engineering and in situ gene activation) 4. Prevention of tissue loss (2000–present) Bioactive bioceramics have largely contributed to this revolution in medicine. Numerous innovations in this field are now appearing; it is the beginning of bioceramics and not the “has-been medical device.”

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Bioactive Glass Bone Grafts: History and Clinical Applications Larry L. Hench

Contents History of Bioactive Glasses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Bone Grafts: Regeneration of Tissues . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Glass Bone Grafts: Genetic Control of Tissue Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Glass Bone Grafts: Clinical Applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Implications for the Future . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

The second generation of biomaterials concept, based upon a unique composition range of calcium, sodium phosphosilicate (CSPS) glasses, and glass-ceramics, involved bonding of the implant material to the bone. It is essential to recognize that no man-made material can respond to changing physiological loads or biochemical stimuli, as do living tissues. This compromise limits the lifetime of all man-made body parts. Recognizing this fundamental limitation also signals that we have reached a limit to current medical practice that emphasizes replacement of tissues. For the twenty-first century, it is critical to emphasize a more biologically based method of repair-regeneration of tissues. Third-generation bioactive materials with controlled release of biochemical stimuli provide the starting point for this shift toward a more biologically based approach to repair of diseased or damaged tissues, e.g., regeneration of tissues. Bioactive glass bone grafts are based upon this concept of in situ regeneration of bone with structure, architecture, and mechanical strength equivalent to normal cortical and cancellous bone. We need to remember that only a little more than 40 years ago, the Larry L. Hench: deceased L.L. Hench (*) Department of Biomedical Engineering, Florida Institute of Technology, Melbourne, FL, USA e-mail: [email protected]; [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_5

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concept of a material that would not be rejected by living tissues also seemed impossible. This is now a clinical reality that has benefited tens of millions of people and should stimulate new concepts in the years ahead. Keywords

Bioactive glasses • Bone grafts • Tissue regeneration • Regenerative medicine

History of Bioactive Glasses In 1969 a second generation of biomaterials was discovered for use in repairing and replacing diseased or damaged parts of the body [1–5]. This new concept, based upon a unique composition range of calcium, sodium phosphosilicate (CSPS) glasses, and glass-ceramics, involved bonding of the implant material to the bone. Such materials were called bioactive. Table 1 summarizes the composition of bioactive glasses and glass-ceramics and their level of bioactivity. The glasses with highest levels of bioactivity, class A, bond to both bone and soft connective tissues by means of osteostimulation and osteoconduction. Slower bonding materials, class B bioactivity, bond only to bone by a process of osteoconduction. Discussions of classes of bioactivity are presented in other reviews. Bioactive materials elicit a controlled action and reaction in the physiological environment. The mechanism of bonding of bioactive glasses (composed of Na2OCaO-P2O5-SiO2) to living tissue, published in 1971 [3], involves 11 reaction steps [4]. As illustrated in Fig. 1, the first five steps occur on the surface of 45S5 Bioglass. The reactions begin by rapid ion exchange of Na+ with H+ and H3O+ from the surrounding physiological solutions in contact with the surface of the bioactive glass. The ion exchange is followed by a polycondensation reaction of a very large concentration of surface silanols (Si-OH HO-Si) to create a very high surface area silica (SiO2) gel, which provides a large number of surface reactive sites for heterogeneous

Table 1 Composition and properties of bioactive glasses and glass-ceramics used clinically for medical and dental applications Composition (wt%) Na2O CaO CaF2 MgO P2O5 SiO2 Phases

45S5 Bioglass (NovaBone) 24.5 24.5 0 0 6 45 Glass

S53P4 (AbminDent1) 23 20 0 0 4 53 Glass

Class of bioactivity

A*

A

A/W glass-ceramic (cerabone) 0 44.7 0.5 4.6 16.2 34 Apatite Beta-wollastonite Glass B

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Fig. 1 Sequence and rates of reaction stages to grow new bone at the interface with 45S5 bioactive glass

nucleation and crystallization of a biologically reactive hydroxyl-carbonate apatite (HCA) layer equivalent to the inorganic mineral phase of bone. The growing HCA layer on the surface of the material is an ideal environment for six cellular reaction stages, shown in sequence in Fig. 1. The cellular mechanisms include colonization by osteoblast stem cells (stage 8) followed by proliferation (stage 9) and differentiation (stages 10 and 11) of the cells to form new bone that has a mechanically strong bond to the implant surface. The sequence of cellular stages only occurs on the bioactive glass surface if the preceding five stages of surface reactions have progressed rapidly, within minutes to hours, as indicated by the logarithmic time scale in Fig. 1. If the surface reacts too slowly, the cellular stages are delayed and the material bonds only slowly to bone via osteoconduction along the interface of the material and host bone. By the mid-1980s bioactive materials had reached clinical use in numerous orthopedic and dental applications [4]. Synthetic hydroxyapatite (HA) ceramics had begun to be routinely used as porous implants, powders, and coatings on metallic prostheses to provide bioactive fixation [5, 6]. The presence of sparingly soluble HA coatings led to a tissue response (termed osteoconduction) where bone grew along the coating and formed a mechanically strong interface [5, 6]. Bioactive glasses and

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glass-ceramics, based upon the original 45S5 Bioglass® formulation [3], were being used as middle ear prostheses to restore the ossicular chain and treat conductive hearing loss and as endosseous ridge maintenance implants to preserve the alveolar ridge from the bone resorption that follows tooth extraction [7]. The mechanically strong and tough bioactive A/W glass-ceramic, developed at Kyoto University, was used for replacement of vertebrae in patients with spinal tumors [8]. In 1998 a centennial feature article of the American Ceramic Society documented the rapid growth of clinical use of first- and second-generation bioceramics [4].

Bioactive Bone Grafts: Regeneration of Tissues The clinical success of first-generation, bio-inert, and second-generation bioactive and resorbable implants has met many of the medical needs of a rapidly aging population. However, survivability analyses of most prostheses [7, 9] show that a third to half of medical devices fail within 15–25 years. Failures require patients to have revision surgery that is costly to the patients and to society and comprises a significant contribution to the rapidly rising costs of healthcare. Thirty years of research has had relatively small effects on failure rates of medical devices made from first- and second-generation biomaterials [7]. For decades the approach to healthcare has been based upon trial and error experiments that require use of many animals and large numbers of human clinical trials. This approach needs to be replaced with a more affordable and more reliable alternative for the younger, 40–70 years old, patients. Improvements of either firstor second-generation biomaterials are limited in part because “All man-made biomaterials used for repair or restoration of the body represents a compromise” [1, 4]. It is essential to recognize that no man-made material can respond to changing physiological loads or biochemical stimuli, as do living tissues. This compromise limits the lifetime of all man-made body parts. Recognizing this fundamental limitation also signals that we have reached a limit to current medical practice that emphasizes replacement of tissues. For the twenty-first century, it is critical to emphasize a more biologically based method of repair-regeneration of tissues. Third-generation bioactive materials with controlled release of biochemical stimuli provide the starting point for this shift toward a more biologically based approach to repair of diseased or damaged tissues, e.g., regeneration of tissues [10]. Bioactive glass bone grafts are based upon this concept of in situ regeneration of bone with structure, architecture, and mechanical strength equivalent to normal cortical and cancellous bone.

Bioactive Glass Bone Grafts: Genetic Control of Tissue Regeneration Third-generation biomaterials, such as bioactive glass particulates of 45S5 Bioglass, are designed to stimulate specific cellular responses at the level of molecular biology [10]. During the first decade of the twenty-first century, the concepts of bioactive

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materials and resorbable materials converged; third-generation bioactive glasses and hierarchical porous foams are being designed to activate genes that stimulate regeneration of living tissues. The key design principle of such third-generation materials is to control the rate of release of specific ionic stimuli that activate or regulate the function of genes in progenitor cells. In Situ Tissue Regeneration. This approach involves the use of biomaterials in the form of powders, solutions, or doped micro- or nanoparticles to stimulate local tissue repair [11, 12]. Certain formulations of bioactive materials release chemicals in the form of ionic dissolution products at controlled rates that activate the cells in contact with the stimuli. The cells produce additional growth factors that in turn stimulate multiple generations of growing cells to self-assemble into the required tissues in situ, along the biochemical and biomechanical gradients that are present. The advantage offered by regenerative medicine is genetic control of the tissue repair process. The result is equivalent to repaired natural tissue in that the new structure is living and adaptable to the physiological environment. There is growing evidence to support the hypothesis governing design of thirdgeneration biomaterials, i.e., generation of specific cell responses to controlled release of biochemical stimuli. For example, when a particulate of bioactive glass is used to fill a bone defect, there is rapid regeneration of bone that matches the architecture and mechanical properties of bone in the site of repair. Both osteoconduction and osteoproduction [13] occur as a consequence of rapid reactions on a bioactive glass surface [4]. The surface reactions release critical concentrations of soluble Si, Ca, P, and Na ions that give rise to both intracellular and extracellular responses at the interface of the glass with its cellular environment. Attachment and synchronized proliferation and differentiation of osteoblasts rapidly occur on the surface of bioactive materials [14]. Osteoprogenitor cells capable of forming new bone colonize the surface of highly bioactive materials. Slow release of soluble ions from the material stimulates cell division and production of growth factors and extracellular matrix proteins. Mineralization of the matrix follows and the mature osteoblast phenotype, encased in a collagen-HCA matrix (osteocytes), is the final product both in vitro and in vivo [13–24]. Numerous studies have established that there is genetic control of the cellular response to the most reactive of the bioactive glasses (45S5 Bioglass). Seven families of genes are upregulated when primary human osteoblasts are exposed to the ionic dissolution products of bioactive glasses [17–24]. The gene expression occurs within 48 h and includes enhanced expression by more than twofold of the families of genes listed in Table 2. See Xynos et al. for a listing of the genes and the extent of their upregulation [17]. The upregulated genes encode nuclear transcription factors and cell cycle regulators. Potent growth factors, especially insulin-like growth factor II (IGF-II), were increased by 3.2-fold along with IGF binding proteins and proteases that cleave IGF-II from their binding proteins. Similar bioactive induction of the transcription of at least five extracellular matrix components (2–3.7-fold) and their secretion and

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Table 2 Families of genes upregulated or activated by ionic dissolution products from bioactive glass 1. 2. 3. 4. 5. 6. 7.

Transcription factors and cell cycle regulators Signal transduction molecules Proteins in DNA synthesis, repair, recombination Growth factors and cytokines Cell surface antigens and receptors Extracellular matrix components Apoptosis regulators

2 to 5 fold 2 to 6 fold 2 to 3 fold 2–3.2-fold 2 to 7 fold 2–3.7-fold 1.6–4.5-fold

self-organization into a mineralized matrix is responsible for the rapid formation and growth of bone nodules and differentiation of the mature osteocyte phenotype [14–16, 24]. Studies have confirmed the results of the early Xynos et.al. findings and extended the generality to include several types of precursor cells and differing sources of biologically active Ca and Si ionic stimuli [14–17]. Bone biology and gene array analyses of five different in vitro models using four different sources of inorganic ions provide the experimental evidence for a genetic theory of osteogenic stimulation [17–25]. All experiments showed enhanced proliferation and differentiation of osteoblasts toward a mature, mineralizing phenotype without the presence of any added bone growth proteins, such as bone morphogenetic proteins (BMPs). Shifts in osteoblast cell cycles were observed as early as 6 h for most experiments, with elimination (by apoptosis) of cells incapable of differentiation [14]. The remaining cells exhibited enhanced synthesis and mitosis. The cells quickly committed to generation of extracellular matrix (ECM) proteins and mineralization of the matrix. Gene array analyses showed early upregulation or activation of seven families of genes (Table 2) that favored both proliferation and differentiation of the mature osteoblast phenotypes, including: transcription factors and cell cycle regulators (six with increases of two- to fivefold); apoptosis regulators (three at 1.6–4.5-fold increase); DNA synthesis, repair, and recombination (four at two- to threefold); growth factors (four at two- to threefold) including IGF-I1 and VEG F; cell surface antigens and receptors (four at two- to sevenfold, especially CD44); signal transduction molecules (three at two- to sixfold); and ECM compounds (five at 2–3.7fold).

Bioactive Glass Bone Grafts: Clinical Applications Figure 2 summarizes the time line for development of clinical products of bioactive glass from the date of the first discovery of 45S5 Bioglass in November 1969. The first products approved by FDA in the mid-1980s were bulk implants cast from molten 45S5 Bioglass: the MEP (middle ear prostheses) for replacement of diseased or damaged or missing ossicles and the ERMI (endosseous ridge maintenance

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1969 1971 1972 1975 1981 1981 1985 1987 1988 1993

1995 1996 1999 2000 2000 2000 2001 2004 2011

29

Discovery of bone bonding to 45S5 Bioglass at University of Florida First peer reviewed publications of bonding of bone to bioactive glasses and glass-ceramics [1–3] Bonding of Bioglass bone segments and coated femoral stems in monkeys Bioglass dental implants bonded in baboon jaws Discovery of soft connective tissue bonding to 45S5 Bioglass Toxicology and biocompatibility studies (20 in vitroand in vivo) published to establish safety for FDA clearance of Bioglass products First medical product (Bioglass Ossicular Reconstruction Prosthesis) (MEP) cleared by FDA via the 510(k) process Discovery of osteoproduction (osteostimulation) in use of Bioglass particulate in repair of periodontal defects Bioglass Endosseous Ridge Maintenance Implant (ERMI) cleared by FDA via the 510 (k) process Bioglass particulate for use in bone grafting to restore bone loss from periodontal disease in infrabony defects (Perioglas) cleared by FDA via the 510 (k) process Perioglas obtained CE Mark in Europe Use of Perioglas for bone grafts in tooth extraction sites and alveolar ridge augmentation cleared by FDA via the 510 (k) process European use of 45S5 particulate for orthopedic bone grafting (NovaBone) FDA clearance for use of NovaBone in general orthopedic bone grafting in nonload bearing sites Quantitative comparison of rate of trabecular bone formation in presence of Bioglass granules versus synthetic HA and A/W glass-ceramic Analysis of use of 45S5 Bioglass ionic dissolution products to control osteoblast cell cycles Gene expression profiling of 45S5 Bioglass ionic dissolution products to enhance osteogenesis FDA clearance of 45S5 particulate for use in dentinal hypersensitivity treatment (NovaMin) Acquisition of NovaMin technology by Glaxo-Smith-Kline and world launch of Sensodyne Repair and Protect toothpaste for prevention of dentinal hypersensitivity and gingivitis

Fig. 2 Chronology of science and clinical product development of 45S5 Bioglass

implant) for replacement of the roots of teeth for patients wearing dentures. The clinical results of the second-generation bulk 45S5 implants were outstanding with >90 % success over extended time periods. While the second-generation Bioglass® materials performed admirably in replacing diseased or missing hard tissue, the discoveries that Bioglass® could positively affect osteoblasts and, in fact, “stimulate” them to produce more bone tissue earlier than other synthetic biomaterials led to the concept of “osteoproduction” and “osteostimulation” [13]. In order to take advantage of this property, and of the need to regenerate diseased or missing tissues, the development of third-generation Bioglass® products focused on using particles rather than monolithic shapes [2]. The first NovaBone ® particulate material cleared for sale in the USA was PerioGlas ®, which was cleared via the 510(k) process in December 1993 [2].

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In 1995, PerioGlas® obtained a CE mark and marketing of the product began in Europe. The initial indication for the product was to restore bone loss resulting from periodontal disease in infrabony defects. In 1996, additional indications for use were cleared by FDA, including use in tooth extraction sites and for alveolar ridge augmentation [2]. The first paper to describe potential use of 45S5 Bioglass ® particulate in repair of periodontal defects was published in 1987 by Dr. June Wilson and Professor Sam Low, Department of Periodontology, and colleagues at the University of Florida [13]. A detailed study of the monkey model and clinical results was documented many years ago and published in previous papers [2, 4]. Over its nearly 20-year clinical history, PerioGlas® has demonstrated excellent clinical results with virtually no adverse reactions to the product. Numerous clinical studies have demonstrated the efficacy of the product in multiple uses [2]. To date, PerioGlas® is sold in over 35 countries, and the manufacturer estimates that the product has been used in more than one million surgeries (Data on file at NovaBone Corporation, Alachua, Florida, USA). Building on the successes of PerioGlas® in the market, a Bioglass ® particulate for orthopedic bone grafting was introduced into the European market in 1999, under the trade name NovaBone ® [2]. The product was cleared for general orthopedic bone grafting in non-load-bearing sites in February 2000. The material 45S5 Bioglass is now widely used in many types of orthopedic and dental applications as shown in Table 3. The osteoblast cell culture results reviewed above correlate with clinical results using the same bioactive material, 45S5 Bioglass [2, 25, 26]. An especially important finding is clinical equivalence of results from synthetic bioactive glass bone grafts of the composition 45S5 to use of autogenous grafts in the same clinical applications. Clinical studies that compare the success of autogenous bone grafts versus grafts of the gene-activating glasses show equivalent rates of bone regeneration and fewer side effects with the bioactive glasses [25, 26]. For example, iliac crest autograft is currently the gold standard for spinal fusion. However, there are disadvantages of an autogenous graft including increased blood loss, increased operative time, second-site morbidity, and pain. A comparative study of bioactive glass (45S5 Bioglass) versus iliac crest autograft for spinal fusion in adolescent idiopathic scoliosis (AIS) has been reported for a group of 88 consecutive patients [26]. Forty patients received iliac crest autograft, and 48 received Bioglass synthetic bone graft (NovaBone particulate) with a minimum of 2-year follow-up. The results showed fewer infections (2 % vs. 5 %) and fewer mechanical failures (2 % vs. 7.5 %) in the Bioglass group. Loss of correction of the main thoracic curve was also less for the Bioglass group (11 % vs. 15.5 %). The conclusions for this retrospective study were: 1. Bioglass is as effective as iliac crest graft to achieve fusion and maintain correction in AIS.

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Table 3 Medical and dental products based upon 45S5 Bioglass Orthopedics Trauma Long bone fracture (acute and/or comminuted); alone and with internal fixation Femoral nonunion repair Tibial plateau fracture Arthroplasty Filler around implants (acetabular reconstruction) Impaction grafting General Filling of bone after cyst/tumor removal Spine fusion Interbody fusion (cervical, thoracolumbar, lumbar) Posterolateral fusion Adolescent idiopathic scoliosis Cranial-facial Cranioplasty Facial reconstruction General oral/dental defects Extraction sites Ridge augmentation Sinus elevation Cystectomies Osteotomies Periodontal repair Dental-maxillofacial-ENT Toothpaste and treatments for dentinal hypersensitivity Pulp capping Sinus obliteration Repair of orbital floor fracture Endosseous ridge maintenance implants Middle ear ossicular replacements (Douek MED)

2. Fewer complications were seen in the bioactive glass group of patients. 3. The morbidity of iliac crest harvesting can be avoided by use of bioactive glass in spinal fusion. These are important conclusions for the twenty-first-century challenge of affordable healthcare for the aged. Elimination of need for second-site (iliac crest) surgery in elderly patients that require spinal fusion or other reasons for a bone graft means less exposure to anesthesia and potential for infection. Use of the synthetic bone graft of bioactive glass also avoids pain and healing of the second site. An extensive range of medical and dental clinical applications of third-generation bioactive glasses, marketed as NovaBone and PerioGlas, has evolved over the last

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two decades, as summarized in Table 3. Clinical success is excellent for these applications, as discussed by Gaisser and Hench in Chapter 11 in Introduction to Bioceramics, 2nd edition [25].

Implications for the Future A genetic basis for development of a third generation of biomaterials provides the scientific foundation for molecular design of bioactive materials for in situ tissue regeneration and repair, preferably using minimally invasive surgery. There are significant economic and humanistic advantages to each of these new approaches that may aid in solving the problems of care for an aging population. It should be feasible to design a new generation of gene-activating biomaterials tailored for specific patients and disease states. New, predictive analytical methods are becoming available that can aid in developing such innovative approaches to affordable healthcare. These noninvasive analysis methods can be developed to make it possible for patients to be diagnosed and prescribed specific treatments based upon molecular biological data from their own cells rather than have to rely upon statistical trial and error prescriptions. Perhaps of even more importance is the possibility that bioactive stimuli can be used to activate genes in a preventative treatment to maintain the health of tissues as they age. Only a few years ago this concept would have seemed impossible. We need to remember that only a little more than 40 years ago the concept of a material that would not be rejected by living tissues also seemed impossible. This is now a clinical reality that has benefited tens of millions of people and should stimulate new concepts in the years ahead.

References 1. Hench LL (1991) Bioceramics: from concept to clinic. J Am Ceram Soc 74:1487–1510 2. Hench LL, Wilson J, Greenspan DC (2004) Bioglass: a short history and bibliography. J Aust Ceram Soc 40:1–42 3. Hench LL, Splinter RJ, Allen WC, Greenlee TK Jr (1971) Bonding mechanisms at the interface of ceramic prosthetic materials. J Biomed Mater Res 2(1):117–141 4. Hench LL (1998) Bioceramics. J Am Ceram Soc 81(7):1705–1728 5. Yamamuro T, Hench LL, Wilson J (eds) (1990) CRC handbook of bioactive ceramics, vol 2: Calcium phosphate and hydroxylapatite ceramics, CRC Press, Boca Raton, Florida 6. Klein CPAT, Wolke JGC, deGroot K (1993) Stability of Calcium Phosphate Ceramics and Plasma Sprayed Coating. In: Hench LL, Wilson J (eds) An introduction to bioceramics. World Scientific, London, p 199 7. Hench LL, Wilson J (eds) (1996) Clinical performance of skeletal prostheses. Chapman and Hall, London, pp 214–236 and 255–270 8. Yamamuro T (1996) A/W Glass-Ceramic: Clinical Applications. In: Hench LL, Wilson J (eds) An introduction to bioceramics. World Scientific, London, p 89–105 9. Wrobelewski BM, Fleming PA, Siney PD (1999) Charnley low-frictional torque arthroplasty of the hip. 20-to-30 year results. J Bone Joint Surg Br 81(3):427–430 10. Hench LL, Polak JM (2002) Third-generation biomedical materials. Science 295:1014–1017

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11. Jones JR (2012) Review of bioactive glass: from Hench to hybrids. Acta Biomater. doi:10.1016/ j.acbio.2012.08.023 12. Hench LL, Jones JR, Fenn MB (2012) New materials and technologies for healthcare. Imperial College Press, London 13. Wilson J, Low SB (1992) Bioactive ceramics for periodontal treatment: comparative studies. J Appl Biomater 3:123–169 14. Xynos ID, Hukkanen MVJ, Batten JJ, Buttery ID, Hench LL, Polak JM (2000) Bioglass 45S5 stimulates osteoblast turnover and enhances bone formation In vitro: implications and applications for bone tissue engineering. Calcif Tissue Int 67(4):321–329 15. Hench LL, Xynos ID, Buttery LD, Polak JM (2000) Bioactive materials to control cell cycle. J Mater Res Innov 3:313–323 16. Xynos ID, Edgar AJ, Buttery DKL, Hench LL, Polak JM (2000) Ionic products of bioactive glass dissolution increase proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression and protein synthesis. Biochem Biophys Res Commun 276:461–465 17. Xynos ID, Edgar AJ, Buttery DKL, Hench LL, Polak JM (2001) Gene-expression profiling of human osteoblasts following treatment with the ionic products of bioglass (R) 45S5 dissolution. J Biomed Mater Res 55:151–157 18. Bielby RC, Christodoulou IS, Pryce RS, Radford WJP, Hench LL, Polak JM (2004) Time- and concentration-dependent effects of dissolution products of 58S sol-gel bioactive glass on proliferation and differentiation of murine and human osteoblasts. Tissue Eng 10:1018–1026 19. Bielby RC, Pryce RS, Hench LL, Polak JM (2005) Enhanced derivation of osteogenic cells from murine embryonic stem cells after treatment with ionic dissolution products of 58s bioactive sol-gel glass. Tissue Eng 11:479–488 20. Christodoulou I, Buttery LDK, Saravanapavan P, Tai GP, Hench LL, Polak JM (2005) Doseand time-dependent effect of bioactive gel-glass ionic-dissolution products on human fetal osteoblast-specific gene expression. J Biomed Mater Res B Appl Biomater 74B:529–537 21. Christodoulou I, Buttery LDK, Saravanapavan P, Tai GP, Hench LL, Polak JM (2005) Characterization of human foetal osteoblasts by microarray analysis following stimulation with 58S bioactive gel-glass ionic dissolution products. J Biomed Mater Res B Appl Biomater 77B:431–446 22. Gough JE, Jones JR, Hench LL (2004) Nodule formation and mineralisation of human primary osteoblasts cultured on a porous bioactive glass scaffold. Biomaterials 25:2039–2046 23. Hench LL (2003) Glass and genes: the 2001 W. E. S. Turner memorial lecture. Glass Technol 44:1–10 24. Jones JR, Tsigkou O, Coates EE, Stevens MM, Polak JM, Hench LL (2007) Extracellular matrix formation and mineralization of on a phosphate-free porous bioactive glass scaffold using primary human osteoblast (HOB) cells. Biomaterials 28:1653–1663 25. Hench LL (ed) (2013) Introduction to bioceramics, 2nd edn. Imperial College Press, London 26. Ilharreborde B, Morel E, Fitoussi F, Presedo A, Souchet P, Pennecot G, Mazda K (2008) Bioactive glass as a bone substitute for spinal fusion in adolescent idiopathic scoliosis: a comparative study with iliac crest autograft. J Pediatr Orthop 28:347–351

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Fundamental Properties of Bioceramics and Biocomposites Maria Grazia Raucci, Daniela Giugliano, and Luigi Ambrosio

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Physicochemical Properties of Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Porosity Properties of Bioceramic Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biological Properties of Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Mechanical Properties of Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biocomposites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Sol–Gel Approach to Prepare Biocomposite Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Types of Reinforcements Used for Biocomposites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biological and Mechanical Properties of Biocomposites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Incorporation of Biomolecules . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Several varieties of ceramics, such as bioglass-type glasses, sintered hydroxyapatite, and glass-ceramic A–W, exhibit specific biological affinity, i.e., direct bonding to surrounding bone, when implanted in bony defects. These bonebonding ceramics are called bioactive ceramics and are utilized as important bone substitutes in the medical field. However, there is a limitation to their clinical applications because of their inappropriate mechanical properties. Natural M.G. Raucci (*) • D. Giugliano Institute of Polymers, Composites and Biomaterials (IPCB), National Research Council of Italy (CNR), Naples, Italy e-mail: [email protected]; [email protected] L. Ambrosio Department of Chemical Sciences and Materials Technology, National Research Council of Italy, Rome, Italy e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_3

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bone takes a kind of organic–inorganic composite, where apatite nanocrystals are precipitated on collagen fibers. Therefore, problems with the bioactive ceramics can be solved by material design based on the bioactive composites. In this chapter, an overview of fundamental properties of ceramics and biocomposite materials for biomedical application was reported. Keywords

Bone tissue • Bioceramics • Biocomposites • Biomimetic materials • Biocompatibility • Bioactivity • Osteoconductive • Hydroxyapatite • Porosity • Scaffolds • Mechanical properties

Introduction Native bone tissue possesses a nanocomposite structure, mainly composed of nonstoichiometric hydroxyapatite (HA; Ca10(PO4)6(OH)2) and collagen fiber matrix, that provides appropriate physical and biological properties, especially mechanical support and protection for the vertebrate skeleton. However, bone needs to be repaired or regenerated upon damage. The demand in the surgical market is highlighted by the fact that there are approximately four million operations involving bone grafting or bone substitutes performed around the world annually [1]. Currently, different types of bone grafts and bone graft substitutes, such as autografts, allografts, and alloplastic or synthetic bone grafts, are used for surgical treatments [2, 3]. Autografts compose approximately 58 % of the bone substitutes, whereas allografts constitute approximately 34 % [2]. As autografts have all the properties necessary for new bone growth, they are considered the gold standard for bone repair. However, the disadvantages of autografts include limited availability, donor site morbidity, and risk of disease transmission from donor to recipient [3]. Therefore, allografts are attractive alternatives to autografts; however, they are not usually osteoinductive or osteogenic, they have risks of an immunological reaction or disease transmission, and they have insufficient mechanical properties for loadbearing bone applications. Therefore, there is a great demand for synthetic grafts for fracture repair, and there is also scope to improve existing graft materials. Therefore, it gives rise to an abundant demand for safe and effective materials for use in tissue regeneration [4]. Biomaterials in the form of implants (sutures, bone plates, joint replacements, ligaments, vascular grafts, heart valves, intraocular lenses, dental implants, etc.) and medical devices (pacemakers, biosensors, artificial hearts, blood tubes, etc.) are widely used to replace and/or restore the function of traumatized or degenerated tissues or organs, to assist in healing, to improve function, to correct abnormalities, and thus to improve the quality of life of the patients. The world market for biomaterials is estimated to be around $12 billion per year, with an average global growth of between 7 % and 12 % per annum. Biomaterials are expected to perform in our body’s internal environment, which is very aggressive. For example, the pH of body fluids in various tissues varies in the range from 1 to 9. During daily activities, bones are subjected to a stress of approximately 4 MPa,

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whereas the tendons and ligaments experience peak stresses in the range 40–80 MPa. The mean load on a hip joint is up to three times the body weight (3000 N), and the peak load during jumping can be as high as ten times body weight. More importantly, these stresses are repetitive and fluctuating depending on the activities such as standing, sitting, jogging, stretching, and climbing [5]. In the early days, all kinds of natural materials such as wood, glue, and rubber and tissues from living forms and manufactured materials such as iron, gold, zinc, and glass were used as biomaterials based on trial and error. The host responses to these materials were extremely varied. Some materials were tolerated by the body, whereas others were not. Under certain conditions (characteristics of the host tissues and surgical procedure), some materials were tolerated by the body, whereas the same materials were rejected in another situation. Over the last 30 years, considerable progress has been made in understanding the interactions between the tissues and the materials. Researchers have coined the words “biomaterial” and “biocompatibility” [6] to indicate the biological performance of materials. Materials that are biocompatible are called biomaterials, and the biocompatibility is a descriptive term which indicates the ability of a material to perform with an appropriate host response, in a specific application [6]. In simple terms, it implies compatibility or harmony of the biomaterial with the living systems. Therefore, structural compatibility refers to the mechanical properties of the implant material, such as elastic modulus (or E, Young’s modulus) and strength, implant design (stiffness, which is a product of elastic modulus, E and second moment of area, I), and optimal load transmission (minimum interfacial strain mismatch) at the implant/tissue interface. Optimal interaction between biomaterial and host is reached when both the surface and structural compatibilities are met. Clinical experience clearly indicates that not all off-the-shelf materials (commonly used engineering materials) are suitable for biomedical applications. The various materials used in biomedical applications may be grouped into (a) metals, (b) ceramics, (c) polymers, and (d) composites made from various combinations of (a), (b), and (c). This chapter is intended to provide an overview of bioceramics and organic– inorganic composites obtained by sol–gel technology used in the biomedical field.

Bioceramics No material implanted in an organism is absolutely bioinert – each one stimulates a reaction in living tissue. The selection and improvement of biomaterials are a long process of trial and error which, according to contemporary research, is defined by four basic types of reaction between the material and the living organic tissue with which it is in contact [7]. There is not a sharp distinction between the various types of materials: they may be biotoxic, bioinert, bioactive, or bioresorbable to various degrees. Among inorganic substances, the only bioactive and bioresorbable materials are ceramics. A bioactive ceramics occupy an intermediate position between bioinert and bioresorbable [8]. A material can be designated as bioactive when it stimulates a specific biological reaction at the material–tissue interface, occurring with the formation of biochemical bonds between the living tissue and the material.

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The formation of such bonds with the surface of hydroxyapatite ceramics was first observed in the late 1960s [9]. The distinction between a bioactive and bioresorbable ceramics might be associated with only a structural factor: a nonporous hydroxyapatite ceramics behave as a bioinert material and is retained in an organism for at least 5–7 years without change, while a highly porous ceramics of the same composition can be resorbed in a time of the order of 1 year. Several varieties of ceramics have been found to exhibit bone-bonding performance. They are called bioactive ceramics, meaning that they can elicit biological activity. Bioglass-type glasses in the system Na2O–CaO–SiO2–P2O5 [10], sintered hydroxyapatite (Ca10(PO4)6(OH)2) [11], and glass-ceramic A–W are known to be bioactive ceramics. In the glass-ceramic A–W, oxyfluorapatite (Ca10(PO4)6(O,F2)) and β-wollastonite (CaO  SiO2) crystals are dispersed in a MgO–CaO–SiO2 glassy matrix. Most bioactive ceramics bond to bone through a low crystalline apatite layer formed on their surfaces in the body environment, created by a chemical reaction with body fluid. The apatite formation in vivo can be also observed in simulated body fluid (SBF; Na+ 142.0, K+ 5.0, Mg2+ 1.5, Ca2+ 2.5, Cl 147.8, HCO3 4.2, HPO42 1.0, and SO42 0.5 mol/m3) with a similar concentration to inorganic ions as human blood plasma [12]. In order to construct fundamental knowledge on novel bioactive materials design, apatite formation behavior on materials with different surface structures has been investigated in SBF by using metal oxide hydrogels and self-assembled monolayers (SAMs). Based on these results, several surface functional groups, such as Si–OH, Ti–OH, Zr–OH, Ta–OH, COOH, and SO3H [13], are found to be effective for triggering the heterogeneous nucleation of the apatite. In addition to surface functional groups, apatite formation is governed not only by the functional groups but also several other factors, such as ion release enhancing the apatite nucleation from the materials, and by a spatial gap constructed on the material surfaces [14].

Physicochemical Properties of Bioceramics Ceramics include a large class of nonmetallic materials whose chemical compositions, bond types, and properties vary over a very wide range. Therefore, it was logical to search among ceramics for bioactive and bioresorbable as well as bioinert materials. Bioceramics are more biocompatible with an organism than other implanted materials; have less effect on the immune system; have a broader range of biochemical, mechanical, and other properties; and can be adapted to a wider range of functional possibilities and lifetime requirements in the organism. A characteristic feature of the reaction of a material with a biosystem is the defining role of its structure. Therefore, ceramics (which can be distinguished as a class of materials in which structure is the basic physical parameter which determines service properties) by their nature are optimally suited for use in biosystems and additionally offer a wide range of possibilities to vary properties under such conditions. Thanks to these two factors, the use of bioceramics in surgery is substantially more effective than that of other implanted materials.

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Bioinert Ceramics. The term bioinert refers to any material that once placed in the human body has minimal interaction with its surrounding tissue; examples of these are stainless steel, titanium, alumina, partially stabilized zirconia, and ultrahigh molecular weight polyethylene. Generally, a fibrous capsule might form around bioinert implants; hence, its biofunctionality relies on tissue integration through the implant. Many of the more “traditional” ceramics have been used for bioceramic applications. Alumina and zirconia, for example, have been used as inert materials for a range of applications from the 1960s. Their high hardness, low friction coefficient, and excellent corrosion resistance offer a very low wear rate at the articulating surfaces in orthopedic applications. Microstructures are controlled to inhibit static fatigue and slow crack growth while the ceramic is under load. Alumina is currently used for orthopedic and dental implants. It has been utilized in wear bearing environments such as the total hip arthroplasties (THA) as the femoral head generating reductions in wear particles from ultrahigh molecular weight polyethylene (UHMWPE). Other applications for alumina encompass porous coatings for femoral stems, porous alumina spacers (specifically in revision surgery), and in the past as polycrystalline and single crystal forms in dental applications as tooth implants [15]. Bioactive and Resorbable Ceramics. A very typical case in surgical practice is the healing of bone defects which are formed in the operational removal of cysts, tumors, and genetic defects. For this purpose a bioactive ceramics able to reliably integrate with bone and maintain its strength over a long period of time is required, or else a resorbable which gradually disappears and is replaced with healthy bone. Porous granules and powders of bioactive and resorbable ceramics are very often used to fill bone defects. This is possible because of the bonding of dispersed ceramic material in the wound with blood fibrin and thrombin and after some time also with collagen fibers and elements of newly formed bone tissue, resulting in the creation of a so-called bone–ceramic complex with adequate strength. The formation of such a complex is observed also when bone defects are filled with porous granules of a bioinert ceramics. However, the use of resorbable ceramics is particularly effective, thanks to the eventually complete replacement of this complex with whole bone [16]. Bioactive and resorbable ceramics are used in all types of bone reconstruction, in particular for the fabrication of implants which densely fuse with bone (e.g., in skull restorations after operations or trauma), tooth-root implants, biological tooth fillings, cure of diseases of the periodontia (tissue around teeth), maxillofacial reconstruction, grafting and stabilizing skull bone, joint reconstruction, for the endoprosthesis of hearing aids, cosmetic eye prostheses, etc. Resorbable ceramics also aid in the restoration of tendons, ligaments, small blood vessels, and nerve fibers. Hydroxyapatite (HA) enters into the composition of bioactive and bioresorbable ceramics or substances close to it in composition which form HA crystals by the reaction with the organism at the implant–biomedium interface. Synthetic HA (Ca10(PO4)6(OH)2) is a complete chemical and crystallochemical analogue of bone mineral. Synthetic and natural HA differ only in structure. In bone, HA is present in the form of microcrystalline platelets with the approximate dimensions (1.5–3.5)  (5.0–10.0)  (40.0–50.0) nm. The strongest cortical bone is a

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composite material consisting of HA crystals (70 mass %) bonded with collagen fibers (30 mass %) [17]. Collagen is fibrous albumin composed of three peptide chains, each of which is twisted into a left-rotated spiral, intertwined to form a rightrotated spiral. For every turn of the triple spiral, there are ten turns of each individual chain. Various forms of synthetic HA (from highly dispersed powder to nonporous ceramics) do not exactly reproduce the structure of the natural HA crystals of bone but nevertheless are included in the metabolic processes of a living organism and reprocessed by the organism at a rate which depends on the structure, chemical composition, and specific surface area of the HA. The compound β-tricalcium phosphate (TCP) Ca3(PO4)2 is also a bioactive and resorbable ceramic material and, to a lesser degree, so are other calcium phosphates. All calcium phosphates convert to HA in the internal media of the organism. Certain calcium phosphate glass-crystalline materials and glasses which, like HA ceramics, are able to develop strong biochemical bonds with adjacent bone tissue, thanks to the formation of HA microcrystals on their surfaces upon reaction with the physiological fluid, are also considered as bioactive ceramics. Glass-crystalline materials may be considered as in essence ceramics with high concentrations of glass phase, while bioactive glasses are materials which contain small amounts of crystalline phase, or nuclei, formed by selective chemical solution and (or) annealing [18]. The basis for such a generalization is the property which all of these materials have in common – bioactivity. This signifies that in a biological medium, after a certain period of time, a biologically active layer consisting of microcrystals of hydroxyapatite plus a small amount of carbonate groups – hydroxycarbonate apatite (HCA) – forms on their surfaces. These crystals are structurally and chemically identical to the mineral component of bone and form strong chemical bonds with adjacent bone tissue. Most current bioceramic research utilizes this solution to measure the bioactivity of an artificial material by examining apatite-forming ability on its surface in simulated body fluids (SBF). This synthetic body fluid is highly supersaturated in calcium and phosphate in respect to apatite under even normal conditions. Therefore, if a material has a functional group effective for the apatite nucleation on its surface, it can form the apatite spontaneously [14]. It is widely accepted that the essential requirement for an artificial material to bond to living bone is the formation of bonelike apatite layer on its surface. Formation of the bonelike apatite layer on the bioactive materials can be produced in a SBF with ion concentrations almost equal to those of the human blood plasma. Osteoblasts have been shown to proliferate and differentiate on this apatite layer (Fig. 1). The mechanism of formation of a direct chemical bond between dense HA ceramics and bone can be described as follows: a cellular bone matrix composed of differentiated osteoblasts appears on the surface of the ceramics, which forms a 3–5 mm thick amorphous zone of increased electron density with a higher concentration of phosphate ions and calcium. Following this collagen bundles appear which connect the amorphous zone and bone cells. Next, crystals of bone mineral form in the amorphous zone. As this zone matures, the bonding region shrinks to 0.05–0.2 mm. As a result, the living bone is

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Fig. 1 Biomineralized scaffold of polycaprolactone/hydroxyapatite (PCL/HA) after treatment in simulated body fluid (7 days) at different magnification

joined to the implant by a thin epitaxially bonded layer [19]. Transmission electron microscopic analysis of the crystal structure of the bone–HA interface indicates that there is practically complete epitaxial reproduction by the growing crystals of bone–HA of the orientation of apatite crystals on the implant surface. The mechanism of formation of HCA crystals on the surface of bioceramics with a glass phase is similar and also includes several steps [20]: formation of an amorphous silica gel, partially dissolved by the organism via exchange of alkaline cations of the glass with protons of the biomedia; condensation and repolymerization of the silicate layer surface impoverished in alkaline and alkaline-earth cations; migration of Ca2+ and PO43 ions to the surface through the silicate layer; adsorption of calcium and phosphate ions from the biomedium with formation, on the silicate layer surface, of an amorphous film enriched in calcium phosphate; and crystallization of the calcium phosphate film with the participation of OH and CO32 ions to form HCA microcrystals. The parameters of this process (bioactivity characteristics) are determined basically by the chemical, not the phase, composition of the material (ratio of glass and crystalline phases), although the phase and structural states strongly affect the strength and other properties. The strength properties of bioactive ceramics are substantially lower than those of bioinert. Highest strength is found in nonporous HA and in the glass-ceramic Cerabone A–W, but the bend strength and crack resistance of these materials are five to seven times lower than that of high-strength ZrO2 ceramics. By synthesizing bioactive ceramic–biopolymer composites, it is possible to somewhat increase tensile strength and impact toughness and at the same time decrease Young’s modulus, bringing the elasticity of the biomaterial close to that of bone.

Porosity Properties of Bioceramic Scaffolds In spite of the serious mechanical limitations, bioceramics of calcium orthophosphates are available in various physical forms: powders, particles, granules

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Fig. 2 (a) Scaffold material based on PCL/HA colonized by human mesenchymal stem cells after 14 days cell culture; (b) hMSC as bridge between wall pores

(or granulates [21]), dense blocks, porous scaffolds, injectable formulations, selfsetting cements, implant coatings, and composite component of different origin (natural, biological, or synthetic). Furthermore, custom-designed shapes like wedges for tibial opening osteotomy, cones for spine and knee, and inserts for vertebral cage fusion are also available [22]. Surface area of porous bodies and scaffolds is much higher, which guarantees good mechanical fixation in addition to providing sites on the surface that allow chemical bonding between the bioceramics and bones [19]. Furthermore, pore sizes are directly related to bone formation, since they provide surface and space for cell adhesion and bone ingrowth. On the other hand, pore interconnection provides the way for cell distribution and migration, as well as it allows an efficient in vivo blood vessel formation suitable for sustaining bone tissue neoformation and possibly remodeling. Namely, porous HA bioceramics can be colonized by bone tissues [23]. Therefore, interconnecting macroporosity (pore size >100 μm) [24], which is defined by its capacity to be colonized by cells, is intentionally introduced in solid bioceramics (Fig. 2). Macroporosity is usually formed due to release of various volatile materials, and, for that reason, incorporation of pore creating additives (porogens) is the most popular technique to create macroporosity. The porogens are crystals or particles of either volatile (they evolve gases at elevated temperatures) or soluble substances, such as paraffin, NaCl, NaHCO3, gelatin, poly(methyl methacrylate), or even hydrogen peroxide [25]. Obviously, the ideal porogen should be nontoxic and be removed at ambient temperature, thereby allowing the ceramic/porogen mixture to be injected directly into a defect site and allowing the scaffold to fit the defect [24]. Sintering particles, preferably spheres of equal size, is a similar way to generate porous threedimensional (3D) bioceramics of calcium orthophosphates. However, the pores resulting from this method are often irregular in size and shape and not fully interconnected with one another. A wetting solution, such as polyvinyl alcohol, is usually used to aid compaction, which is achieved by pressing the particles into cylinders at approximately 200 MPa [26].

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Several other techniques, such as replication of polymer foams by impregnation, dual-phase mixing technique, freeze casting, stereolithography, and foaming of gel-casting suspensions, have been applied to fabricate porous calcium orthophosphate bioceramics [27]. In vivo response of calcium orthophosphate bioceramics of different porosity was investigated, and hardly any effect of macropore dimensions (150, 260, 510, and 1220 μm) was observed [28]. In another study, a greater differentiation of mesenchymal stem cells was observed when cultured on 200 μm pore size HA scaffolds when compared to those on 500 μm pore size HA [29]. The latter finding was attributed to the fact that the higher pore volume in 500 μm macropore scaffolds might contribute to the lack of cell confluency leading to the cells proliferating before beginning differentiation. Besides, the authors hypothesized that bioceramics having a less than optimal pore dimensions induced quiescence in differentiated osteoblasts due to reduced cell confluency. Already in 1979, Holmes suggested that the optimal pore range was 200–400 μm with the average human osteon size of ~223 μm [30], while in 1997 Tsurga and coworkers suggested that the optimal pore size of bioceramics that supported ectopic bone formation was 300–400 μm [31]. Therefore, there is no need to create calcium orthophosphate bioceramics with very big pores; however, the pores must be interconnected [24]. Interconnectivity governs a depth of cells or tissue penetration into the porous bioceramics, as well as it allows development of blood vessels required for new bone nourishing and wastes removal [32]. Meanwhile, the presence of microporosity provides both a greater surface area for protein adsorption and increased ionic solubility. Studies showed that increasing of both the specific surface area and pore volume of bioceramics might greatly accelerate the process of biological apatite deposition and, therefore, enhance the bone-forming bioactivity. More importantly, the precise control over the porosity, pore sizes, and internal pore architecture of bioceramics on different length scales is essential for understanding of the structure–bioactivity relationship and the rational design of better bone-forming biomaterials.

Biological Properties of Bioceramics The most important differences between bioactive bioceramics and all other implanted materials are inclusion in the metabolic processes of the organism, adaptation of either surface or the entire material to the biomedium, integration of a bioactive implant with bone tissues at the molecular level, or complete replacement of resorbable material by healthy bone tissues. All of the enumerated processes are related to the effect of an organism on the implant. Nevertheless, another aspect of implantation is also important – the effect of the implant on the organism. For example, using of bone implants from corpses or animals, even after they have been treated in various ways, provokes a substantially negative immune reactions in the organism, which substantially limits the use of such implants. In this connection, it is useful to dwell on the biological properties of bioceramic implants, particularly

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those of calcium orthophosphates, which in the course of time may be resorbed completely [33]. It has been accepted that no foreign material placed within a living body is completely compatible. The only substances that conform completely are those manufactured by the body itself (autogenous) and any other substance that is recognized as foreign initiates some types of reactions (host–tissue response). The reactions occurring at the biomaterial/tissue interfaces lead to time-dependent changes in the surface characteristics of the implanted biomaterials and the tissues at the interface [34]. In order to develop new products, it is desirable to understand the in vivo host responses. Like any other species, biomaterials and bioceramics react chemically with their environment, and, ideally, they should not induce any change or provoke undesired reaction in the neighboring or distant tissues. Generally, both bioactivity and bioresorbability phenomena are fine examples of chemical reactivity and calcium orthophosphates (both non-substituted and ion-substituted ones) fall into these two categories of bioceramics [35]. A bioactive material will dissolve slightly but promote formation of a surface layer of biological apatite before interfacing directly with the tissue at the atomic level which results in the formation of a direct chemical bond with bone. Such an implant will provide a good stabilization for materials that are subject to mechanical loading. A bioresorbable material will dissolve and allow a newly formed tissue to grow into any surface irregularities but may not necessarily interface directly with the material. Consequently, the functions of bioresorbable materials are to participate in the dynamic processes of formation and reabsorption that take place in bone tissues; so bioresorbable materials are used as scaffolds or filling spaces allowing tissue infiltration and substitution. A distinction between the bioactive and bioresorbable bioceramics might be associated with a structural factor only. For example, bioceramics made from nonporous, dense, and highly crystalline HA behave as a bioinert (but a bioactive) material and are retained in an organism for at least 5–7 years without changes, while a highly porous bioceramics of the same composition can be resorbed approximately within a year. Before recently, it was generally considered that, alone, any type of synthetic bioceramics possessed neither osteogenic nor osteoinductive properties and demonstrated minimal immediate structural support. When attached to the healthy bones, osteoid is produced directly onto the surfaces of bioceramics in the absence of a soft tissue interface. Consequently, the osteoid mineralizes, and the resulting new bone undergoes remodeling. However, several reports have already shown some osteoinductive properties of certain types of calcium orthophosphate bioceramics [36]. Although in certain in vivo experiments an inflammatory reaction was observed after implantation of calcium orthophosphate bioceramics [37], the general conclusion on using calcium orthophosphates with Ca/P ionic ratio within 1.0–1.7 is that all types of implants (bioceramics of various porosities and structures, powders, or granules) are not only nontoxic but also induce neither inflammatory nor foreignbody reactions. An intermediate layer of fibrous tissue between the implants and bones has never been detected. Furthermore, calcium orthophosphate implants display the ability to directly bond to bones [33].

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When a bioceramic implant is fixed in the human body, a space filled with biofluids exists next to the implant surface. With time, proteins will be adsorbed at the bioceramic surface that will give rise to osteoinduction by proliferation of cells and their differentiation toward bone cells, revascularization, and eventual gap closing. Ideally, a strong bond will be formed between the implant and surrounding tissues. Osteoblasts cultured on HA bioceramics are generally reported to be completely flattened, and its cytoplasmic edge is difficult to distinguish from the HA surfaces after 2 h incubation [38]. Osteoblasts cultured on porous HA bioceramics appeared to exhibit a higher adhesion, an enhanced differentiation, and suppressed proliferation rates when compared to nonporous controls. Furthermore, formation of distinct resorption pits on HA [39] and β-TCP [40] surfaces in the presence of osteoclasts was observed. A surface roughness of calcium orthophosphate bioceramics was reported to strongly influence the activation of mononuclear precursors to mature osteoclasts [39]. Mesenchymal stem cells are one of the most attractive cell lines for application as bone grafts. Early investigations by Okumura et al. indicated an adhesion, proliferation, and differentiation, which ultimately became new bone and integrated with porous HA bioceramics [41]. Recently, Unger et al. showed a sustained coculture of endothelial cells and osteoblasts on HA scaffolds for up to 6 weeks [42]. Furthermore, a release of factors by endothelial and osteoblast cells in coculture supported proliferation and differentiation was suggested to ultimately result in microcapillary-like vessel formation and supported a neo-tissue growth within the scaffold [42].

Mechanical Properties of Bioceramics In the body, the mechanical properties of natural bone change with their biological location because the crystallinity, porosity, and composition of bone adjust to the biological and biomechanical environment. The properties of synthetic calcium phosphates vary significantly with their crystallinity, grain size, porosity, and composition (e.g., calcium deficiency) as well. In general, the mechanical properties of synthetic calcium phosphates decrease significantly with increasing amorphous phase, microporosity, and grain size. High crystallinity, low porosity, and small grain size tend to give higher stiffness, compressive and tensile strengths, and greater fracture toughness. It has been reported that the flexural strength and fracture toughness of dense HA are much lower in a dry condition than in a wet condition [43]. Moreover, by comparing the properties of HA and related calcium phosphates with those of cortical bone, it was found that bone has a reasonably good compressive strength though it is lower than that of HA. But bone has a significantly higher fracture toughness than HA. The mechanical properties are even lower for porous HA structures [43]. The high tensile strength and fracture toughness of bone are attributed to the tough and flexible collagen fibers reinforced by HA crystals. Hence, calcium phosphates alone cannot be used for load-bearing scaffolds despite their good biocompatibility and osteoconductivity.

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Biocomposites Even bioactive ceramics cannot substitute in load-bearing portions of bone, because their fracture toughness is lower and their Young’s modulus is higher than those of human cortical bone. In addition, these ceramics are difficult to form into the desired shapes during implantation. Novel bioactive bone substitutes with high flexibility and high machinability are desired in the medical field. Natural bone takes a kind of organic–inorganic composite, where apatite nanocrystals are precipitated on collagen fibers. Therefore the organic–inorganic composites inspired by the bone structure are expected to solve these problems. In fact, the development of composite scaffold materials is attractive as advantageous properties of two or more types of materials can be combined to suit better the mechanical and physiological demands of the host tissue. For instance, natural and synthetic polymers such as collagen, polylactic acid (PLA), polyglycolic acid (PGA), copolymers from the grafting of PLA and PGA (PLGA), or polycaprolactone (PCL) showed suitable properties for the application in tissue engineering [44]. By taking advantage of the formability of polymers and including controlled volume fractions of a bioactive ceramic phase, mechanical reinforcement of the fabricated scaffold can be achieved [45]. At the same time, the poor bioactivity of most polymers can be counteracted. Probably the most important driving force behind the development of polymer/bioceramic composite scaffolds for bone tissue engineering is the need for conferring bioactive behavior to the polymer matrix, which is achieved by the bioactive inclusions or coatings. The degree of bioactivity is adjustable by the volume fraction, size, shape, and arrangement of fillers [47]. It has been shown that increased volume fraction and higher surface-area-to-volume ratio of fillers favor higher bioactivity; hence, in some applications the incorporation of fibers instead of particles is favored [46]. Addition of bioactive phases to bioresorbable polymers can also alter the polymer degradation behavior, by allowing rapid exchange of protons in water for alkali in the glass or ceramic. This mechanism is suggested to provide a pH-buffering effect at the polymer surface, modifying the acidic polymer degradation. In related research, it has been reported that polymer composites filled with HA particles hydrolyzed homogeneously due to water penetrating the interfacial regions [47]. Ideally, the degradation and resorption kinetics of composite scaffolds are designed to allow cells to proliferate and secrete their own extracellular matrix, while the scaffolds gradually vanish, leaving space for new cell and tissue growth. The physical support provided by the 3D scaffold should be maintained until the engineered tissue has sufficient mechanical integrity to support itself. Recently, composite materials comprising a bioactive phase within a biodegradable polymer matrix have been developed. In particular, the challenging idea to design tissue-inspired composite materials moves toward the synthesis of ceramic–polymer composites which show advantages over either pure ceramic or polymer [48, 49], resulting in a superior material for the specific application. Traditionally, calcium phosphate-based ceramics have proved to be attractive materials for biological applications. Among these bioceramics, particular attention has been given to hydroxyapatite Ca10(PO4)6(OH)2 whose atomic calcium to phosphorus (Ca/P) ratio is 1.67.

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In recent years, design studies of scaffold composite materials were performed to successfully reproduce the microenvironment required to support and nurture the molecular interactions which occur within tissues, between cells, and within the mineralized extracellular matrix (ECM). It is well known that scaffold architecture plays a crucial role in initial cell attachment and subsequent migration into and through the matrix and in mass transfer of nutrients and metabolites, providing sufficient space for development and later remodeling of the organized tissue [50]. An adequate definition of the morphological features (i.e., pore size) is strongly required to assure cell adhesion, molecular transport, vascularization, and osteogenesis. For instance, small pores (with diameters of few microns) favor hypoxic conditions and induce osteochondral formation before osteogenesis occurs [51]. In contrast, scaffold architectures with larger pores (several 100 μ in size) rapidly induce a well-vascularized network and lead to direct osteogenesis [51]. There are numerous ways to synthesize HA. Widely used processes include aqueous colloidal precipitation, sol–gel, solid-state, and mechanochemical methods. These may be synthesized at room temperature and provide the ability to control directly the particle and grain sizes [48]. In comparison with traditional strategies involving physical mixing of HA, they assure a more controlled and fine distribution of crystallites of compounds into polymer matrices. This provides an improved mechanical response in terms of strength, stiffness, toughness, and fatigue resistance to reach the complete mechanical compatibility [46].

Sol–Gel Approach to Prepare Biocomposite Materials A sol–gel method enables the powder less processing of glasses, ceramics, and thin films or fibers directly from solution. Precursors are mixed at the molecular level, and variously shaped materials may be formed at much lower temperatures than it is possible by traditional methods of preparation. One of the major advantages of sol–gel processing is the possibility to synthesize hybrid organic–inorganic materials. Combination of inorganic and organic networks facilitates the design of new engineering materials with exciting properties for a wide range of applications. The organic–inorganic hybrid materials may be prepared in various ways. The simplest one relies on dissolution of organic molecules in a liquid sol–gel [52]. The other way uses the impregnation of a porous gel in the organic solution. In the third type, the inorganic precursor either already has an organic group or reactions occur in a liquid solution to form chemical bonds in the hybrid gel. The sol–gel process itself leads to formation of gels from mixtures of liquid reagents (sols) at room temperatures. It involves several steps: the evolution of inorganic networks, formation of colloidal suspension (sol), and gelation of the sol to form a network in a continuous liquid phase (gel). During the “aging” step (after gelation and before drying), the sol–gel-derived material expulses the liquid phase (solvent which can be water or alcohol) in the process called syneresis. Drying of the obtained gels, even at room temperature, produces glass-like materials called xerogels (i.e., xeros). The process generates a porous material, where the pore size depends on such factors as

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Fig. 3 Hydroxyapatite nanocrystals (HA) synthesized by sol–gel technology

time and temperature of the hydrolysis and the kind of catalyst used. The diameter of the pores is directly related to the shrinkage of the “wet” gels. During the drying process, the gel volume decreases even several times (which is the main reason of cracking). In particular, the sol–gel process for preparing HA usually can produce fine-grained microstructure containing a mixture of nano-to-submicron crystals (Fig. 3). These crystals can be better accepted by the host tissue. The sol–gel product is characterized by nanosize dimension of the primary particles. This small domain is a very important parameter for improvement of the contact reaction and the stability at the artificial/ natural bone interface. Moreover, the high reactivity of the sol–gel powder allows a reduction of processing temperature and any degradation phenomena occurring during sintering [53]. Moreover, the low temperature of process allows to introduce bioactive molecules (i.e., growth factors, peptides, dendrimer, antibiotics) sensible to high temperature [54]. The major limitation of the sol–gel technique application is linked to the possible hydrolysis of phosphates and the high cost of the raw materials. On the other hand, most of the sol–gel processes require a strict pH control, vigorous agitation, and a long time for hydrolysis. These problems were solved by using a non-alkoxidebased sol–gel approach where the calcium and phosphate precursors are calcium nitrate tetrahydrate and phosphorous pentoxide, respectively [53]. More importantly, gel formation is achieved without the need for any refluxing steps. In this case, the P2O5 reacts with alcohol to form P(O)(OR)3 oxyalkoxide with the liberation of water, which in turn partially hydrolyzes the oxyalkoxide precursors. The presence of phosphorus hydroxyl alkoxide is not sufficient in itself to form a gel, indicating the important role of Ca(NO3)2  4H2O. It may be speculated that Ca(NO3)2  4H2O probably results in the

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generation of alkoxy-nitrate salts which participate in a polymerization reaction with the partially hydrolyzed phosphate precursors, the polymerization reaction thereby resulting in the gel. For the synthesis of the organic–inorganic composite material PCL/HA, the polymer may be added during the production of inorganic phase in order to allow the chemical interactions between the components [55, 56]. Molecular-level mixing of calcium and phosphorous precursors with the polymer chains derived from the sol–gel process resulted in composites having enhanced dispersion and exhibiting good interaction between the inorganic phase and the polymer matrix. Several studies have demonstrated by FTIR and AFM analyses the presence of hydrogen bond in the composite materials synthesized by sol–gel method [56]. Moreover, the presence of HA particles in the composite material beneficially offsets the acidic release from the polymer through the alkaline calcium phosphate and mitigates erosion problems associated with the release of acidic degradation products. In vivo and in vitro measurements of pH in bone chambers have shown that the pH drop is 0.2 units near the eroding polyesters [57]. Furthermore, it is possible to evaluate a homogeneous distribution of n-HA crystals with 10–30 nm as diameter and 40–50 nm as length obtained by sol-gel process. A homogeneous distribution of nanoscale hydroxyapatite particles in the polymeric matrix allows an increase of bioactive potential of materials. Many investigations of nanophase materials to date have illustrated their potential for bone repair. For example, increased osteoblast adhesion on nano-grained materials in comparison to conventional (micron grained) materials has been reported [58]. Osteoblast proliferation in vitro and long-term functions were also enhanced on ceramics with grain or fiber sizes 70  C) or extremely aggressive chemical conditions during processing, being a challenge to the scaffold fabrication process. To achieve this aim, a sol–gel processing might be a strategy to incorporate biomolecules during scaffold fabrication [54]. To the authors’ knowledge, however, sol–gel-derived bioactive organic–inorganic hybrids have not yet been formed into highly interconnected porous structures, which is essential for application of these composites as scaffolds.

Conclusions Bone tissue engineering has emerged as a new area of regenerative medicine, and biomaterials have an essential function concerning cell adhesion, spreading, proliferation, differentiation, and tissue formation in three dimensions. Material design based on organic–inorganic composites not only improves weak points in ceramic biomaterials but also provides various biological functions such as drug delivery and tissue regeneration. From the materials science perspective, the present challenge in tissue engineering is to design and fabricate reproducible bioactive and bioresorbable 3D scaffolds of tailored porosity and pore structure, which are able to maintain their structure and integrity for predictable times, even under load-bearing conditions. To improve the osteoconductive and osteoinductive material properties, the incorporation of biomolecules such as growth factors with the aim to accelerate local bone healing is promising and currently under extensive research. Incorporating biomolecules during scaffold processing however is not simple as biomolecules are sensitive to elevated temperatures and extreme chemical conditions. This aim is achieved by using a sol–gel route that also allows a good dispersion of ceramic particles in biocomposites allowing to enhance the fundamental properties of materials.

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Bioinert Ceramics: Zirconia and Alumina Corrado Piconi and Alessandro Alan Porporati

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina: Physical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina: Mechanical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina Stability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia Physical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia Mechanical Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia Stability . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia Radioactivity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina–Zirconia Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconia-Toughened Alumina . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina-Toughened Zirconia . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biocompatibility of Bioinert Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . In vitro Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . In vivo Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Carcinogeneticity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Reaction to Ceramic Wear Debris . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Processing of Bioinert Ceramics for Joint Replacements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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C. Piconi (*) Medicine and Surgery Department, Clinical Orthopedics and Traumatology Institute, Catholic University, Rome, Italy e-mail: [email protected] A.A. Porporati Medical Products Division, CeramTec GmbH, Plochingen, Germany e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_4

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Abstract

Alumina and zirconia are used as biomaterials since long. The use of alumina, especially, as a dental implant and porous bone substitute was reported in the first half of the 1960s. Both these ceramics exhibit a high chemical inertness, which is the reason for their high biological safety even at the higher specific surfaces. Due to their hardness being higher than the other metal alloys, alumina and zirconia found their main application as biomaterials in the articular surfaces of joint replacements. Today, the use of zirconia ball heads has practically ceased in hip arthroplasty. Zirconia ceramics are used mostly in dentistry, while in orthopedics are still in use in some niche products. In association with alumina, zirconia is used in the ceramic composites that are presently the reference bioinert ceramic for clinical applications. This chapter is an overview of the development of bioinert ceramics, as well as of their physical and mechanical properties. The behavior of the present composite bioinert ceramics is also described in detail, and a review is given of the studies carried out to assess the biological safety of alumina and zirconia. Finally, the flow sheet of a possible production process for the manufacture of bioinert ceramic components is outlined.

Introduction At the end of the 1960s, arthroplasty was becoming a well-accepted procedure in orthopedics, thanks to the introduction devices making use of a metal-on-metal (MoM) and metal-on-polyethylene (MoP) bearings. The medium term follow-ups nevertheless were showing the relevance of the reactions to wear debris in the cascade of events leading to osteolysis and to asepting loosening of implants. There was the need of low-wear bearings, and Dr. Pierre Boutin, a surgeon with practice in Pau, a town in southern France, started to investigate the feasibility of an all-ceramic bearing making use of high purity alumina. It is likely that Boutin focused his attention on alumina on the basis of the good clinical results obtained by the alumina dental implants developed by Sandhaus in Switzerland and by Driskell in the USA in the same period [1]. Alumina was by that time the most advanced ceramic material, thanks to the availability of high alumina powder Degussit Al23 in the market that allowed to overcome many limitations of the former ceramic precursors. While it is likely that these pioneers selected alumina for its inertness in the biological environment, this concept has been abandoned since a long while because it is known that whatever material once implanted elicits a host reaction. Nevertheless, the concept of “bioinert” ceramics is still used to distinguish alumina and zirconia from the so-called bioactive or bioreactive materials, and in this way it is also used in the title of this chapter. Zirconia was introduced as a biomaterial during the second half of the 1980s [2]. The mechanical properties of zirconia allowed overcoming some limitation in design of alumina ceramic devices. Moreover, the better mechanical properties of zirconia led to consider the ceramic components made out this material (essentially

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ball heads for Total Hip Replacements) more forgiving about mechanical stresses than the corresponding alumina ones. Unfortunately, the Yttria-stabilized Tetragonal Zirconia Polycrystals (Y-TZP) – i.e., the zirconia ceramic largely used as a biomaterial – is a metastable material whose processing has several criticalities. This led to a number of failures and to the practical abandonment of zirconia in orthopedics around the year 2000, and to a review of the requirements specified in international standards for load bearing ceramics for clinical applications. While the use of zirconia in orthopedics was stopped, there was a growing interest in dentistry where presently this material is used for the production of dental implants, as well for CAD-CAM production of bridges, crowns, and in structures of dentures [3]. So far, alumina and zirconia are used in a synergistic way in alumina-zirconia composites, which are the standard load-bearing bioceramics in hip arthroplasty. “Bioinert” ceramic composites are showing mechanical and wear behavior that allowed designing of a number of innovative medical devices during the last years and other ones to come in the future (see Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacement,” this Handbook).

Alumina Alumina (Aluminum oxide – Al2O3) is one of the biomaterials with longer clinical use. Alumina has about 45 years of clinical record in orthopedics where it is still in use as “pure” alumina either as a component of high-performances ceramic composites. The material characteristics improved through the years thanks to the improvements introduced in raw materials (ceramic precursors) and in ceramic processing. Alumina is so widely accepted in orthopedics that even in scientific journals the use of the term “ceramic” without any other specification is indicating alumina, a wording that seems to ignore that a number of other ceramics are today in clinical use. However, in this chapter, we will follow the current use, then “ceramic” will mean “alumina”, if not differently specified. The German Patent by M. Rock, issued in 1933, contains the first mention of alumina as a material for the construction of “artificial spare parts . . . for humans and animal bodies” [4]. It was uneventful, and the real use of alumina in medical device took place during the 1960s in dental implants thanks to Sandhaus and Driskell [1]. In the same years, Sir John Charnley demonstrated the advantages of “low friction arthroplasty” of the hip over former devices. It was in this atmosphere that Dr. Pierre Boutin with the help of one of his patients, who was a top manager of a CGE France factory that was nearby the Boutin’s practice in Pau, developed the first alumina-on-alumina bearing in total hip arthroplasty. In this device, both the stem and the cup were cemented into the host bone, while the ball head was fixed on a cylindrical trunnion at the extremity of the stem by epoxy glue, a solution that was source of a number of early failures in the first series and was soon abandoned. Besides Boutin, several other researchers in Germany, Japan, and in the USA were investigating the biomedical applications of alumina during the same years. In Germany, Langer in Keramed, Heimke in Friedrichsfeld, Dörre in Feldm€ule, and

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Maier in Rosenthal were working in a wide research program funded by the Ministry of Research to develop alumina clinical implants. In Japan, Oonishi and Kawahara developed alumina devices in close cooperation with Kyocera engineers. Especially relevant are the development of alumina knee replacements (see Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacement”) and the use of alumina single crystal hip replacement stems implanted after resection of bone tumors. In the USA, Smith, Hulbert, Driskell, and Ducheyne were investigating the potential of alumina in replacing segmental bone defects and in dental implants [5]. The German research program was aimed to develop hip replacements designed to achieve cementless fixation in the host bone, i.e., by a screw-like profile of the cup. The clinical use of these early large and heavy monolithic devices allowed surgeons and engineers to assess the clinical advantages of alumina-on-alumina bearings in terms of wear reduction and absence of particle-induced osteolysis. Poor outcomes were due to the component design or to failures in the cementation. Nevertheless, in younger patients with good bone quality – in which good primary stability in the cortical bone was achieved – several positive clinical results had been reported [6, 7].

Alumina: Physical Properties Alumina (Aluminum oxide, Al2O3) applications as a biomaterial are based on the microstructural properties of this oxide that may occur in many metastable phases, which irreversibly transform into alpha-alumina if heated above 1200  C. Alphaalumina is a close packed hexagonal arrangement of oxygen ions that constitute a very high thermodynamically stable phase. Alpha-alumina (also known as corundum, or emery if containing impurities), that is, the material selected for biomedical application, is well known in nature as being the matrix of several gemstones, e.g., ruby red in color for the presence of chromium, or sapphire, blue in color due to both titanium and iron. The alumina molecule contains strong ionic and covalent chemical bonds between Al3+ and O2 ions that are the origin of the high melting point, hardness, and resistance to the attack of strong inorganic acids of this material. In other words, alpha-alumina is aluminum metal in its highest oxidative state with high chemical and physical stability: namely, it also has a marked resistance to the attack of strong inorganic acids, like e.g., orthophosphoric or hydrofluoric acid. The low electric and thermal conductivity and the high melting point are also due to the energy of the ionic-covalent bonds in the solid. In the lattice of alpha-alumina, each aluminum cation Al3+ is surrounded by oxygen anions O2 forming two regular triangles on both sides, twisted by 180  and lying on parallel planes. The surface layer of O2 anions allow the chemisorption on the surface of OH+ groups, then the bonding of water molecules or proteins. In other words, the surface has high wettability, e.g., higher of several metallic alloys.

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Fig. 1 Microstructure of the monolithic alumina BIOLOX ®forte (Courtesy CeramTec GmbH, Plochingen, Germany)

Surface wettability along with high hardness (2000–2200 GPa) makes alphaalumina the ideal material in industrial wear applications (e.g., ruby ball bearings) as well as for the components of the bearings for arthroprostheses joints, whose surface, thanks to the high hardness of alpha-alumina, is extremely resistant to scratching. On the other hand, the high hardness that characterizes this material results in complex and costly machining (Fig. 1). The extremely smooth surface finish achievable by diamond polishing (typical values: Ra = 0.02 μm, Rt = 0.5 μm), the hardness, and the wettability of alphaalumina are the reasons of the low wear of the joint replacement bearings making use of alumina components. A further physical property to be controlled to achieve the best clinical outcomes in alumina bioceramics is the residual porosity of the products. The porosity fraction influences the mechanical properties as explained in the following but especially the open porosity fraction has to be controlled. Namely, liquids and gases can penetrate the ceramic by open porosity, and then it is mandatory to avoid open porosity in alumina devices to be used in the human body.

Alumina: Mechanical Properties Alumina, as most ceramics, has moderate tensile and bending resistance and a brittle fracture behavior. Unlike in metals, there is no yielding at the tip of cracks in alumina to dissipate stress. The unavoidable presence of surface defects, notches, and internal flaws increases locally the stress concentration. Especially tensile stresses stimulate the growth of the size of defects, until they become real cracks, and the ceramic fractures. The brittle fracture behavior of alumina (its low toughness) is the main limit in designing alumina components. In addition, because the population of defects in ceramics components has a large scatter in size and in location (see Table 1), the relationships between stress distribution, failure probability, and strength in ceramics need to be discussed on a statistical basis [8].

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Table 1 Physical properties of alpha-alumina Property Crystallography Lattice parameter a Lattice parameter c Melting point Density (theoretical) Thermal expansion coeff. Thermal conductivity (298 K, >99.9%TD) Specific heath (298 K)

Units – nm nm K g/cm3 – W/m K J/Kg K

Value Hexagonal 0,476 1,299 2310 3,986 6,5 30 8–10 104

Table 2 Defects in ceramics and their origin Defect type Delamination

Defect size 1–100 μm

Organic inclusion Inorganic inclusion Surface defects after green machining Porosity due to powder aggregation Residual porosity Large grains

0.1–5 μm 0.01–1 μm 0.01–0.05 μm

Surface defects after hard machining Inhomogeneous grain size

0.03–1 μm 0.001–1 μm Depending on microstructure 0.001–0.5 μm 0.0005–1 μm

Origin Powder consolidation. Sintering cycle. Uncontrolled shrinkage Powder (batch) contamination Powder (batch) contamination Binder selection, batch preparation Binder selection, batch preparation Powder consolidation Powder consolidation, incomplete sintering Powder selection. Batch preparation. Sintering cycle Grinding tools selection. Machining parameters Powder selection. Batch preparation. Sintering cycle

Alumina is an ionic-covalent solid that does not yield under load as metal and alloys do. The strong chemical bonds in alumina are the roots of several of its behaviors, e.g., the low electric and thermal conductivity, the high melting point that makes it practically impossible to shape alumina by casting, and high hardness that characterizes this material which makes its machining complex and costly. Ceramics used in joint replacements are polycrystals obtained by the consolidation of powders. Their structure contains a number of “defects” (See Table 2) that may develop their size under stress without any apparent strain at macroscopic level until they reach a critical size, giving origin to the real crack eventually that leads to the catastrophic fracture of the piece. An evident difference between ceramic and metals is that ceramics do not show plastic deformation or yielding before fracture. In other words, ceramic under stress accumulate elastic energy with negligible strain. The energy stored in the material will be

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released suddenly would its level overcome a given threshold, originating the fracture of the ceramic part. The defects – especially the ones that are long and thin in shape – are “sensitive” to tensile stresses, which stimulate their growth (the defective surfaces are “thrown apart”) while compressive loads result safer because they tend to “close” the defect. This behavior can be controlled by the selection of a proper high-purity raw material, by the reduction of the grain size (meaning smaller grain boundaries) and an increase in density (meaning less porosity, less flaws) either by the introduction of toughening mechanism (e.g., by the transformation of selected phase in the material). Nevertheless, although the tensile (bending) strength of several ceramics could exceed the one of many metallic alloys, these materials show fracture energies lower than the metal ones, a fact that has to be carefully taken into account in the design of ceramic components. One can assume that a ceramic piece will break as a chain: a chain fails if a single link fails. In the same way, ceramic fails if a single “defect” grows under stress until failure. It is noted that the distribution of defects in the structure of a ceramic is not casual, but depends on the overall production process. The design of ceramics must then be made by a probabilistic approach, taking into account the probability that above a given stress level, a single defect can develop its size and grow leading the piece to failure. This is made by applying the theory developed by W. Weibull [9]. Briefly, he developed a mathematical model which is today incorporated in the ISO standard 20501 [10] that allows to determine the probability that a given number of ceramic parts obtained by a given production process can survive until a given stress threshold. This is obtained by the mathematical processing of a set of fracture stress measured on a series of samples identical in shape and volume. The results are the Weibull characteristic strength, σ 0, and the Weibull modulus, m. They are representing respectively the stress threshold for the survival of 37 % of the samples and the variability of the fracture stress among the samples. The higher the m, the lower the variability in strength and the reliability of the production process of the material. As a biomaterial, alumina ceramic has undergone deep improvements in mechanical properties during 40 years of clinical use as described in Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacement” of this Handbook. The improvements in the selection of raw materials used as precursors and in sintering process resulted in marked reduction in grain size and in the increase in density, close to the theoretical one [11]. The microstructural improvements gave remarkable increase in bending strength (from less than 400 MPa to more than 630 MPa in pure alumina components – see Table 3). However, only the introduction in clinics of the alumina composite allowed overcoming the limits of alumina in terms of mechanical properties. The toughness and bending strength of the most used alumina composite BIOLOX ®delta (CeramTec GmbH, Plochingen, Germany) are more than twice the former market leader BIOLOX ®forte (CeramTec GmbH, Plochingen, Germany), which is a monolithic alumina.

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Table 3 Mechanical properties of medical grade alumina and alumina composites

Property Al2O3 content ZrO2 Content Density Average grain size Al2O3 Young modulus Hardness Flexural strength Fracture toughness

Units Vol %

Alumina (1970s) 99.1–99.6

BIOLOX ® (since 1974) 99.7

BIOLOX ® forte (since 1995) >99.8

BIOLOX ® delta 80

Ceramys ® ATZ 20

Vol %

–

–

–

17

80

g/cm3 μm

3.90–3.95 4.5

3,95 4

3.97 1.75

4,37 0.56

5.51 0.4

GPa

380

410

407

358

–

GPa MPa

17.7 >300

20 (HV1) 400

20 (HV1) 631

19 (HV1) 1384

15 (HV20) 1039

MPa m½

3.5

4

4.5

6.5

7.4

Alumina Stability The long-term stability of alumina in the body environment is one of the cornerstones of its selection as a biomaterial. Strength degradation may occur in alumina due to permeation of liquids through open porosity, which in this way can interact with the microstructural defects (flaws) that are intrinsic to ceramics microstructure, as discussed above. High density and zero open porosity are characteristics relevant for the reliability of alumina components to be used in joint replacements. The defects in the microstructure of the material have a size (subcritical size) that does not influence the mechanical properties. Nevertheless, depending on several variables (e.g., chemical species concentration, environment, temperature, and stress level) they can grow at a very low speed reaching a critical size, leading the piece to failure. This behavior is known as Subcritical Crack Growth (SCG). Especially in the late 1970s, SCG was a relevant problem in some alumina devices used in joint replacements, because of the impurity of alkali, calcium and silicon oxides contained in precursors. Strength degradation can take place for calcium segregation at grain boundaries, as well as for CaO interactions with the aqueous environment. Silica limits densification and promotes grain growth during sintering, decreasing the material strength. Alkalis (e.g., NaO) may segregate at grain boundaries and their dissolution in physiologic liquids may lead to accelerated fatigue fracture [5]. Although the incidence of this problem was not the same for all the materials manufactured at that time, the control of the strength degradation due to chemical impurities in alumina was achieved only after the release of the standard ISO 6474 in 1980 that specified the minimum purity requirements of alumina for clinical implants.

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Zirconia The metal zirconium is known from very ancient times: the name is a derivation of the Arabic Zargon, (golden in color), which derives in turn from the Persian Zar (Gold) and Gun (Color). Zirconium gemstones are cited in the St. John Apocalypse among the ones forming the walls of the Heavenly Jerusalem. Zirconia, the metal dioxide (ZrO2), was identified in 1789 by German chemist Martin Heinrich Klaproth, and was used for a long time as pigment for ceramics. The development of zirconia (zirconium dioxide: ZrO2) ceramics as a biomaterial started in the mid-1980s, to manufacture ball heads for total hip replacements able to overcome the mechanical limits of the alumina ceramics produced on that time. The early developments were oriented toward Magnesia-Partially Stabilized Zirconia (MgPSZ), in which tetragonal phase is present as small acicular precipitates within large cubic grains (Ø 40  50 μm) forming the matrix. As this feature may negatively influence the wear properties of the polyethylene used in the acetabular components that are coupled to zirconia ball heads in total hip replacements, most of the developments were focused on Yttria-stabilized Tetragonal Zirconia Polycrystal (Y-TZP), a ceramic formed almost completely by submicron-sized grains, which become a standard material for clinical applications (Fig. 2) [2]. In addition, it is noted that the tetragonal acicular phase is nucleated within cubic zirconia grain during aging steps on cooling, maintaining precisely controlled time/temperature conditions. This implies a tough process control that may not be compatible with mass production requirements.

Fig. 2 Microstructure of Y-TZP (Cortesy Dr. L.Pilloni, ENEA, Italy)

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The use of zirconia in hip joint arthroplasty is now almost abandoned, since the worldwide recall that affected the product that was the leader of the market in 2000. So far, Y-ZTP ceramics are used especially in dentistry as blanks for CAD-CAM machining of crowns, bridging either structures for full partial dentures, as well as dental implants and abutments for titanium fixtures [1]. In orthopedics, Y-TZP is used only in the femoral components of two total knee replacements models. Readers are addressed to Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacement,” this Handbook, for further details.

Zirconia Physical Properties Zirconia (Zirconium dioxide, ZrO2) is an allotropic oxide, i.e., with the same chemical formula it may occur in three crystal structures – monoclinic, tetragonal, and cubic – depending on temperature. Its crystal lattice is monoclinic up to 1170  C when it shifts to tetragonal, a crystal shape which is maintained up to 2370  C. At this temperature the tetragonal lattice shifts into cubic, and it maintains this form up to melting (2680  C). The changes in crystal structure of zirconia are reversible and are associated to changes in volume and shape of the crystal cells. On cooling, the change in volume of unconstrained crystals in the cubic-tetragonal transition is about 2.5 %, while it is about 4 % in the tetragonal-monoclinic phase transition. Especially the t ! m transformation that takes place at a temperature in the range 1000–1200  C on cooling is of technological interest (Table 4). In a polycrystal the strain generated by the lattice expansion and by the shear strain due to the change in shape of the tetragonal cell implies an increase of stress on the neighboring, untransformed grains. At a macroscopic level, this can shatter sintered products of pure zirconium dioxide. For these reasons, pure unstabilized Zirconium dioxide was used mainly either as a refractory or as a pigment until 1929, when the stabilization of cubic phase to room temperature was demonstrated by

Table 4 Physical properties of zirconia. Data on alumina reported for comparison Material Crystallography Lattice parameter a Lattice parameter b Lattice parameter c Density (theoretical) Melting point Thermal expansion coeff. Thermal conductivity (298 K) (>99.9 %TD)

– nm nm nm g/cm3 K – W/m K

Zirconia Monoclinic 0.5156 0.5191 0.5304 5.83 2790 5–10 2.5

Hexagonal 0.5094 0.5177 – 6.1

Cubic 0.5124 – – 6.06

Alumina Hexagonal 0.476 – 1.299 3,98 2310 6.5 30

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Ruff et al. [12]. Cubic zirconia is still used as an abrasive in jewelry to replace diamonds thanks to its lower cost and high refractive index. Engineering applications of zirconia ceramics were made possible by the discovery of the stabilization of the tetragonal phase to room temperature. This ceramic that was composed of zirconia in cubic and tetragonal phase was named Partially Stabilized Zirconia (PSZ) by its discoverers [13] because the low concentration of the stabilizing oxide does not allow the full stabilization of the cubic phase. PSZ was obtained using calcium oxide as stabilizer, but successively either magnesium oxide (magnesia, MgO) or yttrium oxide (yttria, Y2O3) was used for this role. The real breakthrough in zirconia came in 1975 with the publication in Nature of the paper “Ceramic Steel?” by Garvie et al. [14]. They reported an observation of increased toughness due to the transformation of the tetragonal phase into monoclinic phase in MgO-stabilized PSZ (Mg-PSZ). This transformation performs as an effective dissipative mechanism for fracture energy. It takes place in a “martensitic” way as in some steels, hence the title of the paper. The phase transition (i) is not associated with transport of matter (it is diffusionless); (ii) is athermal, because it takes place at a temperature range and not at a given temperature; and (iii) involves a change in shape of the crystal lattice. Shortly after Rieth et al. [15] and Gupta et al. [16] showed successfully that a zirconia ceramic would be almost completely formed by tetragonal grains at room temperature, using Yttrium Oxide (Yttria – Y2O3) as stabilizer of the tetragonal phase was made feasible. These ceramic materials, that were called Yttria-Tetragonal Zirconia Polycrystal (Y-TZP), contained 2–3 mol% yttria (Y2O3). Y-TZP is almost completely constituted by tetragonal grains some hundreds of nanometers in size (0.3–0.5 μm). Cubic zirconia as well as monoclinic grains can be present as minor phases. It is noted that a number of oxides have been studied as stabilizers for zirconia (CaO, MgO, Y2O3, CeO2, Er2O3, Eu2O3, Gd2O3, Sc2O3, La2O3, and Yb2O3). In zirconia applied in biomedical applications, studies were focused on materials stabilized by CaO, MgO, and CeO but the ceramic that was developed industrially for the production of a medical devices was the one stabilized by Y2O3 [17]. The tetragonal grains have the ability to transform into monoclinic; the transformation implies some 4 % volume expansion in free grains. This is the origin of the toughness of the material, e.g., of its ability to dissipate fracture energy. When the constraint exerted on the grains is relieved, i.e., by a crack advancing in the material, the grains next to the crack tip shift into monoclinic phase. The neighboring grains then experience some strain to accommodate both the change in volume of the transformed grain and the shear strain due to the change in shape of the crystal cells. These microstructural changes take place at the expense of the elastic stresses that lead to the advancement of the crack. At a macroscopic level, this results in the increase of toughness of the material. It is noted that, besides the energy dissipated in the phase transformation as described, the crack has to also overcome the compressive stress field due to the volume expansion of the grains to advance further. There are several very accurate reviews of the models proposed to explain the mechanisms of transformation toughening in zirconia ceramics, [18–20] which the reader can refer for details.

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Briefly, the model proposed by Lange [18] explains the transformation in terms of change in total free energy (ΔGt-m). This model describes the change in free energy as the sum of three quantities: (i) change in chemical free energy (ΔGc), (ii) strain energy, due to the expansion of the transformed particles (ΔGSE), and (iii) change in surface energy depending on the shape change of the particle and on microcracking, e.g.,: ΔGtm ¼ ΔGc þ ΔGSE þ ΔGs While the change in chemical free energy (ΔGc) depends on temperature and chemical composition, the strain energy change ΔGSE (>0) depends on stiffness of the matrix and on the residual stress due to processing that is unavoidable in a sintered polycrystal due to the mismatch of thermal expansion coefficients of the two phases. The third term (ΔGs) represents the change in free energy due to the change in the surface of grains during phase transition, as well as the formation of new surface due to microcracking. As remarked by Lughi and Sergo [17], the contribution to the free energy change of this term becomes relevant only when the size of the grains is in the order of magnitude of some tens of nanometers or so, because it is a function of the square of the crystal size, while the other two terms scale with its cube. In a summary, the energy dissipated in the transformation and then the contribution of the phase transformation to mechanical properties depends on three main features of the material: – The concentration of the stabilizing oxide (Yttria) – The constraint exerted by the matrix onto grains – The size and shape of grains and their homogeneous distribution The equilibrium among these characteristics should be carefully tuned, to maintain the tetragonal phase in Y-TZP in a metastable state, optimizing in this way the contribution of the transformation to the mechanical properties of the material. The grains size should be fine and evenly distributed to avoid the increase of energy required for the transformation, and to decrease the size of grain boundaries. Yttria concentration too should be tuned, to maximize the fraction of tetragonal phase retained at room temperature and on its “transformability.” The residual stresses in the matrix generated on cooling by the differences in shrinkage in the crystal lattice and the stiffness of the material are also giving a marked contribution to the energy threshold occurring to the tetragonal grains to transform into monoclinic. Care is to be given to maintain the structure of Y-TZP in a metastable state: it has been demonstrated experimentally that there is a decrease in bending strength when the material is overstabilized.

Zirconia Mechanical Properties The transformation toughening mechanism described in the former part of this chapter results in the remarkable mechanical behavior of zirconia (Table 5). Zirconia

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Table 5 Mechanical properties of zirconia. Data on alumina behavior reported for comparison Property Al2O3 content Density Average grain size Flexural strength Young modulus Fracture toughness Hardness

Units Vol % g/cm3 μm MPa GPa MPa m½ GPa

Y-TZP – 6 0,3–0,5 900–1200 210 7–9 12,5

Mg-PSZ – 5,5 40–70 400–650 210 8–11 12,5

Alumina (1970s) 99.1–99.6 3.90–3.95 4.5 >300 380 3,5 18

HIP Alumina >99.8 3,97 1,75 650 400 4,5 20–21

is an ionic ceramic polycrystal, then there is not the plastic strain before fracture characteristic of metals and of metallic alloys that is the dissipative mechanism for fracture energy in these materials. Although zirconia has a high fracture strength, toughness is lower than in metals, although twice the one of high-density alpha-alumina (9–12 MPa m½ Vs. 4–6 MPa m½). Y-TZP bending strength is above 900 MPa (4-point bending), than in the order of the one of several alloys, while the elastic modulus of zirconia – about 200 GPa, similar to the modulus of Titanium alloys – is a feature that makes easy to design metal to ceramic connections because this minimizes microstrains under load. Besides bending strength and toughness, the hardness and thermal conductivity of zirconia are the behavior of zirconia especially relevant for its use in structural application in dentistry. Zirconia has a very low thermal conductivity (e.g., about one-tenth of alumina) and this has to be taken into account during hard machining of zirconia components by grinding because the thermal spikes that may arise would trigger the progressive degradation (LTD) of the mechanical properties of the material. This can happen also in total hip replacements bearing surfaces in case of starved lubrication. This behavior, along with the hardness of zirconia, which is lower than that of alumina (12–13 GPa Vs. 20–21 GPa) and sensitive to scratching, has discouraged the use of zirconia in ceramic-on-ceramic THR bearings. As a conclusive remark on zirconia, it is worthwhile to underline that the behavior of this material strongly depends on the selection of the ceramic precursors and on their tailoring for the consolidation process during the batch preparation step, on the machining and surface finish of the componenets, all characteristics peculiar to each ceramic manufacturer. Overall, end users must keep in mind that zirconia is a metastable material whose behavior depends not only on the processing conditions but also on the environmental conditions that the components will experience during their lifetime.

Zirconia Stability It is an established behavior that the metastable tetragonal phase in Y-TZP ceramics can spontaneously transform into monoclinic. The discovery of this behavior – termed “aging”, hydrothermal transformation, or Low Temperature Degradation (LTD) – is due to Kobayashi et al. [21] who observed the degradation in the

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mechanical properties with the progress of transformation. In the following of this chapter we use the three terms indifferently. LTD is characterized by several well-assessed features: (i) the transformation starts at the surface of the material and progresses into it; (ii) the monoclinic transition of the tetragonal grains in the surface results in a surface uplift; (iii) the stresses lead to grain pull-out and surface cracking; (iv) the transformation penetrates into the material, enhancing the extent of the process; (v) as the cracks extension progresses, one can observe the decrease of the material density and the reduction in strength and toughness of the ceramic. There is experimental evidence that LTD can take place also at room temperature in a humid environment. However, it is enhanced at temperatures above 100  C under saturated steam. Due to its mechanical relevance, aging in zirconia ceramics has been investigated since the first attempts to use this ceramic as a biomaterial [2]. Nevertheless, none of the models developed so far give a comprehensive representation of this behavior [17]. The models more debated were proposed by Sato et al. [22], Yoshimura et al. [23], and more recently by Chevalier et al. [24]. Sato’s and Yoshimura’s models are based on the local reaction interaction of preferential sites in the material surface with water molecules. The first model [22] is based on the reaction between water and Zr-O-Zr bridges at the grain boundaries, resulting in the formation of Zr-OH bonds. Following this model, this is leading to the change in the conditions for the stabilization of the tetragonal phase at room temperature, triggering the transformation into monoclinic and the cascade of events of LTD. Also, the Yoshimura model [23] takes into account the formation of Zr-OH bonds, which give origin to stress fields in the crystal lattice. Once higher than a given threshold, the transformation starts leading to material degradation with its progress. The model proposed by Chevalier [24] is based on the filling by O2 ions of the oxygen vacancies due to the presence of the trivalent Yttrium in the ZrO2 lattice. LTD is especially evident on Y-TZP, while other zirconia (Mg-PSZ, Ce-TZP) are not affected by this problem. In contrast, the mechanical behavior of both these zirconia are largely surpassed by the Y-TZP ones, a fact that makes the latter ceramic the one preferred for clinical use, formerly in orthopedics and now in dentistry. The rate of strength degradation can be controlled acting on several parameters: high density, small and uniform grain size, creation of a spatial gradient of yttria concentration and introduction of grains within the matrix of Alumina that can delay the onset of the transformation. All the above parameters are controlled by the manufacturing process and by the physicochemical behavior of the precursors selected for the production of the ceramic. Then LTD is peculiar to each Y-TZP material and of each manufacturing process.

Zirconia Radioactivity The issue of radioactivity of zirconia used in medical devices was raised during the second half of the 1980s after the sudden increase in the level of attention of

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the general public to the presence of radionuclides in the environment. This happened after two severe accidents in nuclear power plants (Three Miles Island, Chernobyl) that occurred at that time. In fact, zirconia ores are associated to 238U, 232Th, and to their decay products in concentrations depending on the source. The presence of radionuclides in low-purity, low-cost zirconia, e.g., sands used in smelting works or in refractories – is monitored by the public health authorities since long. The production of the chemical used in the preparation of the high- purity ceramic precursors of zirconia enacts the effective separation of radioactive contaminants from the powders. However, the use of zirconia in medical devices imply to characterize this feature, and a significant number of studies was published during the first half of 1990s during the setup of the standard ISO 13356 concerning zirconia for clinical implants [2]. It has been demonstrated that separation technologies used in the purification of nuclear wastes [25] allows to obtain chemicals suitable for the production of high-purity zirconia precursors with specific gamma activity lower than the one of human body, i.e., less than 50 Bq/kg [1]. It is noted that some bioceramics used in dentistry since a long while have specific activity far higher than the one of zirconia, e.g., feldspathic porcelains due to the natural abundance of 40 K in Potassium.

Alumina–Zirconia Composites The addition of alumina to zirconia, forming an alumina–zirconia composite (Al2O3–ZrO2) lowers significantly the zirconia transformation kinetics. For this reason, the two phase alumina–zirconia material has been marketed for orthopedic applications after zirconia material was abandoned. The explanation of the retardation of phase transformation kinetics afforded by only minor alumina concentration when zirconia is the major phase is not satisfying yet. Such ceramic composite materials are commonly defined as Alumina- Toughened Zirconia (ATZ) ceramics. Instead, alumina containing small amount of zirconia, commonly referred as zirconia-toughened alumina (ZTA) or alternatively named alumina matrix composite (AMC), is also exceptionally resistant to hydrothermal aging. Moreover, because of topological reasons, the slower aging kinetics is perhaps clearer in these materials, where alumina is the major phase. In fact, its presence in the zirconia phase, finely distributed into the material and consequently the zirconia grains, isolated by the stiff zirconia matrix, the percolative pathway risk is reduced and consequently degradation cannot continue deep into the material. As a consequence, the ZTA doesn’t increase in roughness as the monoclinic zirconia increases after hydrothermal treatment. Furthermore, the transformation is hindered by stiff predominant alumina matrix. However, besides the hydrothermal aging problems, the two design approaches put in evidence the basic material strategy adopted by the material scientists: the high-fracture toughness may be the leading feature in an ATZ, whereas hardness may be the main feature in a ZTA. The ZTA material had achieved in the last decade the

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predominance, because of highly suitable properties for couple bearing application in THA, but perhaps, also because of the fear of mostly zirconia-based materials. In following, these two types of material design approaches will be discussed.

Zirconia-Toughened Alumina The basic concept of a Zirconia-Toughened Alumina (ZTA) material is to keep or improve the alumina properties as hardness, stiffness, and thermal conductivity, which made successful for more than 40 years its use as implantable ceramic in THA and at the same time, by using zirconia as reinforcing element, to substantially increase the material fracture toughness and strength by exploiting its tetragonal to monoclinic phase transformation as described in section “Zirconia Physical Properties” of this chapter for monolithic zirconia components. The enhancement of mechanical properties is aimed to increase the reliability of alumina THA components but also to allow a higher degree of freedom in the design of THA components. The ability of zirconia phase transformation (t-ZrO2 ! m-ZrO2) in ZTA ceramics is an indispensable prerequisite for their excellent mechanical properties and the degree of stabilization of the zirconia tetragonal phase at body temperature is essential for the desired toughening mechanism. Y2O3 is the most widely used tetragonal zirconia chemical stabilizer, whose microstructure and grain size contribute to tetragonal zirconia phase stabilization. However, stabilization must be achieved such that no material degradation will occur in body environment, i.e., in aqueous liquid (synovia), which is known to potentially trigger phase transformation at the surface of ceramic components as elucidated in section “▶ Zirconia Stability”.

Structure of Zirconia-Toughened Alumina The Zirconia-Toughened Alumina is basically a 2-phase alumina matrix ceramic composite consisting of fine homogeneously dispersed zirconia grains. The zirconia volume content needs to be precisely balanced to obtain the optimum resistance against crack extension while assuring an optimal chemical stability. In this way, it may be possible to increase fracture toughness and strength of the material to the levels equal to and in some cases above those seen in monolithic zirconia. The most successful representative of this demanding type of material in arthoplasty, is the ZTA marketed under the brand name of BIOLOX ®delta (ISO 6474–2) manufactured by CeramTec Medical Products Division in Plochingen, Germany. BIOLOX ®delta THA components are clinically used for more than 10 years and have been implanted in over 3.5 million cases. This material is today considered the ceramic golden standard in THA. BIOLOX ®delta consists of finely grained (with an average Al2O3 diameter of about 0.5–0.6 μm) high-purity alumina matrix, which is very similar to the wellknown monolithic alumina material BIOLOX ®forte (ISO 6474). Common to other composite materials, its basic physical properties such as stiffness, hardness, thermal conductivity, etc. are mainly predetermined from the dominating phase. Fracture toughness and strength are increased with the addition of reinforcing elements to the

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Fig. 3 SEM image of the microstructure of the ZTA composite ceramic BIOLOX ®delta (Courtesy CeramTec GmbH, Plochingen, Germany). The gray grains represent the alumina matrix, the white grains the zirconia phase and the elongated grains consist of reinforcing strontium hexaaluminate (SHA) platelets

microstructure. The best compromise between the mechanical properties and the chemical stability is achieved for this material by incorporating about 17 Vol.% of zirconia in the alumina matrix. Such zirconia content would be slightly higher of percolation pathway threshold limit of 16 Vol.% [24]; nevertheless, Pecharromán et al. [25] have analyzed the threshold limit for hydrothermal stability of zirconia in alumina matrix composites by steam sterilization and revealed that a relevant aging level for such material is given at 18–22 % volume content of zirconia. As further reinforcing element, strontia (SrO) is added to the composition; this additive reacts with alumina during the sintering stage to grow in situ elongated strontium hexaaluminate (SHA) crystals with a magnetopumbite structure. These SHA platelet-shaped grains are homogenously dispersed in the ceramic composite microstructure and provide an additional barrier to crack propagation. Indeed, their task is to deflect any subcritical cracks created during the lifetime of the ceramic. The maximum platelets length is of about 5 μm with an aspect ratio of 5–10. Figure 3 shows the microstructure of BIOLOX ®delta, where the gray grains represent the alumina matrix and the white grains the zirconia phase. In addition, the elongated grains consist of reinforcing SHA platelets. These two well-known reinforcing effects in material science, transformation toughening and crack deflection, give BIOLOX ® delta a unique strength and toughness, so far unattained by any other ceramic material extensively used in a structural application in the human body. Additionally, stabilizing elements are also added to the reinforcing elements of the material. A minor amount of chromia (Cr2O3), which is completely soluble in the alumina matrix, donates to the material its pink color; and yttria (Y2O3), which is soluble in the finely distributed zirconia phase, supports the stabilization of the zirconia tetragonal phase. Its concentration has been optimized in order to obtain the maximum resistance to crack propagation, while keeping an optimal chemical stability. However, as noted in section “▶ Zirconia”, the best performance is achieved by properly triggering the tetragonal zirconia transformation by means of [24]:

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Fig. 4 Reinforcing mechanism in BIOLOX ®delta at crack initiation and propagation. Yellow particles represent tetragonal zirconia. Color change to red indicates monoclinic phase transformation. Arrows show the region of compressive stresses due to phase transformation (Courtesy CeramTec GmbH, Plochingen, Germany)

– – – –

The size, shape, and distribution of zirconia particles The addition of yttria The amount of stiff surrounding alumina matrix The internal and external stresses

It is a matter of fact that the above concurring factors make the design of a ZTA composite particularly difficult; higher tetragonal zirconia stabilization as a consequence, for example, of a too high yttria concentration would lead to the suppression of zirconia phase transformation annihilating almost all the improvements of ZTA. On the other hand, poor zirconia stabilization caused, for example, by zirconia uncontrolled grain growth due to a noncorrect sintering would destabilize the tetragonal phase enhancing the low-temperature degradation susceptibility. Nevertheless, in the later case, the mechanical properties might be outstanding, but the material could be unreliable leading to catastrophic consequences due to LTD affecting the product that was leader of the Y-TZP for THA market in the year 2000. Figure 4 represents a realistic part of the microstructure. The gray particles refer to the alumina matrix and yellow to the tetragonal zirconia. The zirconia t ! m transformation is indicated by the change to red color. In the case of severe overloading, crack initiation and crack extension will occur. High tensile stresses in the vicinity of the crack tip will trigger the phase transformation of the zirconia particles. The accompanied volume expansion leads to the formation of compressive stresses which will stop the crack extension.

Mechanical Properties of Zirconia-Toughened Alumina The necessary requirements of a ZTA for a use in THA are expressed by the standard for composite materials ISO 6474–2. BIOLOX ®delta as representative of material category fulfill the ISO requirements as shown in Table 6 (see also Table 3). The excellent properties of the material BIOLOX ®delta support advantageous properties of the final product, namely, ceramic hard-hard bearings for hip

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Table 6 Comparison of regular BIOLOX ®delta mechanical properties with the ISO 6474–2 requirements Property Al2O3 content ZrO2 content Density Average grain size Al2O3 Average grain size ZrO2 Flexural strength Young modulus Fracture toughness Hardness (HV1)

Units Vol % Vol % g/cm3 μm μm MPa GPa MPa m½ GPa

ISO 6474-2 – – – 1.5 0.6 1000 320 4 16

BIOLOX ®delta 81.6 17 4,37 0.56 0.3 1384 358 6,5 19

Table 7 Comparison of strength and burst load of BIOLOX ®forte and BIOLOX ®delta Parameter Strength Burst load Burst load

Test/design 4-point bending 28–12/14 L 36–12/14 M

Unit MPa kN kN

BIOLOX ® forte 620 54 110

BIOLOX ®delta 1400 85 131

Ratio delta/forte 2.3 1.6 1.2

arthroplasty. Indeed, in general, the performance of any hip replacement system depends on the intrinsic material properties, the design and manufacturing quality of the whole device, the external load and the particular environment, and finally the quality of mounting and installation. The use of high-performance materials inevitably promotes the performance of a hip replacement system – however, the latter factors may be even more decisive for the success of a system, and for this reason, they were also taken into account in the ISO 6474–2. These complex correlations must be necessarily evaluated by design analysis, modeling, simulations, risk analysis, and many other tools. In order to eliminate any influences of design features most of the material testing is performed using 4-point bending bars. The burst load of the components is significantly increased as shown in Table 7 [11]. All burst loads are far above the required value of 46 kN and the maximum in vivo load at worst-case conditions of approximately 10 kN. Kuntz et al. [11] obtained the data of the burst tests given in Table 7 from ball heads with identical geometry, titanium alloy test taper, and the same test setup. In this way, they showed that the advantage of BIOLOX®delta ball heads in the burst load came only from the higher strength of the material in comparison to the monolithic alumina BIOLOX ®forte. However, although the strength of BIOLOX ®delta resulted to be more than twice the strength of BIOLOX ®forte, the ratio of the burst strength values was lower. The authors justified the result with the ductile deformation of the titanium alloy taper during the burst test which steadily increases the contact area of the conical bore of the ceramic ball head and the metal taper. However, generally speaking, the burst load of identical ball heads is always higher when a high-strength material is used.

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Fig. 5 Typical scratches on BIOLOX®forte and BIOLOX ®delta performed according to the standard EN 1071–3:2005. The zirconia phase transformation in BIOLOX ®delta enhances the ceramic scratch resistance avoiding the extensive grain pull-out shown by the monolithic alumina

It is also of particular interest to see that the burst strength resulted to be higher for the larger ball head size. As a consequence, a clear benefit of high-performance material is also the possibility to apply a challenging design, such as components with lower wall thickness, which would be otherwise unaffordable. We need to remember that a high-performance material always increases the safety margin of the component. As result, according to the Australian Orthopaedic Association National Joint Replacement Register (AOA 2014), the fracture rate of ceramic components, with the adoption of the alumina composites, decreased drastically, making the breakage an extremely rare occurrence (i.e., 0.17/10’000 procedures for ZTA, where alumina shows 6.48 fractures for 10’000 procedures). The wear behavior of the ZTA material is also interesting; Clarke et al. [27] showed that there is no statistical difference between ZTA and alumina BIOLOX ®forte wear couples, but under off-normal conditions as in the microseparation tests, the ceramic composite material BIOLOX ®delta wear couple showed significantly lower wear rate. Such behavior is justified by the phase transformation toughening acting in the ZTA material as clarified by Piconi et al. [28] with the scratch resistance measurements performed according to standard EN 1071–3:2005. Figure 5 shows a scratch comparison between BIOLOX ®forte and BIOLOX ®delta.

Alumina-Toughened Zirconia Begand et al. [29] presented during 2004 a new alumina-toughened material (ATZ) for THA. A commercial material was developed and manufactured by Mathys Orthopädie GmbH (Mörsdorf, Germany) and was introduced in the international market in 2010. Similarly to the ZTA solution presented by CeramTec GmbH a couple of years earlier, Mathys tried with its ATZ ceramic composite – marketed with the brand name Ceramys® – to combine the advantages of the two monolithic materials, Al2O3 and ZrO2 with the aim to avoid the disadvantages. The material and the components’ properties seem to be remarkable and in some cases better than the ZTA alternative, as specified by Mathys. Nevertheless, the ATZ material has up to date not achieved the broad use of its competitor; in about 4 years Mathys declared 30,000 components sold worldwide in the market. The reason might be researched first of all in the fear of zirconia, that still exist in the THA field, and on the other

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hand Mathys being a prosthetic company with whole THA implants in its portfolio, making difficult to sell its ceramic bearing solution to other prosthetic companies. Moreover, the price pressure in THA field might also be a possible explanation, with alumina being a much cheaper raw material than zirconia.

Structure of Alumina-Toughened Zirconia The ATZ Ceramys ® consists of 80 w.% of zirconia and 20 w.% of alumina. More in detail, Schneider et al. [30] describe the ATZ Ceramys ® as formed by 61 % tetragonal zirconia, 17 % cubic zirconia, approx. 1 % monoclinic zirconia, and alpha-alumina as the balance. The tetragonal zirconia phase is stabilized with 3 mol.% of yttria as for the standard monolithic zirconia (3Y-TZP). The alumina grains are finely dispersed in the zirconia matrix and the average grain size diameter approach is 0.4 μm, both for ZrO2 and Al2O3. Indeed, on the paper, ATZ may appear of relatively simple design, nevertheless as pointed out previously, the properties of these special materials depend significantly onthe production steps and conditions which are different for all single manufacturers. Mechanical Properties of Alumina-Toughened Zirconia Ceramys ® properties are summarized in Table 8. According to Oberbach et al. [31] the biaxial flexural strength (acc. to norm ISO 6474) surpassed the 1200 MPa with a Weibull modulus of 18. The burst strength performed on hip joint heads with a diameter of 28 mm with 12/14 L and 12/14 M necks and titanium alloy tapers achieved values [32] comparable to that showed by ZTA material in Table 7. The pin-on-disk wear test with ATZ/ATZ couplings using serum as fluid test medium showed to be comparable to Al2O3/Al2O3 couplings (i.e., 0.152 and 0.157 mm3 of weight loss, for ATZ and Al2O3, respectively) [32]. Instead, similarly with the ZTA case [27], the ATZ material showed a net improvement among Al2O3 under adverse in vivo–like loading conditions (i.e., microseparation) using the hip simulator, where ATZ/ATZ outperformed the Al2O3/Al2O3 couple bearing with a wear rate of 0.06 mm3/million cycles compared with the 0.74 mm3/million cycles of monolithic alumina alternative [33].

Table 8 Mathys ATZ Ceramys ®mechanical properties

Property Al2O3 content ZrO2 content Density Average grain size Al2O3 Average grain size ZrO2 Flexural strength Young modulus Fracture toughness Hardness (HV20)

Units Vol % Vol % g/cm3 μm μm MPa GPa MPa m½ GPa

Ceramys ® ATZ 20 80 5,51 0.40 0.40 >1200 n.s. 7.4 15

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Stability of Alumina–Zirconia Composites The resistance to environmental degradation in an alumina–zirconia composite (BIOLOX ®delta) has been discussed in depth by Chevalier et al. [34]. The results show that even if an increase of monoclinic content is expected after long aging duration in vivo, the impact of the transformation is far from being the same that of monolithic zirconia, due to the absence of percolative phenomena in this material, as described in section “Structure of Zirconia-Toughened Alumina.” After accelerated hydrothermal treatment, the composite does not show large surface uplifts, as it is the case for 3Y-TZP, and increase in surface roughness is not detectable. Also, the mechanical behavior remains unaffected. As stated at the beginning of the chapter, the delay in the onset of the low temperature degradation of the zirconia phase in zirconia-rich alumina–zirconia composites was not elucidated yet. Also for this reason, Ceramys ® ATZ composite ceramic was extensively tested in adverse aging conditions. However, the presence of 3 mol.% yttria stabilization as that one usually used in monolithic zirconia materials may put in evidence the wish of the designer of the material to slightly overstabilize the zirconia phase, in order to avoid the occurrence of low temperature degradation, limiting on the other hand the mechanical performances. The presence of a zirconia cubic phase and very low monoclinic zirconia content in pristine conditions [30] is a consequence of the above assertion. Thus, the material exhibits a much slower overall aging kinetics in comparison with the 3Y-TZP. Nevertheless, because of the presence of the percolation mechanism, the zirconia composite seem to show a significant roughness (i.e., Ra) increase as the monoclinic phase volume content increases during the hydrothermal exposure [30]. However, the remarkable mechanical properties with the controlled zirconia t ! m transformation kinetics might make of ATZ material an appealing alternative in THA.

Biocompatibility of Bioinert Ceramics The biological safety of alumina and zirconia are well established since a long time. The high oxidative status gives high chemical inertia to both ceramics, especially to alumina which is frequently taken as a negative reference in biocompatibility tests of candidate biomaterials. Tests on alumina, zirconia, and alumina–zirconia composites were performed varying a number of different test parameters. The first remark to be made is that the materials tested were different both in physical forms (powders or dense ceramics) and in chemical and physical characteristics (e.g., reactive surface, chemical composition, impurity content, etc.). Further remarks are concerning test conditions: the in vitro assays were performed using extracts in various media, either in direct or indirect contact, while the cell lines used (e.g., macrophages, lymphocytes, fibroblasts, and osteoblasts) were stable from primary cultures. Similar consideration can be made on the in vivo tests that had been

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performed in several sites of implantation in different animal models, to analyze adverse reactions in soft tissue and/or bone, as well as systemic toxicity. Comprehensive reviews are reported in [2, 5, 35].

In vitro Tests In testing materials in vitro, the specific surface (i.e., surface for a given volume or mass of the material) was found to be a relevant parameter, because particulate can develop some square meters per gram of reactive surface. For the same reason, the size of the particulate is a relevant parameter in these tests, because a given mass of particulate develops a higher surface as the particle size decrease. Similar consideration can be made in tests performed on porous materials. The concentration of powders resulted relevant for cellular response. Dosedependent cytotoxicity was observed on human lymphocytes stimulated with phytohemoagglutinine in presence of partially stabilized zirconia and alumina powders as well as on fibroblasts cocultured with ceramic extracts. The concentration of ceramic powders has been reported as relevant to cellular apoptosis, e.g., in macrophages by Catelas et al. [36]. Kranz et al. [37], while challenging corundum particles to murine macrophages also observed cytotoxic reactions depending on the particle size, the more cytotoxic being the nanometric ones because of their very high specific surface (Table 9).

In vivo Tests The tissue reactions to injection of alumina or zirconia particles are much lower than after injection of metal or polyethylene particles, also at high specific surface of the material under test, e.g., after injection of alumina particles, the fibrosis observed after injection of polymeric or metallic particles was absent [38, 39]. Implants of either alumina or zirconia in bone performed in nonloaded sites show that bone comes into contact with the ceramic without fibrous interposition [35, 38]. Absence of acute systemic adverse tissue reactions to ceramics were also reported after subcutis, intramuscularor intraperitoneal, and intra-articular injection of alumina and zirconia powders in rats and/or mice, as well as by implants made into the paraspinal muscles of rabbits, rats, or in rabbit’s bone (Table 10).

Carcinogeneticity In 40 years of clinical use of alumina and zirconia, only one author discussed the possible link between soft tissue sarcoma and the presence of alumina ceramic bearing in a hip implant [40]. During the following years, the issue of carcinogenesis induced by ceramics has been raised especially for zirconia, due to the

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Table 9 Summary of the tests in vitro performed on bioinert ceramics and composites First author Harms

Publication year 1979

Material tested Alumina

Physical form Powder

Pizzoferrato

1982

Alumina

Ceramic

Bukat

1990

Ceramic

Greco

1993

Powder

Human lymphocytes

Ito Li

1993 1993 1994

Wear debris Powder & ceramic Powder

L929 Human oral fibroblasts

Dion Harmand

1995

Catelas

1998

Mebouta

2000

Torricelli

2001

Lohman

2002

Karlsson

2003

Alumina & Zirconia Alumina & Zirconia Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Alumina

Cell type Macrophages Lymphocytes HeLa Fibroblasts 3 T3 Fibroblasts 3 T3

Gutwein

2004

Alumina

YagilKelmer Carinci Hao

2004

Powder

HUVEC Fibroblasts 3 T3 Fibroblasts 3 T3

Powder

Macrophages J774

Powder

Human Monocytes

Powder

Osteoblasts

Powder

Osteoblasts MG63 Primary osteoblasts

Alumina

Nanoporous ceramic Nanoporous ceramic Powder

2004 2005

Zirconia Zirconia

Ceramic discs Ceramic discs

Baechle

2007

Zirconia

Ceramic discs

Sollazzo Kranz

2008 2009

Zirconia Alumina

Colloidal coating Powder

Att Kohal

2009 2009

Zirconia Zirconia & ATZ

Ceramic discs Ceramic discs

Maccauro Roualdes

2009 2010

ZTA ZTA

Ceramic discs Ceramic discs

Cho

2014

Zirconia

Ceramic discs

Osteoblasts CRL11372 Human monocytes U937 Osteoblasts MG63 Human fetal osteoblasts Osteoblast-like CAL72 Osteoblasts MG63 Macrophages RAW264.7 Primary osteoblasts Primary human osteoblasts Fibroblasts C3H10T½ Primary human fibroblasts Osteoblast-like MC3T3-E1

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Table 10 Summary of in vivo biocompatibility studies on bioinert ceramics First author Helmer Hulbert

Publication year 1969 1972

Material tested zirconia Alumina & Zirconia

Griss

1973

Alumina

Physical form Pellets Discs, tubes Dense & porous Slurry

Animal model Monkey Rabbit

Harms

1979

Alumina

Powder

Swiss mice Mice

Garvie Wagner

1984 1986

Ceramic bars Pins

Rabbit Rat

Christel

1988

Pins

Rat

Paraspinal muscle

Christel

1989

Pins

Rabbit

Femur

Maccauro Specchia Warashina

1992 1995 2003

Powder Pins Powder

Mice Rabbit Rat

Intraperitoneal Femur Calvaria

Scarano Sennerby Graci

2003 2005 2011

Tibia Tibia, femur Knee joint

2013

Screws Screws Low-cohesion Cramic Screws

Rabbit Rabbit Rabbit

Marques

Zirconia Alumina & Zirconia Alumina & Zirconia Alumina & Zirconia Zirconia Zirconia Alumina & Zirconia Zirconia Zirconia Alumina & Zirconia Zirconia

Subcutis, Knee joint Intraperitoneal, intramuscular Paraspinal muscle Femur

Rabbit

Tibia

Implant site Femur Paraspinal muscle

presence of radioactive contaminants, as discussed in section “Zirconia Mechanical Properties” of this chapter. Covacci et al. [41] investigated the possible mutagenic and oncogenic effects of a zirconia ceramic (Y-TZP) sintered from high purity powders. The results obtained showed that Y-TZP ceramic does not elicit either mutagenic or transforming effect on mouse embryo-derived fibroblasts C3H10T½, confirming the suitability of Y-TZP for application in medical devices. Maccauro et al. [42], who also challenged C3H10T½ with alumina–zirconia composite ceramic samples, has reported the absence of cellular DNA damage, mutagenicity, and carcinogenicity as well.

Reaction to Ceramic Wear Debris Because alumina, zirconia, and their composites are used mainly in hip replacement bearings (see Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacement” in this Handbook), the biocompatibility of their wear debris was accurately characterized. Wear debris trigger a cascade of biological reactions that can eventually result in osteolysis – i.e., the loss of the mineral part of the

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bone – and eventually in the aseptic loosening of the implants, making mandatory a revision surgery. Due to the economic and social relevance, researchers studied this pathology since a long time. Since 1977, Willert and Semlitsch [43] observed granulomas associated with necrosis and fibrosis due to particles accumulating in periarticular tissues that can migrate by the blood stream giving rise to systemic reactions. This was also reported by Schmalzried and Callaghan [44] who identified an “effective joint space” where the wear debris dispersed in the fluid surrounding the joint can be transported due to the pressure gradients generated by the motion of the articular joint during the patient’s daily activity. Because ceramic joint replacement bearing is especially suited for active patients, like e.g., the younger ones with longer life expectations, the long-term effects of ceramic wear debris becomes more and more relevant. It is acknowledged that local and systemic reactions to ceramic wear debris are reduced in comparison to the ones observed in presence of metal and polyethylene (UHMWPE) debris. Alumina wear debris has been found less active than UHMWPE debris in promoting the differentiation of osteoclasts, especially in the low concentration typical of the low wear of alumina bearings [45]. These findings are well corresponding to the ones reported by Fisher et al. [46] who demonstrated the negligible biological activity of ceramic-on-ceramic wear debris in comparison with standard polyethylene (60) and to cross-linked polyethylene (13). The migration of wear debris to peripheral organs implies a high vascularization of the periprosthetic membrane that contains the debris. This feature is depending on inflammatory tissue response related to the cytotoxicity of the particles and to the release of humoral factors after phagocytosis by macrophages. A specific aspect of the tissue reaction to ceramic debris is the formation of fibrous tissue poor in vessels in absence of contaminations from metals or polymers. Membrane surrounding ceramic debris was observed by Maccauro et al. [47], as formerly did Lerouge et al. [48] who reported that foreign body cells are typically associated to polyethylene, PMMA, and large metal particles, but rarely to ceramics. The absence of vessels in the periprosthetic tissue surrounding alumina particles in alumina-alumina joints led De Santis et al. [49] to hypothesize that ceramic debris did not migrate from the production site. Graci et al. [50] observed membrane poor in vessels surrounding ceramic debris. This finding was not depending on the nature of the ceramic tested (alumina either zirconia Y-TZP), nor on the release time of the material. Namely, periprosthetic tissue membrane surrounding ceramic debris was poor in vessels either if powders were massively inserted directly into the joint, either if debris are progressively released for a long time, by means of an original testing method developed by these Authors. All these findings are explaining the absence of local systemic toxicity of ceramic debris in the long term, and the positive outcomes reported in the long-term followups of several series of implants with ceramic-on-ceramic either ceramic-on-polyethylene bearings (see Chap. 28, “▶ Perspective and Trends on Bioceramics in Joint Replacemen,” this Handbook).

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Processing of Bioinert Ceramics for Joint Replacements The bioceramics used in joint replacements are belonging to the wide class of materials known as hi-tech ceramics (advanced ceramics, fine ceramics), that are obtained by precursors in form of powders obtained through the synthesis of high purity chemicals. Many synthesis processes are in use, based on solid state reactions, on gas –solid reactions, or on wet routes, e.g., coprecipitation from salt solutions, or sol-gel processes. Precursors are formed by crystallites having size ranging from 0.01 to 0.1 μm that are forming small particles (grains, or domains) 0.1–1 μm in size. Grains may be packed together forming bigger agglomerates 10–100 μm in size. Chemical, crystallographic, morphologic and bulk characteristics of powders are relevant in ceramic processing and are depending on the process used for their preparation. Grain size is the parameter most frequently reported to characterize a powder, for its strong influence on the microstructure of the final product. A relevant parameter linked to grain size is the specific surface area that gives a rough indication of the reactivity of the powder: current values for ceramic precursors are in the order of some square meters per gram. Also the flowability of the powder is relevant for processing, as it indicates its suitability to free flow in feeding and filling the dyes used for powder consolidation, without forming coarse agglomerates. Some of the additives that determine the characteristics of the ceramic are introduced in the powder during its synthesis, like e.g., yttria Y2O3 in Y-TZP, or magnesia MgO in PSZ precursors. The precursors’ main characteristics are: – High specific surface, in the order of some tens square meters per gram, a feature that makes them very reactive – Uniformity of crystallite size – Homogeneous chemical composition – Flowability, e.g., the ability of the powder to flow freely, without formation of agglomerates or voids during the powder consolidation – Ultralow content of impurities A general flow sheet of the process used to transform the precursors in a solid body can be summarized in four main steps: 1. Batch preparation: the precursors are blended with a suitable mix of process additives depending on the consolidation route and on the desired final behavior of the product. Magnesia (MgO) is a sintering aid commonly used to limit grain growth in alpha-alumina. Binders and lubricants are also added to the powders for dye pressing in the compaction process used in shaping the product. The ceramic composition with a consistence of slurry is then eventually milled to homogenize the composition and narrow the particle size distribution. The blend is then treated by spray drying to obtain the feedstock (e.g., in form of granules) to be used in the next step of the process.

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2. Compaction: the feedstock (e.g., ready-to-press powders) is shaped in simple geometry (e.g., a cylinder of adequate size) by dye pressing. The product of this step (the greens) is then machined to near-net-shape. An alternative route is the near-net shaping by injection molding, a process for obtaining more complex geometries but that implies an high content of plasticizers (waxes) in the batch to achieve the necessary flowability and the proper filling of the mold. Typical density of the greens at the end of this step is about 50 % of the Theoretical Density (or 50 %TD). 3. Consolidation. The greens reach the reference density by sintering. This process involves minimization of the surface energy through a high- temperature thermal treatment, resulting in a solid body. In the first phase of a sintering cycle, binders are evacuated via the open porosity then porosity is evacuated from the body, leaving a polycrystalline solid with a density nearing 100 % T.D. (Typical densities of ceramic head and inlays are 99.5 %T.D. or more). The process is carried out in a furnace where the green body is fired to temperatures about 2/3 of the melting temperature of the oxide. The sintering aids introduced into the feedstock allow enhancing transport phenomena and limit grain growth The sintering process usually is carried out in two steps. The first step (presintering) consists in heating the green body to intermediate temperature to evacuate the organic additives. As a result, the green body shrinks reaching final density some 70–80 % TD. Presintered ball heads or inlays may then be obtained by mechanical tooling, which is currently made using numerical control machines Machined parts are then inserted in a high-temperature furnace (some 1500  C) where in an appropriate time/temperature sequence they reach the final density (e.g., > 99.9 %TD). Sintering is usually performed in oxidizing atmosphere in gas-heated kilns but continuous (tunnel) furnaces are also in use for high volume production. In these furnaces, the green bodies are loaded on trays that are transported across a tunnel which is heated at different temperatures: timetemperature sequence depends on the trays’ speed. In both cases, final density is achieved by HIP (Hot Isostatic Pressing) treatment in high pressure batch furnaces. HIP consists in a high temperature plus high pressure treatment (e.g., 1100  C, 1000 Bar). HIP is a relevant innovation introduced in bioceramics processing to obtain ceramics with density near to the theoretical one. Because the densification takes place at relatively low temperature, the deleterious grain growth that takes place when the process is made at higher temperature is avoided. In addition, HIP treatment allows minimizing the residual stresses within ceramic pieces thus further improving the strength and reliability of products. 4. Machining and finishing. The classic process to obtain ceramic heads or inlays is based on the machining of green cylinders obtained by bidirectional dye pressing of ready-to-press powders. As by this process a big volume of valuable materials is wasted, some producers use the alternative cold isostatic pressing method. After cooling, the spherical surface is polished to the desired grade, the inner cone is ground to the expected tolerance (Table 11), identification marks made by laser, and the lot is sent to final controls prior to be released to market.

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Table 11 Typical values of finish of a THR ceramic ball head Part Spherical surface

Inner taper

Characteristic Deviation from sphericity Deviation from nominal diameter Roughness Ra Roughness Rmax Deviation from straigthness Deviation from roundness Roughness Ra

Value [μm] 10 % abundance and at >75 % ubiquity) can be summarized in two groups: Streptococcus (OTU, 2, 6) and unclassified Pasteurellaceae (OTU, 19, , 16). Further major core genera (>1 % abundance at >80 % ubiquity) include Gemella (OTU, 11), Veillonella (OTU, 4), Prevotella (OTU, 10), Fusobacterium (OTU, 9), Porphyromonas (OTU, 7), Neisseria (OTU, , 8), Capnocytophaga (OTU, ), Corynebacterium (OTU, , 15), unclassified Neisseriaceae (OTU, 21), Actinomyces (OTU, 14), and unclassified Lactobacillales (OTU, 13) [3]. A single OTU dominated nearly all oral mucosal sites of this large cohort: Streptococcus (OTU, 2). Thus, about half of the total cultivable flora comprises oral streptococci, which can be detected on almost all surfaces of the oral cavity; these are dominated by S. mutans, i.e., the pathogen primarily responsible for dental caries. Other Grampositive cocci, such as enterococci and staphylococci, are usually in less abundance, as are Actinomyces and lactobacilli, in turn the most frequently detectable Gram-positive rods. The Gram-negative cocci species Neisseria, seldom implicated in dental diseases, are also a common finding, with an abundance equal to that of Actinomyces spp. The most frequent Gram-negative oral rods are Haemophilus spp. and Aggregatibacter spp. Of note, Aggregatibacter actinomycetemcomitans has been associated with aggressive forms of periodontal diseases, and the relative abundance in the oral microbiome of healthy subjects appears negligible. Considering other bacteria involved in the pathogenesis of periodontal diseases, i.e., Porphyromonas species, Treponema denticola, and Fusobacterium nucleatum, evidence exists to support their drastic increase in dental plaque when the appropriate mechanical oral hygiene procedures are not performed. (ii) Fungi. Candida is the main fungal component of the oral environment, also being found in healthy people because of its commensal feature, together with further genera, such as Aspergillus and Saccharomyces, detectable as a minor component. Candida albicans is the most commonly isolated yeast species, followed by other clinically relevant “non-Candida albicans species,” which include Candida tropicalis, Candida krusei, and Candida glabrata. Candida is a usual component of the oral biofilm, and its relative abundance in dental plaque is particularly high in patients with oral candidiasis: as mentioned above, one of the most common clinical pictures is prosthetic stomatitis, occurring in areas beneath the resin base of removable dentures. (iii) Mycoplasmas. These pleomorphic microorganisms differ from other oral bacteria in that they lack an outer membrane. They have been isolated from the oral cavity, the most typical species being Mycoplasma pneumoniae, considered a surface parasite; it may be an etiological factor in infections of the upper respiratory tract, mainly in immune-compromised patients.

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Development of the oral microbiome – The oral microbiome as it exists today can be seen as the product of microorganisms’ long adaptation in cohabiting within the human body. From this fascinating perspective, microorganisms have been tailored to live in human organisms under a mutually beneficial symbiosis between microorganisms and human tissues [3]. This coevolution of the oral microbiome with the human host has resulted in a process known as “colonization resistance”: this term describes the ensemble of host-associated microbial communities, fully equipped with mechanisms enabling them to prevent colonization by and establishment of foreign microbes. Five principal types of interactions among oral bacteria have been identified, namely, competition for nutrients, synergy, antagonism, neutralization of virulence factors, and interference in signaling mechanisms. Bacterial interspecies communication is a cornerstone of colonization resistance, together with a broader inter-kingdom communication, both processes being crucial in oral microbial ecosystem homeostasis. For instance, biofilm formation by C. albicans appears to be partially regulated by certain bacteria that produce a range of selective signaling molecules; C. albicans’s metabolites are, in turn, compounds known to be able to influence bacterial growth. The acquisition of a normal, beneficial oral microbiome, including the process of colonization resistance, is an essential step in the growth of newborns. The oral microbiome in infants is closely connected to that of the gastroenteric tract, but after 2 weeks of life, it already differs as the oral cavity is rapidly colonized by bacteria originating from the environment where the newborn lives. Bacterial transfer from the mother, or from other external sources, including other people sharing the same environment, greatly affects dental biofilm morphogenesis. As recently reviewed by Zaura et al. [3], the key aspects pivotal for the physiological acquisition of a normal microbiota during development are: (i) Vertical transmission of the microbiome from mother to child starts with delivery, whether vaginal or through Caesarian section, which to a large extent determines which microorganisms initially colonize the infant’s oral cavity (vagina- or skin-derived). Infants born by Caesarian section acquire Streptococcus mutans earlier than vaginally born infants, while vaginal birth enables newborns to acquire a greater bacterial taxonomic diversity by the third month of life. Similarly, breast-feeding versus infant formula feeding appears to influence acquisition of the oral microbiome; breastfeeding gives the infant “beneficial” oral lactobacilli that are not detectable in formula-fed infants. (ii) Preservation of the oral microbiome, after acquisition during the first stage of life, involves bidirectional interactions between the microbiome and the host; thus, the human immune system develops in a continuing dialogue with the commensal populations of microbiota. This communication exploits the following three main ways: the first includes the host pattern recognition receptors (PRRs), especially the toll-like receptor (TLR) family, expressed by oral mucosa cells, i.e., keratinocytes, macrophages, mucosal dendritic cells (DCs, which belong to the Langerhans cell subtype), polymorphonuclear leukocytes, and natural killer cells. Altered expression patterns of TLRs have been found in

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several dental and oral diseases, suggesting their specific role in pathogenesis, while it has been suggested that mucosal DCs are peculiar intermediaries able to avoid infectivity of the oral cavity by commensal microbiota. A second tool to stimulate antigenic tolerance, and thus avoid the risk of local infectious disease, is the expression of lipopolysaccharide (LPS) receptors CD14, TLR2, and TLR4 by DCs, at the level of the non-inflamed oral epithelium. Finally, chemical sensing is the third pivotal tool that the host can exploit to monitor microbial activity. In recent decades, studies have suggested there may be a direct link between secreted bacterial products and chemosensory activation mechanisms for mucosal clearance. The fundamental role of the immune system in preserving oral health becomes increasingly evident when examining the impaired situation due to the patient’s pathological status; typical examples are those of patients receiving hematopoietic stem cell transplant and who require immunosuppressive therapy or of patients affected by head and neck carcinoma and treated with local radiotherapy. One of the most severe and painful adverse effects is the mucosal damage known as “severe mucositis,” which is potentially associated to life-threatening viral and fungal supra-infections. (iii) The secretory immunoglobulin A (S-IgA), usually delivered via the saliva and gingival crevicular fluid, limits and controls microbial adhesion and colonization. Conversely, bacterial ability to evade S-IgA guarantees their survival within the oral cavity, again highlighting their ongoing symbiotic coevolution with the human body host. S-IgA elusion is mainly achieved through bacterial IgA proteases, which neutralize the immunoglobulin. These proteases are known virulence factors of several human pathogens, such as Neisseria meningitidis and Streptococcus pneumoniae, and of other commensal streptococci (Streptococcus mitis, Streptococcus oralis, and Streptococcus sanguinis). The latter have been defined as “primary colonizers” and are also the foremost species in infants. (iv) Salivary flow rate and saliva composition also play key roles in maintaining the healthy oral microbiome. Focusing on protein composition of the saliva, microbial homeostasis is strongly affected by the presence of salivary glycoproteins, because they contain glycans that may act as traps to prevent pathogens from adhering to epithelial cells. Other salivary proteins that influence the oral microbiome include lysozyme, peroxidase, mucins, lactoferrin, defensins, and agglutinins. The oral cavity as a biological niche – From the topographical standpoint, two main subniches are described in the oral cavity: the supragingival niche and the subgingival niche. The supragingival niche includes the teeth or implants and the mucosal tissue outside the gingival sulcus. The plaque recovered from this area in healthy subjects generally comprises aerobic Gram-positive bacteria, mostly Streptococcus spp. (Streptococcus sanguinis, Streptococcus mutans, Streptococcus mitis, Streptococcus salivarius) and lactobacilli. In contrast, the subgingival niche (i.e., the gingival sulcus) is characterized by the presence of some Gram-negative microaerophilic bacteria, in addition to Gram-positive and aerobic species. Many

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Table 2 Bacterial species recoverable subgingivally in healthy subjects Gram-positive Actinomyces Clostridium Lactobacillus Staphylococcus Streptococcus

Morphotypes Rod Rod Rod Coccus Coccus

Gram-negative Bacteroides Fusobacterium Neisseria Prevotella Treponema Veillonella Wolinella Eikenella Aggregatibacter Porphyromonas Tannerella Campylobacter Capnocytophaga

Morphotypes Rod Rod Coccus Rod Motile rod Coccus Rod Coccus Coccus Rod/coccus Rod Long rod Rod

of these are rods, with some motile bacteria and facultative intracellular bacteria (e.g., Porphyromonas gingivalis) [4]. Table 2 lists the most frequent genera and species recoverable from the sulci of healthy subjects. These microorganisms are natural commensals of the oral cavity, where they are found either in their planktonic form or within structured and complex 3D biofilm communities. The formation of the biofilm community is a key factor in the transition of bacteria from commensals to putative pathogens. When bacteria grow in the biofilm, they may accumulate high concentrations of bacterial metabolites (e.g., fatty acid end products, ammonia, hydrogen peroxide, oxidants, and carbon dioxide) in their local environment, which influence the prevalence of species both within the microbial community and in the host. For instance, as already mentioned, carious lesions are closely related to certain biofilm-forming bacteria, mainly Streptococcus mutans, which is able to adhere to the teeth, proliferate, and produce lactic acid, which in turn can dissolve the mineralized components of enamel and dentine. In the presence of sugar, S. mutans overwhelms the other non-acid-producing Streptococcus spp. that make up the supragingival plaque. Actinomyces spp. are among the dominant taxa in both the supra- and the subgingival plaque, from both healthy subjects and periodontitis patients. Porphyromonas gingivalis, Bacteroides forsythus, and Treponema denticola have been detected in supragingival and subgingival plaque samples of both healthy subjects and individuals affected by periodontitis, although they are significantly more prevalent in both supra- and subgingival plaque samples from the latter.

The Oral Microbiome in Oral Diseases: A Focus on Implant and Prosthetic Dental Materials In the light of the concepts described above, biological properties that confer stability to the microbiome are important for the prevention of disease-related

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“dysbiosis,” producing the microbial shift toward periodontitis or carious lesions. Although the processes underlying the healthy equilibrium of a normal microbiome remain poorly understood, the mechanisms that underlie oral diseases have been investigated in depth; in particular, research has focused on oral microbiome changes that occur on the surface of implants and prosthetic dental materials. In addition, the surface adhesion of microbes, such as bacteria and fungi, and the subsequent formation of biofilms contribute to multidrug-resistant infections in humans and, consequently, to the failure of medical devices. Peri-implant microbiome – Having been the object of numerous high-quality studies, osseointegrated dental implants are today a therapeutically successful option in prosthetic dentistry, for the rehabilitation of complete, partial, and single edentulism. Oral implantology is based on technologically advanced devices, highly customized to replace missing teeth, satisfying both functional and esthetic requirements. Implant rehabilitation may be considered one of the foremost discoveries of the twentieth century; however, from the oral microbiome perspective, dental implants also represent new artificial surfaces within the oral cavity, which appear more prone than natural tooth surfaces to form bacterial biofilms. Dental plaque, similar to what occurs on natural teeth, can easily accumulate. Biofilm formation on the implant surface is a trigger factor for the further inflammatory process of peri-implant tissues, namely, peri-implant mucositis (when the inflammation only involves the periimplant mucosa) or peri-implantitis (when the inflammation progresses toward the surrounding alveolar bone). The oral microbiome in peri-implant infections has been studied by conventional, molecular, and metagenomic analyses. Using the 16S rRNA-based PCR detection method on crevicular fluid samples, the biofilm adhering to abutments showed the presence of both A. actinomycetemcomitans and P. gingivalis [5]. Moreover, the oral microbiota growing on dental and implant surfaces has recently been investigated in partially edentulous patients, in a large, 10-year retrospective clinical trial, on 504 implants and 493 adjacent teeth [6]. The microbiota analyses of dental plaque specimens, collected after the placement of sandblasted and acid-etched implants, revealed the presence of some bacterial species associated with periodontitis, such as aerobic Gram-negative rods and staphylococci, although abundances were very wide ranging (from 6.2 % to 78.4 % of implants). The study authors reported a higher abundance of Tannerella forsythia, Parvimonas micra, Fusobacterium nucleatum/ necrophorum, and Campylobacter rectus at implant sites than on dental surfaces. Based on these data, the prevalence of Prevotella intermedia, Treponema denticola, Campylobacter rectus, and Staphylococcus warneri has been suggested to be associated with peri-implantitis. In addition, comparing smokers versus nonsmokers, the latter showed higher counts of periodontopathogenic species; similar to the comparison between periodontal versus non-periodontal patients. These latter findings again support the role of the two major risk factors, i.e., smoking and periodontal disease, in the pathogenesis of peri-implant inflammation. Considering the composition of the oral microbiome, although some evidence suggests that the miscellaneous microbial flora of peri-implant infections may bear a

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resemblance to that of periodontal infections, some recent studies suggest there may be certain differences. It is likely that future breakthroughs will occur with the increasing application of metagenomics and metatranscriptomics. These innovative technologies have recently been applied to evaluate the microbiota associated with osseointegrated implants and to investigate peri-implant disease pathogenesis. The current state of the art was reviewed by Charalampakis and colleagues, who analyzed the existing knowledge on peri-implant microbiology and the diversity of the microbial communities associated with peri-implantitis [7]. The peri-implant microbiome in healthy individuals includes a preponderance of Gram-positive cocci and nonmotile bacilli, with a small number Gram-negative anaerobic species, similar to what occurs in the normal gingival tissue. The switch to the first step of inflammation around implants, i.e., peri-implant mucositis, correlates with the increased presence of cocci, motile bacilli, and spirochetes, to an extent equivalent to that of gingivitis. Conversely, the further shift to peri-implantitis is mainly related to the appearance of Gram-negative, motile, and anaerobic species, which are frequently detected in periodontitis. Through molecular biology, A. actinomycetemcomitans, Porphyromonas gingivalis, Prevotella intermedia, and several Fusobacterium spp. have been detected in dental implant plaque specimens. It has been suggested that, in general, the bacterial profile of peri-implantitis derives from periodontitis, since most peri-implant lesions shares common features with periodontal disease. In particular, the so-called “red complex” group of periodontopathogens (Porphyromonas gingivalis, Tannerella forsythia, and Treponema denticola) was found to be more abundant at sites affected by peri-implant disease than at healthy ones. Conversely, the count of S. aureus, markedly higher at implant sites than at others, supports the possibility of detecting unique and distinctive microbiological features related to peri-implantitis. Indeed, although the metagenomics approach has yet to provide robust data, owing to the paucity of investigations, emerging data support the view that the peri-implant microbiome is a specific entity, different from the periodontal microbiome. Interestingly, further evidence suggests that implant sites and adjacent teeth appear to share similar microbiota, probably because they are spatially close and comparable ecological niches. Indeed, in the case of fully edentulous patients, healthy implants displayed similar bacterial colonizers as do healthy periodontal sites. However, in the case of partially edentulous patients, the implant surface was colonized by the same species as the adjacent teeth and oral mucosa. In addition, A. actinomycetemcomitans and P. gingivalis, usually detectable only in the presence of teeth, were detected in periimplantitis in fully edentulous patients, indicating that the bacterial species might originate from niches in the oral cavity other than the subgingival sites, such as the soft tissues or saliva; alternatively, these bacteria might remain in place after tooth extraction and subsequently colonize the oral surfaces, including dental implants [8]. Dental caries and the oral microbiome – Dental caries have been investigated microbiologically, at the molecular level, in a number of studies, and the principal findings have been summarized in a recent review by Nyvad and colleagues [1]. In order to extend scientific knowledge of this common disease, cariology

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progressively exploits the new and complementary approaches, including metagenomics, metatranscriptomics, metaproteomics, and metabolomics analyses of dental biofilms, along with refined microbial sampling techniques. One of the priorities for caries microbiologists in the near future will be to verify the performance, and not just the composition, of the entire microbial community. In particular, the metabolism resulting from the activities of oral microbiota greatly affects the dynamic processes of caries. Chiefly for this reason, metabolomics is expected to acquire a decisive role in this field, to facilitate research in assessing bacterial functions. Integrated approaches will make it possible to assess which genes are expressed and which phenotypic characteristics of the biofilms are detectable at specific dental sites, since caries is a localized disease. Taking into account that one of the major difficulties is sample collection, it is essential that biofilm specimens be taken from specific and specified tooth sites. Indeed, the use of pooled samples has been found not to be appropriate, since the bacterial inoculum collected from salivary samples of patients with different caries experiences would be unable to provide insight into the cariogenic potential of site-specific biofilms. Unfortunately, most molecular studies on caries have used saliva samples or pooled plaque samples. In a recent study, the 16S rRNA gene was cloned and sequenced, in order to characterize the microbial composition of the oral biofilm in the presence of carious lesions. Custom-made arrays, specifically targeted to individual patient groups, detected a microbial diversity in patients’ subgingival plaque. The main methodological limitation of this technique is that only those microorganisms specifically targeted by the probes can be detected. Samples collected from healthy and carious root surfaces of older patients were analyzed for their taxonomic microarray, showing that great bacterial diversity and the presence of Actinomyces spp. were more frequent at healthy sites, whereas several species of lactobacilli and Pseudoramibacter alactolyticus were associated with root caries [1]. A further study on the transcriptome determined a functional core microbiota, consisting of about 60 species; it identified numerous functional networks and provided support for the hypothesis that interindividual environmental differences affect the selection of microbial groups. Dominant functions of bacteria, such as the capacity of dental plaque microbes to metabolize diverse sugars and to handle the acid production and oxidative stress that result from sugar fermentation, were expressed by the oral microbiota [9]. Metagenomics and metatranscriptomics, via pyrosequencing analyses, can retrieve millions of partial 16S rRNA gene sequences in one sequencing run; they have been used in a cross-sectional study to analyze the oral microbiota of Chinese children with and without dental caries. The findings supported the hypothesis that the presence in the plaque of the genera Streptococcus, Veillonella, and Actinomyces is significantly associated with dental caries. Focusing on adulthood, the comparison between “healthy” and “cariogenic” salivary microbiome revealed that the latter was significantly more variable in terms of community structure. This outstanding result, i.e., that “healthy” microbiomes are more preserved than caries microbiomes, was consistent with other evidence from a study applying microarrays to analyze the microbial composition of saliva in children in relation to their caries status [1].

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Removable denture oral microbiome – Changes of the oral microbiota before and after wearing removable dentures (RD) appear possibly related to a local imbalance of the microbial community, leading to oral candidiasis; however, there is as yet no certainty. Possible variations in the human oral bacterial community related to wearing partial RD have been analyzed in the four main kinds of biological oral specimens: saliva, supra- and subgingival plaque, and oral mucosal surfaces. A recent study collected these four types of plaque samples from RD wearers (n = 10) at three different times, i.e., before and after 1 and 6 months of wearing RD; a further ten healthy adults were selected as control group [10]. After cloning and sequencing, the health-associated genera, such as Streptococcus, Neisseria, Corynebacterium, Gemella, Veillonella, Selenomonas, and Actinomyces, showed a decreasing trend in RD wearers, while species associated to disease, mainly Streptococcus mutans, appeared to increase. Considering that Candida-related prosthetic stomatitis is correlated to a marked elevation in the number of Candida species cells present on the acrylic base of dentures, an interesting recent trial investigated the relationship between the Candida load and the bacterial diversity, in the saliva of older patients [11]. Patients were partially edentulous, with or without partial RD, or edentulous, with total upper and lower RD: almost all subjects were positive for Candida, with a negative correlation between Candida load and bacterial profiles of the saliva. When the Candida load increased, the diversity of the salivary microbiome decreased, and its composition shifted toward dominance by streptococci and lactobacilli, while genera within the Fusobacteria and Bacteroidia classes disappeared. Decreased bacterial variety was associated with a lack of equilibrium among the microbiome communities.

Definition, Structure, and Composition of the Biofilm The oral microbiome, adhering to hard substrates, can assemble into threedimensional structures, called “dental biofilm” or “dental plaque”: the soft white material that may be observed on the surfaces of both teeth and dental materials. The term “biofilm” indicates a community of microorganisms adhering to a surface, glued into an extracellular polymer matrix, also known as “slime,” within which there are water channels. These channels generally consist of glycoproteins, proteins, and polysaccharides, which are secreted by the microorganisms themselves. Thanks to this complex and dynamic structure, the microorganisms acquire multiple properties, including improved protection against host defenses and against new invading microbes. Salivary proteins, adhering onto tooth surfaces and forming the dental pellicle, help microorganisms to bind to the surface, which is the first step of biofilm arrangement. Biofilms can form on both body tissues and material surfaces. Although mixed-species biofilms predominate in most environments, including the oral cavity, single-species biofilms exist in a variety of infections and on the surface of implantable medical implants such as orthopedic implants or

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catheters [12]. Indeed, the dental biofilm (corresponding to the oral microbiome) is composed of all the components of the oral microbiota and may thus comprise a single or multiple microbial species, mainly bacteria, fungi, and mycoplasmas. The saliva can also contain certain types of protozoa, such as Trichomonas species, but mainly in immunocompromised subjects. In the process of biofilm development, Gram-positive bacteria, such as Streptococcus spp. and Actinomyces spp., are called “pioneer species,” since they are usually the first to adsorb onto the dental pellicle and start to proliferate. They play an important role in producing conditions suitable for other microbes to further colonize the substrate; indeed, their respiration process reduces the oxygen tension and increases the level of carbon dioxide, resulting in hypoxic conditions that are suitable for anaerobic species. A number of oral microorganisms easily proliferate in this environmental setting: they are facultative anaerobes and account for most oral cavity bacteria, for example, oral streptococci, which survive deep within the dental biofilm.

Biofilm Metabolism Bacterial biofilm metabolism chiefly relies upon carbohydrates as principal source of energy, in order to produce ATP. In particular, glucose is converted to pyruvate via the glycolysis metabolic pathway; pyruvate then follows diverse pathways depending on the oxygen tension and the type of microorganism. For example, glucose is degraded to pyruvate via the central carbon metabolism, following the classic glycolysis reaction; however, under anaerobic conditions, it is further degraded into lactate and acetate, by bacteria including Streptococcus, Actinomyces, and Lactobacillus spp. Conversely, in the presence of oxygen, pyruvate is converted to acetate by Streptococcus and Lactobacillus, and lactate is converted to acetate by Actinomyces. When bicarbonate is also present, as often occurs in the saliva, phosphoenolpyruvate is converted to succinate with bicarbonate assimilation; for instance, Actinomyces follows this metabolic pathway. Lastly, in C. albicans pyruvate is directly metabolized into acetyl CoA by the pyruvate dehydrogenase complex in aerobic conditions, but under hypoxia the anaerobic route is activated, and small amounts of acetaldehyde may be produced. Since ethanol is toxic to microorganisms at high concentrations, the preferential metabolism is aerophilic and avoids ethanol accumulation. As has been said, biofilm metabolism is crucial for several dental diseases, and microorganisms’ metabolic pathways have been elucidated, using single bacterial strains in preclinical studies; they have not yet been confirmed by in vivo analyses on supragingival plaque. The main limitation on these studies is that the amount of supragingival plaque that can be sampled from the oral cavity is insufficient for a conventional metabolic study: metabolome analysis could be an excellent alternative to overcome this difficulty.

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Biofilm-Related Medical Device Infections The most common biofilm-related medical device infections are due to the Gramnegative Pseudomonas aeruginosa, Pseudomonas fluorescens, or Escherichia coli or to the Gram-positive Staphylococcus epidermidis, Staphylococcus aureus, or enterococci. Hospital and health-care facilities are peculiar environments in which dangerous antibiotic-resistant pathogens can live and evolve. Hospital-based pathogens show continuous dynamic change, and this influences their distribution through the body over time and their pathogenicity [13]. To fight the multidrug resistance (MDR) of several bacteria is still the major global challenge. Table 3 summarizes the strains correlated to hospital-based infections; to date, the MDR strains identified are the species Enterococcus faecium, Staphylococcus aureus, Klebsiella pneumoniae, Acinetobacter baumannii, Pseudomonas aeruginosa, and Enterobacter spp., collectively known under the acronym of “ESKAPE.” Hospital-based pathogens may infect the oral cavity and intraoral devices. At the beginning of the antibiotic era, hospital-acquired infections were mainly due to Staphylococcus spp., initially kept under close control by penicillin. Then, as Staphylococci started producing beta-lactamase, beta-lactamase-resistant compounds were synthesized in order to counteract these pathogens. Subsequently, methicillin-resistant S. aureus (MRSA) and Gram-negative bacilli emerged and became the chief bacteria responsible for hospital-acquired infections (HI); however, the use-abuse of antibiotics has favored the selection of bacteria with methicillin resistance combined with resistance to other types of antibiotics. In the late 1960s, Enterobacteriaceae, such as Escherichia spp., became increasingly involved in hospital-based infections, finally leading to the emergence of multidrug-resistant (MDR) Gram-negative Pseudomonas aeruginosa and Acinetobacter spp., causing very difficult therapeutic problems and a frustrating and never-ending search for a solution. The World Health Organization (WHO) has now recognized MDR as one of the three most important problems facing human health [14]. MDR is often due to the presence of specific resistance gene “islands” that, under the pressure of antibacterial agents, can be rapidly switched, developing a dynamic and always novel mechanism of antibiotic counteraction. In most cases, MDR strains attain these “islands” from bacteria of unrelated genera, as confirmed by sequence similarity and phylogenetic analyses [15]. This gives rise to the so-called super bacteria or super bugs, resistant to most, if not all, antibiotic regimes. However, the mechanisms underlying MDR vary in different pathogens, often reflecting the cellular structure of the bacterium.

Biofilm Formation and Propagation Oral biofilm formation is part of a biological cycle that includes four main stages: initiation, maturation, maintenance, and dissolution (Fig. 1).

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Table 3 Principal hospital-based infections, microorganisms, and human body sites involve Hospital-based infection (HI) Surgical skin (SSI) and soft tissue infections (SSTI)

Bloodstream infections (BSI)

Meningitis (MI)

Respiratory infections (RI) in intensive care units (ICU)

Endocarditis (EC) Gastroenteritis (GI)

Urinary infections (UI)

Genital/pelvic infections (GI)

Infective agents at different body sites Microorganism Staphylococcus (S.) aureus and epidermidis; Acinetobacter (A.) baumannii; Escherichia (E.) coli; Pseudomonas (P.) aeruginosa; Enterococcus (E.) faecalis; coagulase-negative Staphylococci; Candida (C.) albicans S. aureus; E. coli; Enterococcus spp.; Streptococcus (S.) spp.; Proteus spp.; Staphylococcus (S.) spp.; P. aeruginosa; Candida (C.) spp.; hepatitis B and C virus; Cytomegalovirus Enterovirus; herpes simplex type II; varicella-zoster virus; Adenovirus; parotitis virus; HIV; Flavivirus, Arbovirus; Neisseria (N.) meningitidis; Streptococcus (S.) pneumoniae; Haemophilus (H.) influenzae; S. aureus; P. aeruginosa; E. coli; Listeria monocytogenes; Cryptococcus neoformans; Histoplasma capsulatum; Coccidioides immitis; Blastomyces dermatitidis; Candida spp. Streptococcus (S.) pneumoniae; Haemophilus (H.) influenzae; Moraxella (M.) catarrhalis; S. aureus; P. aeruginosa; A. baumannii; E. coli; Legionella; Aspergillus (A.) fumigatus; Pneumocystis (P.) jirovecii; Mycobacterium (M.) tuberculosis, Klebsiella (K.) pneumoniae; Serratia (S.) marcescens S. aureus; Streptococcus (S) pyogenes and pneumoniae; E. faecalis; P. aeruginosa; Candida (C.) albicans Rotavirus; Campylobacter; S. aureus; Pseudomonas (P.) aeruginosa; E. coli O157:H7; Salmonella spp.; Giardia (G.) lamblia and intestinalis; E. faecalis and faecium; Norwalk virus, Adenovirus, Astrovirus; Calicivirus; Cryptosporidium parvum E. coli; Pseudomonas (P.) aeruginosa; Klebsiella spp.; coagulase-negative Staphylococci; E. faecalis and faecium; S. aureus; Proteus (P.) spp.; S. marcescens; Citrobacter (C.) spp. Human papillomavirus; Trichomonas vaginalis; E. faecalis and faecium; C. albicans; Proteus (P.) spp.; Klebsiella spp.; E. coli; group B hemolytic Streptococcus; Gonococci; Chlamydia; Herpes Simplex Virus; Mycoplasma

Bacteria appear to initiate biofilm development in response to specific environmental cues, such as nutrient availability: microorganisms undergo a transition from free-living, planktonic cells to sessile, surface-attached cells in response to a nutrient-rich medium. Biofilms continue to develop as long as fresh nutrients are provided, but when bacteria are nutrient deprived, they detach from the surface and return to a planktonic mode of growth. This starvation response is thought to allow the cells to search for a fresh source of nutrients and is driven by well-known adaptations that bacteria actuate when nutrients become scarce.

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Fig. 1 Biological cycle of bacteria, including initiation, maturation, maintenance, and dissolution of the biofilm (Artgraph by Eng. Ettore Varoni and Dr. Silvia Bovo)

Fig. 2 Mature biofilm water channels. Channels can be used for microbial network signaling and to dilute drugs (Artgraph by Eng. Ettore Varoni and Dr. Silvia Bovo)

Conversely, the biofilm is also a complex protected arrangement, self-developed by the bacteria to enable them to survive in a hostile environment more easily than when they are in planktonic form. In particular, it enables them to optimize nutrient uptake, shelters them from removal forces, and protects them from desiccation, from host defense mechanisms, and from potential toxic or harmful agents, including antimicrobial agents. Interestingly, it is easy for the biofilm to develop antibiotic resistance, and it very frequently occurs, because the microbial community can regulate the opening and closing of the water channels biochemically (Fig. 2) and can consequently control the concentration of metabolites within the structure and/or stop the entrance of drugs. Bacteria cells in the biofilm community coordinate efforts with their neighbors, to accomplish cooperative activities such as bioluminescence production, biofilm

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development, and exoenzyme secretion. Coordination occurs through a mechanism of cell-to-cell communication called quorum sensing. This mechanism gives bacteria the capacity to recognize the population density by measuring the accumulation of a specific signaling molecule secreted by members of the community. When this density reaches a certain level, accumulation of the signal in the extracellular environment is sufficient to promptly activate the biofilm response to maintain its correct balance [16]. Moreover, “quiescent cells” have been also found inside the biofilm. These cells cannot be killed by antibiotics because they are at a low metabolic stage, assuring the protection of the structure and sustaining the drug’s ineffectiveness [17–19]. Biofilm formation stages – It is crucial to clarify and understand in depth the events involved in biofilm formation on material surfaces, in order to develop effective control strategies. Adhesion is the first step in colonization and is a cornerstone for starting biofilm formation, since it allows bacteria to grow on certain surfaces and then invade host tissues. The sequence of the interaction, between floating bacteria and a surface, may be summarized as follows [20]: 1. Convective transport of fluids and active bacterial chemotaxis. 2. Van der Waals attractive forces, which operate at separation distances greater than 50 nm. 3. At distances of 10–20 nm, the interaction of van der Waals attractive forces and electrostatic repulsion produces a weak area of attraction, which maintains reversible adhesion. 4. At the same distance and even closer, adhesion between bacterial adhesin and ligands adsorbed onto the biomaterial surface from biological fluids, when the material was installed, begins to operate. After surface colonization by pioneer bacteria, co-aggregation of other bacteria to cells that are already attached can occur. Multiplication of the attached organisms produces confluent growth of microorganisms, and a biofilm starts to form. Figure 3 shows the sequential steps of supragingival biofilm formation on a root cementum surface, through scansion electronic microscope (SEM) images.

Biofilm/Substratum/Environment Interaction Materials science and tissue engineering offer a unique opportunity to investigate biofilm formation. The availability of a stable surface is a prerequisite for the bacteria cells to attach and for consequent biofilm formation, and the properties of the surface can affect the outcome and bacteria/surface interactions. Several aspects can affect biofilm formation and growth; the most important effects related to substrate and environment will be described in the next sections.

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Fig. 3 Sequential steps of bacterial colonization in the oral cavity (SEM images). Clean and sterile cementum of the dental root surface, which is suddenly covered by a salivary pellicle (Adapted from Carrassi [21]): after 2 h, the cementum surface is colonized by a few cocci; after 12 h, the surface is completely covered by cocci and short rods; after 24 h, the biofilm has developed. Many bacteria, cocci, and short and long rods can be observed on the root surface, adhering to the slime layer

The Substrate Effect: Surface Energy and Hydrophilic/Hydrophobic Properties The initial interactions between the bacterial cell wall and a surface (including those of other cell walls) are primarily influenced by interfacial electrostatic forces (repulsion or attraction) and van der Waals forces. However, many different nonspecific interactions and interfacial forces also influence cell attachment, including hydration forces, hydrophobic interactions, and steric forces [22]. Hydrophobic (low surface energy) and electrostatic (charge) interactions are the most widely investigated phenomena. In general, bacteria may be modeled as colloidal particles approaching surfaces with a Brownian motion [48]. The interaction may be described by the DerjaguinLandau-Verwey-Overbeek (DLVO) theory focused on long-range interactions between particles and substrate. This interaction includes the Lifshitz-van der Waals interaction and the interaction resulting from the overlapping of two layers of interactions. The forces are additive, and the energy of adhesion is a function of the distance between the particle and the substrate. In the case of bacteria, the DLVO

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is not fully descriptive, and short-range Lewis acid-base interaction and hydration must also be taken into account (XDLVO). The charge upon the bacteria wall is generally measured as electrophoretic mobility. It is usually electronegative, especially in the case of Gram-negative bacteria, as is that of many material surfaces. Thus, from the theoretical standpoint, bacteria do not adhere closely except to strongly electropositive surfaces. However, in practice they may show paradoxical behavior, because of the ability of the cell wall to dynamically alter its charge in response to environmental conditions, such as pH or ionic strength in the medium. In addition, fibrils, fimbriae, and flagella may expose different charges at their tips. The walls may also be penetrated by solvents, causing dynamic rearrangement of the wall polymers and consequently altering surface charge. These phenomena explain why the bacteria/substrate interaction is not fully described by the DLVO or XDLVO theories, and bacterial behavior in regard to the electrostatic properties of the substrate is not fully predictable (Table 4) [23]. In addition to electrostatic attraction, chemotaxis and possibly haptotaxis also contribute to the initial attachment [24]; this occurs in response to chemoattractants in the environment or adsorbed onto the surfaces, such as amino acids, peptides, and glucides. The interactions between bacteria and surface, as described above, are generally reversible, but they evolve rapidly toward irreversible bonds characterized by molecular-specific reactions between bacterial surface structures and the substratum. The interactions are mediated by bacterial surface polymeric structures, called adhesins, included in the capsules, fimbriae, or pili and in the slime. For instance, S. aureus binds fibronectin, while S. epidermidis has several polysaccharide adhesins that mediate the adhesion of this bacterium to various material surfaces and protein tissues. Of the adhesins, the most important are (i) capsular polysaccharide/adhesion (PS/A), (ii) a biosurfactant known as “surface-active agent” (SAA), (iii) polysaccharide intracellular adhesion (PIA), (iv) a polysaccharide composed of β-1,6-linked N-acetylglucosamines with partly deacetylated residues, and (v) peptidoglycan, an accumulation-associated protein (AAP). PS/A and SAA take part in bacteriamaterial interactions, whereas PIA and AAP are implicated in cell-cell interactions [25] (Fig. 4).

The Environment Effect Temperature, exposure time, bacterial concentration, and the presence of antibiotics or other antibacterial molecules affect bacteria adhesion and biofilm development. In addition, physical stresses, including flow, scraping, or epithelial detachment, have a great influence on biofilm formation. In general, high mechanical stresses inhibit biofilm formation and its maturation. All these phenomena are evident in the oral cavity, where environmental conditions change frequently. It is a common observation that, in subjects who

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Table 4 Representative examples of deviations from the DLVO or XDLVO theory observed in bacterial adhesion studies (Adapted from Poortinga et al. [23]) Strain Arthrobacter, Corynebacterium, Rhodococcus, Pseudomonas, Gordona

Experiment Adhesion to glass and Teflon

E. coli

Adhesion to sludge flocs

Vibrio alginolyticus

Adhesion to hydroxyapatite

Corynebacterium

Accumulation of bacteria at air-water interface

S. salivarius

Adhesion to glass

Marine strains Sphingomonas paucimobilis

Adhesion to hydrophobic and hydrophilic polystyrene Adhesion to bare glass and EPS-coated glass

Pseudomonas

Adhesion to sand

E. coli

Direct measurement of bacterial interaction force with glass, mica, and hydrophobic polymers

Pseudomonas and Burkholderia

Measurement of bacterial interaction with silicon nitride AFM tip

Findings Experimentally obtained energy barriers against adhesion are some orders of magnitude smaller than DLVO predictions at low ionic strength Adhesion does not correlate with bacterial zeta potential but with a fraction of the positive charge present on the bacterial cell surface Bacterial adhesion increases at increasing ionic strength, in accordance with the DLVO theory, but decreases when ionic strength exceeds 0.1 M In contrast to DLVO predictions, under repulsive conditions, accumulation decreases for increasing ionic strength Despite small differences in DLVO interaction energies, adhesion rates of a fibrillated and non-fibrillated strain differ greatly No correlation found between adhesion and ionic strength The XDLVO theory can explain adhesion to glass but cannot explain adhesion to glass coated with bacterial EPS A fraction of the bacteria adheres faster than the rest, while DLVO calculations predict no difference Force measurements do not correlate with XDLVO calculations for a lipopolysaccharide covered strain but do correlate for a strain with truncated lipopolysaccharide chain A repulsive force extending over longer distances (>100 nm) than predicted by the DLVO theory is measured

do not brush their teeth efficiently, plaque accumulation is abundant. In xerostomic patients, who are deficient in saliva amount and flow, plaque accumulates very

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Fig. 4 Molecular interaction between bacteria and substrate (Adapted from Katsikogianni and Missirlis [25])

rapidly, and the clinical consequences consist of a prevalence of caries and periodontal disease.

The Effect of Surface Roughness at Micro- and Nanoscales Certain physical parameters, such as surface roughness and morphology, are thought to closely affect biofilm formation. It is well known that rough restorative materials accumulate more plaque and expose patients to the risk of developing caries and gum diseases at neighboring sites. This is a key aspect in implantology, because most implants available on the market are designed to be rough and grooved, in order to improve primary stability, healing of mineralized and soft tissues, and maintenance of tissue integration around the implants over time, whether in healthy or diseased subjects. However, when rough surfaces are exposed to the oral environment, biofilm formation is swift, mainly because of the roughness and grooving shelter bacteria from physical removal, hindering cleaning procedures. Biofilm formation around implants is an etiological factor for peri-implantitis and implant failure or loss [26]. It has been observed that, although roughness and wettability are related, the roughness parameter is often predominant [22]. The clinical roughness threshold for biofilm formation in the oral cavity has been shown to be Ra = 0.2 μm: below this threshold, for Ra values within the microscale, there is no significant improvement in inhibiting bacterial adhesion [21, 27]. In contrast, at the nanoscale, rough and geometrically determined surface morphology has been shown to produce antifouling properties. At this scale, interaction of the bacteria with the surface remains limited to the surface of physical protrusions, like drops of dew on the leaves of a lotus flower (Nelumbo spp.), and bacteria are repulsed [28].

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The Effect of Protein Absorption As described earlier in this chapter, the first step in the pathogenesis of foreign bodyrelated infections is bacterial adhesion. The mechanisms involved in adhesion lead to passive adsorption of the bacterial cells on the solid material, through physicochemical surface interactions with bacterial structures termed bacterial adhesins. Thus, bacterial behavior varies as a function of material hydrophobicity and electrostatic charge. Chemo-physical properties and functional groups exhibited by the biomaterial surface interact with those of the bacterial cells, determining the kinetics of microbial adhesion. However, in many cases of implanted or invasive medical devices, materials first come into contact with body fluids. This is particularly true in the oral cavity, where installed materials are immediately wetted by the saliva, crevicular fluid, or blood, depending on the anatomic site of application. The components of body fluids, mainly proteins, are rapidly adsorbed onto the material surface. The protein film that quickly forms on the biomaterial surface during the initial exposure to physiologic fluids may thus be considered as the true interface with the bacteria. Nonspecific effects have been described, such as those derived from albumin surface adsorption, thought to alter the physicochemical characteristics of the surface and to increase the degree of hydrophobicity, while competing for the surface with other pro-adhesive host proteins. In addition, various host proteins mediate bacterial adhesion by interacting with bacterial adhesins; these are frequently receptor proteins known as “microbial surface components recognizing adhesive matrix molecules” (MSCRAMMs). The bacteria-binding host proteins include collagen, fibrinogen, fibronectin, laminin, vitronectin, clumping factors A and B, bone sialoprotein, elastin, and IgG. Charged surfaces can also interact electrostatically with other extracellular polymeric components. In addition to polysaccharides, other extracellular polymeric substances are produced by biofilmforming bacteria; these include extracellular DNA, teichoic acids, and amphiphilic molecules, whose production or proportion may depend on the specific growth phase. Effective low-adhesion surfaces are thus hydrophilic, highly hydrated, and non-charged. These types of surface appear to prevent or limit contact between a bacterium and the potential attachment points of the material surface [28]. The adsorption of proteins on a surface can be reduced, either by altering the interaction potential or by slowing down the rate of adsorption through highpotential barriers to interaction. This latter method of controlling the kinetics of adsorption can be achieved by polymer grafting, resulting in the introduction of long-range repulsive forces. Other strategies to achieve lower bacterial adhesion to biomaterials exposed to protein solutions rely on conditioning the surface by pre-adsorption of molecules claimed to increase apolar hydrophilicity and hydrophobicity or to compete with host adhesion adsorption [29]. In addition, the possibility of controlling tissue integration while contrasting bacterial adhesion, simply by acting on the topographical features of the biomaterial surface, is certainly very

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attractive. Specifically patterned surfaces can direct the alignment and spatial distribution of bacterial cells. At the same time, customized superficial nanostructures can reduce the areas of contact where eukaryotic and bacterial cells can anchor. Topographies can achieve a degree of complexity that confers entirely new properties on the material surface [30].

Biofilm Formation on Dental Implants and Prosthetic Dental Materials Biofilm Formation on Dental Implants Biofilm formation on dental implants is the crucial step toward the inflammation of peri-implant tissues, jeopardizing the long-term success of osseointegrated implants. In general, the assessment of the microbiological and immunopathological aspects of peri-implant diseases has shown a microbiological diversity of peri-implantitis biofilms and a specific local immune response of the host [8]. Bacterial colonization and adhesion at the implant surface starts already 30 min after placing the device and lasts for several months. For instance, the presence of S. aureus has been confirmed as long as 1 year later. As pointed out in the paragraph “Peri-implant microbiome,” the bacterial composition of the newly formed implant biofilm closely resembles that of the nearest teeth, suggesting that the microbial flora on dental substrates can act as a “reservoir” for the bacteria that compose the biofilm around implants. Importantly, bacteria of subgingival biofilms, collected from periimplantitis patients, displayed multiple antibiotic resistances in vitro, for example, in the case of Prevotella intermedia, Prevotella nigrescens, or Streptococcus constellatus. Although the qualitative composition of the biofilm in peri-implantitis shows similarities to that of periodontitis, supporting the hypothesis that patients with active periodontal disease are at higher risk for developing peri-implantitis, several further microorganisms, very uncommon in periodontitis, have been recognized in peri-implantitis; these include Staphylococcus aureus, Staphylococcus epidermidis, Escherichia coli, Peptostreptococcus micros, and Pseudomonas spp. [8]. A further peculiar element in peri-implant mucositis, and subsequently in peri-implantitis, is that inflammation acquires typical features defined as the “specialized innate response.” Peri-implantitis displays larger numbers of immune cells, mainly interstitial dendritic cells and related inflammatory mediators. The progression from mucositis to peri-implantitis is characterized by a drastic increase in neutrophils, osteoclasts, macrophages, and lymphocytes, in findings supported by transcriptome analyses. Compared to the inflammatory tissue from periodontitis sites, the periimplant granulation tissue displayed a specific innate response, with greater mRNA expression of pro-inflammatory cytokines, such as interleukin (IL)-1, IL-6, and IL-8. Moreover, resident primary fibroblasts showed increased production of

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vascularization factors, matrix metalloproteases, and complement receptor C1q, with decreased production of metalloprotease inhibitors and growth factors for collagen synthesis [8]. Recent studies have analyzed samples of crevicular fluid collected from the sulcus around abutments and report a significant difference between supra- and subgingival plaque: these findings supported the hypothesis that the cellular adherence of peri-implant tissue to titanium implant, via hemidesmosome, actin filaments, and microvilli, greatly reduces the risk of formation of anaerobic subgingival pockets. Indeed, the biofilm coating observable on supragingival abutment surfaces appeared significantly thicker than that on subgingival sites. Together with surface localization (supra- and subgingival) of oral biofilm, surface modification of biomaterial also appeared to significantly affect the health status of tissues around implant abutments. Two main aspects are particularly involved, i.e., the local immune response to biomaterial and the biofilm adhesion and proliferation on it. With regard to the former, particularly in the case of mucositis, the physicochemical treatment of the implant surface during manufacturing appears to affect the inflammatory response of the adjacent mucosal tissue, in terms of different microvessel density and amount of inflammatory infiltrate. Regarding biofilm adhesion and growth, the surface chemistry and the design features of the implant-abutment configuration can affect biofilm formation. As mentioned, increased surface roughness and surface free energy appear to promote dental plaque formation on implant and abutment surfaces, although this conclusion derives chiefly from descriptive literature, rather than from high-quality meta-analyses. A considerable debate still surrounds the issue, and in particular the precise role played by physicochemical and textural properties of the implant surface on microbial composition is still unknown. It is hypothesized that greater roughness and higher free energy at the implant surface might promote biofilm formation, so that peri-implantitis might occur and progress more quickly. However, and conversely, some evidence also supports the hypothesis that abutments with different surface characteristics do not greatly influence either biofilm formation on the implant surface or the extent and composition of the inflammatory response. No implant system or surface type has been found superior over any other in terms of marginal bone preservation, the main reason for this probably being related to the presence of salivary proteins at the interface between the host tissue and biomaterial. The latter adheres first at the implant surface and can mediate bacterial adhesion: any differences in bacterial adhesion due to surface microstructures may partially be “counteracted” or masked by this salivary pellicle, which mediates the mucosaimplant interconnection [8]. Taken together, the results of these studies suggest that the diversity of the microbial community and the subsequent immunity response of peri-implantitis versus periodontitis might not be as close as has been believed: further investigations targeting the multiplicity of peri-implant-specific microbiota will be needed to identify the best approach for peri-implantitis management, still an important clinical challenge.

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Biofilm Formation on Restorative and Prosthetic Materials As was said in the paragraph “Dental caries and the oral microbiome,” dental caries is chiefly the result of an imbalance in metabolic activity within the oral biofilm, which becomes skewed toward a strong acidification of the milieu at the tooth surface, leading to the dissolution of hard dental tissues (enamel and dentine). From a metabolomics perspective, the cariogenic potential of the microbial community must be described in terms of activities relevant to acid production. Recent studies have shown that metabolomics may explain caries pathogenesis better than a focus solely on microbiome composition; unsurprisingly, sound evidence exists to confirm that carbohydrate metabolism is a cornerstone in caries development, because of its capacity to acidify the environment and dissolve dental tissues, leading to tooth decay. For dental applications, antimicrobial coatings killing bacteria upon contact are more promising than antimicrobial-releasing coatings. Moreover, certain natural polymers, used as biomaterials with intrinsic antibacterial properties, such as chitosan or pectins, could be useful tools, in that they would contextually exert antimicrobial activity during tissue regeneration [31, 49]. Biofilms appear to form in different ways, depending on the different types of biomaterials used in restorative and prosthetic dentistry. On gold and amalgam, the in vivo growth of dental plaque appears thick and can almost completely coat the substrate, but it is also barely viable. Conversely, on ceramics oral biofilms are thin but highly viable. Dental plaque on composites and glass ionomer cements has been reported to produce surface decay, which appears to further enhance biofilm proliferation. In particular, residual monomers released from composites affect plaque development in vitro, but the corresponding in vivo effects are less striking, probably due to the greater dilution of these compounds, which become dissolved in a huge volume of saliva, which is continuously replaced by the flow rate. Dental plaque grows readily on the acrylic bases of dentures, mainly because of their porous structure. The composition of oral biofilms on the mucosal and prosthetic surfaces has been investigated, to determine any differences. A recent study analyzed 61 edentulous subjects with complete maxillary and mandibular dentures [32]: “supragingival” plaque samples were collected from the acrylic base; from the dorsal, lateral, and ventral surfaces of the tongue; from the floor of the mouth, the buccal mucosa, the hard palate, the vestibule/lip, and the attached gingiva; and from the saliva. The microbial profiles of plaque from the soft tissues differed with the site considered, but the main periodontal pathogens, i.e., Aggregatibacter actinomycetemcomitans and Porphyromonas gingivalis, were detectable in all specimens. In particular, samples from the dorsum of the tongue showed the highest bacterial counts, followed by the adherent gingiva and the lingual margins; the lowest counts were recorded for samples from the buccal mucosa and the labial vestibular mucosa. The patterns of microbial colonization versus harvesting site showed three clusters: the first cluster included the saliva, the supragingival plaque, and the lateral and dorsal surfaces of the tongue; the second cluster comprised the six remaining soft tissues; and the third cluster comprised all species on the denture palate.

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The Role of Chemistry in Dental Biofilm Limitation Current Strategies Numerous strategies are currently available to hinder the formation of “pathogenic” oral biofilm on dental biomaterials and the development of related dental disease. These include strategies relating to the materials themselves, substances used to dope materials, and different types of surface coating including bioactive coatings, microand nano-particles, etc. Materials with intrinsic antibacterial properties – Bulk materials that exert antibacterial action without requiring any modification are generally described as intrinsically antibacterial. Numerous metals, such as silver, zinc, and copper, are known to be intrinsically bactericidal. However, their activity is not usually highly specific and is not solely oriented against prokaryotic cells: there is generally a certain degree of cytotoxicity against host cells in peri-prosthetic tissues, reducing their viability. This is often due to the metals becoming corroded in the physiological environment or to its inexorable leaching that leads to the release of high concentrations of active ions, causing local toxicity and, in some cases, accumulation in distant target organs. Silver is certainly the most widely used for biomedical applications; its bactericidal activity is related to the inactivation of critical enzymes of the respiratory chain (e.g., succinate dehydrogenase) by binding to thiol groups and induction of hydroxyl radicals. Recently, the utilization of silver as thin nano coatings, in doped solid or hydrogel materials, in the formulation of bioactive alloys and glasses and its use in the form of micro- or nanoparticles, has progressively advanced, although the possible inactivation of silver-mediated antibacterial activity in physiological fluids and the low biocompatibility index are still debated. Gallium-based treatments provide promising titanium anti-biofilm coatings to develop new bone-implantable devices for oral, maxillofacial, and orthopedic applications [33]. Recent evidence shows that the biological functions of Fe3+ are impaired by replacing iron with gallium; gallium inhibits Fe3+ biological functions by what is known as a “Trojan horse” strategy [33]. Chitosan is another substance known to possess intrinsic antibacterial and antifungal activities [34]. However, chitosan is a polycationic polymer derived from chitin, and it only has bland bactericidal activity, usually enhanced at low pH. Bioactive coatings with bactericidal agents – Bioactive antibacterial coatings have been developed with the purpose of achieving desirable new anti-infective properties at the biomaterial-tissue interface, without compromising the characteristics of the bulk material. In the so-called contact biocides, anti-infective surfaces involve the use of non-leachable substances, such as some antimicrobial peptides, quaternary amines, and N-halamines. These bioactive surfaces only kill bacteria on contact, as the bactericidal substances are not released, and are activated following direct interaction with the bacterial cells. Direct contact-killing is based on extremely high electrostatic forces on the surface that can disrupt bacterial cell membranes by removing anionic lipids [35]. The limit of this strategy is that surfaces can potentially

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be masked and inactivated when filmed by the host proteins present in protein-rich physiologic fluids. Nitrogen monoxide (NO), a natural molecule with pleiotropic functions, usually produced by leukocytes as host defense against microbial pathogens, plays an important role as a bioactive bactericide [36]. However, NO can interact with superoxide in the tissues, in conditions of oxidative stress, generating the highly cytotoxic peroxynitrite (ONOO); this makes it very important to fine-tune the beneficial and toxic effects of NO, by carefully controlling the release kinetics. Great interest is directed toward substrates that become antimicrobial following a process of photoactivation; these include titanium oxide (TiO2). TiO2 surfaces undergo photoactivation upon irradiation, with an adsorption wavelength of 385 nm; this irradiation excites the anatase allomorph, which is one of the three main TiO2 polymorphs. The bactericidal action of irradiated titanium surfaces is due to reactions of photooxidation, which involve O2 and H2O, with the formation of hydroxyl radicals (HOO-) and the direct and indirect oxidation of organic substances. These radicals are highly effective at disrupting bacterial membranes. In particular, AgeTiO2 appears to be a very promising coating, combining the known oligodynamic bactericidal properties of silver ions with an enhanced photocatalytic activity, conferred by facilitating electron-hole separation and/or increasing the surface area for adsorption [24]. Materials delivering antibiotics – An obvious step to produce biomaterials with anti-infective properties is to incorporate antibiotics within the biomaterials. Antibiotics can be incorporated variously into the bulk or coating of a biomaterial, and the incorporation can be either in molecular or in particle form. The release can consequently occur by different modalities, including diffusion to the aqueous phase, erosion/degradation of resorbable loaded matrices, and hydrolysis of covalent bonds. Thus, delivery kinetics depends on the stability of the molecular bonds or on the rate of biodegradation/bioerosion of the matrices entrapping the antimicrobial agent. However, these delivery mechanisms have been widely debated, especially regarding their efficacy over the long term (>3 weeks) [37]. Urinary and central venous catheters provide a significant example of the use of materials delivering antibiotics: a study comparing different types of antibiotic- and metal/antibiotic-doped urinary catheters found no difference in bacteria reduction at 3 weeks between doped and non-doped catheters; however, during the first week, the bactericidal efficacy of the doped catheters was clearly superior to that of their non-doped counterparts. A study examining the bactericidal efficacy of central venous catheters found efficacy to be closely related to the implant site: the infection rate was reduced in the femoral and jugular veins but remained unchanged in the subclavian vein [50]. There is general concern that the routine use of antibiotic-loaded biomaterials will increase the spread of antibiotic resistance: after an initial burst, antibiotic release diminishes and becomes subinhibitory. A number of studies have reported that subinhibitory concentrations of certain antibiotics enhance, rather than inhibit, biofilm formation by bacteria. This leads to the need for new therapeutic agents.

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Antimicrobial peptides (AMPs) are a very interesting emerging class of molecules that occur naturally in the mechanisms of innate immune defenses in multicellular organisms. AMPs show broad-spectrum activity against a large class of pathogens, and their microbicidal action is related to their ability to determine transmembrane pores. Thus, AMPs are considered to be a very promising class of bactericidal agents, and they have been studied in depth and tested in several clinical trials in order to clarify their biocompatibility. Nanostructured anti-adhesion surfaces – Certain nanostructural features of material surfaces have been shown capable of altering the 3D conformation of adsorbed proteins, and this might have an effect on host adhesins that film the biomaterial surfaces [38]. In this connection, one of the most rapidly expanding strategies in the field of nanotechnologies is the exploitation of the antibacterial properties of nanoparticles (NPs). The bactericidal activity clearly depends on the NPs’ characteristics in terms of material, charge, and size. In the case of gold NPs, the bactericidal action has been found to be determined by inhibition of ATP synthase activity associated to the change in membrane potential and by inhibition of the subunit of ribosome for tRNA binding. Silver NPs (AgNPs) appear to interact with the bacterial cell wall, disturbing its permeability, inactivating essential proteins such as thiolcontaining enzymes, causing DNA condensation, and leading to Reactive Oxygen Species (ROS) generation [39]. However, together with these positive bactericidal effects, it must be stressed that NPs can sometimes have toxic effects: the induction of apoptosis and genotoxic effects related to NPs’ translocation to distant tissues/ organs have been reported [40]. Thus, the chemical composition, size, shape, concentration, rate of dissolution/degradation, and surface properties of nanoparticles must be clearly understood and fine-tuned to achieve the best performance in terms of the benefits/drawbacks ratio. Anti-biofilm bioactive molecules – Recent progress in understanding the molecular mechanisms implicated in the physiology of biofilm formation has opened new vistas concerning how to contrast the colonization of bacteria on biomaterial surfaces [41, 42]. This has led to the development of numerous different active substances, including molecules with different mechanisms of action: enzymes capable of selectively degrading extracellular polymeric substances of the biofilm (e.g., dispersin B, rhDNase I), bactericidal molecules capable of killing metabolically quiescent bacterial cells (e.g., lysostaphin, certain AMPs), molecules and other microorganisms interfering with the quorum sensing system and inducing biofilm dispersion (e.g., furanones), and molecules downregulating the expression of biofilm extracellular polymeric substances (e.g., N-acetylcysteine) [43]. All these molecules have a serious defect in common: their efficacy is limited to a single species or at best to a small number of species; this greatly restricts their effectiveness against bacterial communities. Exceptions are the proteolytic enzymes, such as trypsin and proteinase K, which can degrade even host extracellular matrix proteins, and whose internal use in an in vivo physiological environment could obviously have adverse effects on the wound healing process. The most promising therapies now being studied comprise combinations of anti-biofilm molecules and conventional wide-spectrum antibiotics, as, for example, was shown in a study [44] combining

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dispersin B and cefamandole for the treatment of staphylococcal biofilm growth on polyurethanes.

Further Strategies With regard to pathogenesis, the combination of the different “–omics” and related innovative technologies will provide an increasingly comprehensive view of the role that the oral microbiota can play in health and biomaterial-related dental diseases, from peri-implantitis to prosthetic candidiasis. Among others, metabolome analysis is probably the most promising method to monitor these dynamic metabolic activities, helping to clarify pathogenesis. Nonetheless, it may also be applied in examining the effectiveness of both conventional drug therapies and novel compounds and might even provide useful insights for the identification of pioneering biomarkers relevant for the development and progression of biomaterial-related diseases. While ongoing preclinical and clinical studies hope to accumulate more data on the disease pathogenesis, as well as on the efficacy of current anti-infective strategies, new possibilities to counteract biomaterial-associated infections are advancing. Pre-inoculating urinary catheters with nonpathogenic E. coli were found to significantly impede catheter colonization by E. faecalis. However, some practical difficulties surround the introduction of this approach into clinical trials, as it would entail applying non-sterile catheters. The use of phages as “biological weapons” has been attempted, with controversial results: whereas a high inhibition ratio >4 log has been shown in in vitro experiments, no significant result emerged from in vivo studies [45]. Moreover, the use of phages as therapeutic agents is severely limited by (i) their high specificity, (ii) bacterial resistance, (iii) pre-inactivation by the immune system, (iv) poor resistance in the surface immobilization step, and (v) high risk of unpredicted virus expansion using phages as vector. A possible future approach to combating biomaterial-associated infections, while avoiding the use of today’s antibiotics, might be provided by antisense peptide nucleic acids (PNAs). These can interfere with the expression of critical bacterial genes that are involved in antibiotic resistance, biofilm formation, and bacterial reproduction/survival. Gram-positive bacteria are less susceptible to cell-penetrating peptides conjugated with PNAs; however, studies have shown positive results on Gram-negative bacteria by targeting the rpoD gene, which encodes an RNA polymerase primary σ (70) that is essential for bacterial growth [46]. However, a number of critical concerns surrounding the safety of PNAs must be addressed before this technology will be able to enter clinical trials on human patients; in particular, these concern possible mutagenic effects deriving from the complexation of PNAs and their degradation products, which might match DNA and knockdown, or even knockoff, sequences of the human genome. An alternative strategy has been presented, which is based on contrasting bacterial infections by modulating the host’s local immune response, rather than by

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counteracting bacterial colonization directly [47]. Two active cytokines, namely, monocyte chemoattractant protein-1 (MPC-1) and interleukin 12 p70 (IL-12), were tested. The former is a powerful macrophage-recruiting cytokine, while the latter, IL-12, can induce T-helper cells to secrete Th1 cytokines, such as interferon-g (IFN-g), which in turn stimulate the bactericidal activity of macrophages. The results are promising, but no synergic activity between the cytokines was observed. Finally, autologous platelet-rich plasma (PRP) was also found to be bactericidal when used as surface coating: in vitro experiments have shown that PRP can cause a reduction in colony-forming units of two logs. The increasing use, in dentistry as well as in other medical fields, of implantable devices and the apparently unstoppable advance of drug-resistant bacteria are combining to make it imperative that we understand and combat the development of bacterial biofilm on non-biological surfaces. Several interesting approaches are being developed, in the hope that further research will lead to eradicating infectionassociated implant failures.

Summary • The human body contains complex microbial communities with essential functions for the host’s health. • The oral cavity is an example of a dynamic microbial niche. • The increasing use of implantable devices has led to the emergence of biofilmrelated device infections on the part of apparently unstoppable multidrug-resistant bacteria. • New prevention strategies are being developed in order to reduce the frequency of infection-related implant failures. • Emerging approaches are still a matter of debate.

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47. Li B, Jiang B, Boyce BM, Lindsey BA (2009) Multilayer polypeptide nanoscale coatings incorporating IL-12 for the prevention of biomedical device-associated infections. Biomaterials 30:2552–2558. doi:10.1016/j.biomaterials.2009.01.042 48. Israelachvili JN (2011) Intermolecular and surface forces, rev 3rd edn. Academic Press, Elsevier, Waltham, Massachusetts, United States 49. Cochis A, Fracchia L, Martinotti MG, Rimondini L (2012) Biosurfactants prevent in vitro Candida albicans biofilm formation on resins and silicon materials for prosthetic devices. Oral Surg Oral Med Oral Pathol Oral Radiol 113:755–761. doi:10.1016/j.oooo.2011.11.004 50. Jahn P, Beutner K, Langer G (2012) Types of indwelling urinary catheters for long-term bladder drainage in adults. Cochrane Database Syst Rev 10:CD004997. doi:10.1002/14651858. CD004997.pub3

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Mishel Weshler and Iulian Vasile Antoniac

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biomaterials Available for Dental Bone Graft Substitutes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Autogenous Bone Graft . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Osteoinductive Agents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Osteoconductive Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Collagen . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Future Directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Principle of Guided Bone Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biological Principles of Guided Bone Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biological Factors Influencing the Reconstruction of the Alveolar Bone . . . . . . . . . . . . . . . . . Tissue Integration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Membrane Design Criteria and Material Selection . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biocompatibility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Non-resorbable Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biodegradable Barrier Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Resorption Patterns of the Alveolar Ridge . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Surgical Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Post-extraction Sites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Immediate Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Delayed Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Horizontal Defects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Dehiscences and Fenestrations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Autogenous Intraoral or Extraoral Blocks . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Vertical Defects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Vertical GBR with Membrane . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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M. Weshler (*) Laniado Hospital, Netanya, Israel e-mail: [email protected] I.V. Antoniac University Politehnica of Bucharest, Bucharest, Romania e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_52

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Sinus Elevation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Preoperative Antibiotics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Flap Design . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Site Preparation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Graft Material Positioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Membrane Selection and Positioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biodegradable Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Non-resorbable Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Suturing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Follow-Up . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Temporary Dentures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Membrane Removal . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Guided bone regeneration (GBR) membranes were originally developed to promote new tissue growth within a protected volumetric defect for periodontal regeneration. The desire to promote new bone growth without resorting to grafting procedures led to the widespread use of this technique in implant surgery. The main aim is to allow ingress of bone cells to promote bone formation within the defect. Over the last two decades, the development of the technique of guided bone regeneration (GBR) has had a significant impact on esthetic reconstruction in conjunction with implant therapy. This technique involves the use of physical barrier membranes during the healing phase in order to avoid ingrowths of undesired tissue types into a wound area. Different practical aspects related to the use of bone graft and guided bone regeneration for dental implants will be revealed. Keywords

Guided bone regeneration • Dental bone graft • Biomaterial • Bioceramics • Collagen • Membrane • Design • Surgical technique

Introduction Bone grafts are necessary to provide support, fill voids, and enhance biologic repair of skeletal defects. They are used by orthopedic surgeons, neurosurgeons, craniofacial surgeons, and periodontists. Bone harvested from donor sites is the gold standard for this procedure. It is well documented that there are limitations and complications from the use of autograft, including the limited quantity and associated chronic donor site pain. Despite the increase in the number of procedures that require bone grafts, there has not been a single ideal bone graft substitute. Scientists, surgeons, and medical companies, thus, have a tremendous responsibility to develop

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biologic alternatives that will enhance the functional capabilities of the bone graft substitute and potentially reduce or eliminate the need for autograft. The overriding requirement for successful implant placement is to have enough bone volume of sufficient density to enable an implant of the appropriate size and requirements to be placed in a desirable position and orientation. Many of the grafting materials and procedures described in this chapter have been developed as localized procedures to overcome small anatomical limitations. There is also a need occasionally to employ more complex techniques to change the entire alveolar ridge form that may additionally involve an associated change in the skeletal base. Aside from the obvious osseous component to this problem, there are also many situations where the soft tissue in the area of the proposed implant placement is deficient. The soft tissues play a vital role in maintaining the peri-implant environment and long-term health and also contribute greatly to the resulting esthetics, particularly in the anterior region. The peri-implant soft tissues must be able to maintain their structural integrity during normal function and oral hygiene procedures. Bone grafts therefore may be employed to: • • • •

Enable a better implant placement or implant placement at all Restore any jaw defects caused by large or small infection Enhance esthetics and improve soft tissues Change the preexisting jaw relationship

We must remember that the purpose for all of this is dental restoration, thus the initial planning stages take on even greater levels of importance in potential graft cases. By their very nature, they are more difficult to plan and execute, and the end result may fall short of both the clinician’s and patient’s expectations. It is important that all the alternatives are considered and presented to the patient so that they can make an informed decision with regard to their treatment. In particular, it is important to consider whether a compromise solution using prosthetic techniques may be more desirable and achievable as well as more predictable in the long term. Another alternative is to consider whether the utilization of the various implant designs may overcome the problem. The bone is a highly dynamic tissue comprising a mineralized extracellular matrix embedded with bone cells, blood vessels, and nerves. Bone contains three main bone-specific cell types: the osteocyte is a mature cell that sits in the bone lacunae, communicates with other osteocytes through long cellular processes, senses mechanical stress in the bone, and sends signals for bone remodeling as a result of mechanical stress. The responding cells are osteoblasts, cells specialized to secrete the unique collagen-rich extracellular matrix in the bone that enables mineralization, and osteoclasts, macrophage-like cells that degrade the bone structure through a combination of localized acidification (removes the minerals) and protease secretion (breaks down matrix). Osteoclasts tunnel through the bone and are usually followed close behind by osteoblasts. The bone is in a constant state of remodeling in healthy individuals.

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The bone is formed developmentally, and during wound healing, by either endochondral ossification or by intramembranous ossification. In the endochondral ossification process, mesenchymal progenitor cells first form the cartilage. The chondrocytes then hypertrophy and the extracellular matrix mineralizes. Blood vessels invade the site, bringing cells that break down the existing matrix. Osteoprogenitor cells go on to form the bone. Long bones are formed by this process during normal development. Intramembranous bone formation is a more direct process, in which osteoprogenitor cells form the bone directly. Cranial bones are formed by this process during development. Wound healing in the bone may proceed by either process, depending on local environmental factors that include how much the ends of the bone can move relative to each other, with motion favoring the endochondral process. Osteoprogenitor cells are cells that have the potential to become bone cells and reside in the periosteum and the marrow. Osteoprogenitor cells are derived from connective tissue progenitor cells that reside also in the surrounding tissue (muscle) [1]. “Osteogenic cells”: – details of osteogenic (bone-forming) cells. A summary is as follows: Stem cell > osteoprogenitor cell > preosteoblast > osteoblast resting > proliferation > matrix deposition > mineralization In some applications (e.g., dental reconstruction), there are enough progenitor cells in the local area that stimulation of these cells will induce local bone formation. In a compromised site, such as where a tumor was removed and the local tissue irradiated, there may not be many local progenitor cells, and further, it may not be a good idea to release growth factors in a site where a tumor was removed, so alternative approaches must be considered. In addition to the local conditions at the wound site, the patient’s age and lifestyle habits (such as smoking) may influence the wound healing [2]. The use of barrier membranes for the regeneration of bone defects has significantly changed implant dentistry in the past 20 years. This principle, often called guided bone regeneration (GBR or GBR technique), was first described in 1959 by Hurley and colleagues for experimental spinal fusion treatment. In the 1960s, the research teams of Bassett and Boyne tested microporous cellulose acetate laboratory filters (Millipore) for the healing of cortical defects in long bones and for osseous facial reconstruction, respectively. The authors used these filters to establish a suitable environment for osteogenesis by excluding fibrous connective tissue cells from bone defects. However, these pioneering studies did not immediately lead to a broad clinical application of barrier membranes in patients. The clinical potential of the membrane technique was not recognized until the early 1980s, when the research team of Karring and Nyman systematically examined barrier membranes in various experimental and clinical studies for periodontal regeneration. A few years later, barrier membrane techniques were tested in experimental studies on bone regeneration. Based on promising results in these studies, clinical testing of membranes began in implant patients in the late 1980s. Since that time, the GBR technique has

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continued to evolve, necessitating an updated analysis of its scientific basis and clinical applications. The biological principles and indications for the use of guided bone regeneration (GBR) are described with specific focus on the esthetic zones in the upper and lower jaws. The efficacy of barrier membranes in conjunction with bone healing and reconstructive therapy is the result of mechanical, cellular, and molecular mechanisms. Basic studies on guided bone regeneration have shown the same sequence of healing occurring as in regular fracture repair. Different kinds of biological membranes, resorbable as well as non-resorbable, are described in the chapter. Clinical results are presented and a section on complications is also included. The author concludes that the underlying bone structure of the alveolar process plays a key role in the overall esthetic appearance.

Biomaterials Available for Dental Bone Graft Substitutes The term biomaterial generally indicates any substance used to create a medical device destined for diagnosis, prevention, control, mitigation, or therapy of a human disease, on condition that it persists in the body for at least 30 days after implantation. First of all, cytotoxicity, genotoxicity, and hemocompatibility of a biomaterial have to be evaluated. After that, attention has to be paid to its macrostructure and microstructure, by evaluating the isotropy. Finally, its mechanical, physical, and chemical properties should be taken into consideration. Which characteristics should a biomaterial have to be considered for implantation in the human body? They can be summarized as follows: • • • • • • • • • • •

Noncarcinogenic Nonantigenic Hydrophilic Radiopaque Versatile (usable in several clinical fields) Sterile Osteoconductive or osteoinductive Favorable clinical handling Resorption and replacement by host bone Available in sufficient quantities Low in cost

Biomaterials can be obtained from the patient (autogenous), from beings belonging to the same species (homologous), from beings belonging to different species (xenogeneic), or from minerals (alloplasts). Apart from autogenous bone, which has osteoconductive, osteoinductive, and osteoproliferative properties, and homologous bone, whose properties are mainly osteoconductive and slightly inductive, all the biomaterials used for bone regeneration are only osteoconductive scaffolds. Bone substitutes were created in order to promote bone regeneration, avoiding the

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necessity of harvesting the bone from the patient. The first materials on the market were represented by ceramic hydroxyapatite of different macrostructures (coral, bioglass, ceramic hydroxyapatite), and the osteoconductive potential and the resorbability were not excellent with regard to the implant field. A few years later, demineralized freeze-dried bone allograft (DFDBA) from human donors was introduced in the USA; osteoinductive properties were claimed for this, because the demineralization process was able to expose bone morphogenetic proteins (BMP). In addition, some publications confirmed an osteoconductive property for this material [3–5]. Unfortunately, the properties of DFDBA were not confirmed by later histologic and clinical studies in sinus elevation and guided bone regeneration (GBR) procedures. At the same time, xenogeneic anorganic bone was obtained from the cattle [6–8], followed by similar materials from equine or porcine sources. These mineral scaffolds, resulting from a treatment to eliminate any trace of organic material, promote colonization of the bone tissue via osteoconduction. They are slowly replaced by newly formed bone; both the quality and the quantity of lamellar bone are well documented, and at the moment, they are considered a first-choice material in bony defect repair in implantology, with the exception of classes V and VI atrophies (Cawood and Howell1), where the use of autogenous bone, alone or in association with xenogeneic materials, is mandatory to rebuild the bony architecture prior to implant placement. The degree of bone grafting required for implant placement varies from localized deficiencies to cases where there is a need to change the entire arch form and/or jaw relationship. There exist, therefore, a great many techniques and materials to facilitate such grafting procedures, many of which may be used in combination. The interaction between the graft and the surrounding host bone is very important and is the subject of much research. Although some grafts will act merely as space fillers, the ideal graft will be osseoconductive and osseoinductive. Osseoconduction is the property of promoting bone growth from the surrounding host bone onto the surface of the graft material, using the graft as a framework. The graft material in such cases may be resorbed or remain virtually intact, depending on the material used. Osseoinduction is the ability to promote reformation of remote bone from the host bone even within noncalcified tissues. Bone morphogenetic proteins and other bone-promoting factors have this latter property.

Autogenous Bone Graft Autogenous bone graft is considered the very best jaw bone transplant, since it is originally from the same person to be implanted to. There is complete identity between the object to which it should connect, and thus the chances of success of this very type of transplant are the best. The ready availability of autogenous bone has always meant that it is the first choice of bone grafting material for many clinicians. However, patient’s acceptance of autogenous bone harvesting may be low, given the potential morbidity associated with such techniques. Although a great

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Fig. 1 Bone chips from the drilling site

amount of research and clinical time has been spent over many years to develop substitutes for autogenous bone, it remains the gold standard by which all other materials are judged and is the material of choice for the present authors. Its main advantages are: • • • • • •

Availability Sterility Biocompatibility Osseoinductive potential Osseoconductive potential Ease of use

The graft acts as a scaffold for the ingrowth of blood vessels and as a source of osteoprogenitor cells and bone-inducing molecules. The graft is eventually resorbed as part of the normal turnover of the bone. Principal autogenous bone graft sources are ascending mandibular ramus, symphysis mentis, occipital bone, tibia, iliac crest, and rib bone. However there is one more autogenous bone graft source, very handy and simple to obtain but only in small portions, mainly on account of the fact that it is obtained from the drilling site chosen for the implant placement (Fig. 1). Autogenous bone graft has a documented 2–4-month period from integration to regeneration, but with a low capacity of preserving its initial volume at time of integration. Although the gold standard for bone grafting remains the patient’s own bone, the limitations on the amounts available (particularly from sites other than the iliac crest) mean that there remains a great demand for alternative graft materials. Xenografts is a graft of tissue taken from a donor of one species and grafted into a recipient of another species; allogeneic grafts (allografts) are taken from different individuals of the same species; herbal-based bone filler is an alternative biomaterial, with excellent bone regeneration potential generated from soybean; porous titanium granules are a regenerative material treating a local osseous defect around titanium dental implant; and alloplasts are synthetic materials. The macrostructure

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and microstructure of these grafts have an enormous influence on their efficacy. The pore diameter and volume are therefore of great importance. Bone grafts and their substitutes can be divided according to their properties of osteoinduction, osteoconduction, and osteogenesis.

Osteoinductive Agents Osteoinductive agents are bone graft substitutes, generally proteins, which induce differentiation of undifferentiated stem cells to osteogenic cells or induce stem cells to proliferate. Several osteoinductive agents have been identified. Among these compounds are transforming growth factor (TGF) [9] bone morphogenetic proteins (BMPs), [9–12] fibroblast growth factors (FGFs) [13, 14], insulin-like growth factors (IGFs) [15], and platelet-derived growth factors (PDGFs) [16].

Demineralized Bone Matrix Since the initial studies performed by Urist [3], the osteoinductive capacity of DBM has been well established [17]. DBM is produced by the acid extraction of human cortical bone, and the components of the bone that remain behind include the non-collagenous proteins; bone osteoinductive growth factors, the most significant of which is BMP; and type I collagen. DBM provides no structural strength, and its primary use is in a structurally stable environment. Hydroxyapatite, autograft, allograft, or bone marrow cells may be added to DBM. A carrier may be added to DBM to improve its handling characteristics and mechanical properties. DBM obtained from allogeneic human cortical bone shows variable efficacy and osteoinductive index. A reproducible and rapid bioassay, using human cells of osteoblastic lineage, SAOS-2 cells, has been developed to correlate the activity of DBM [18]. Relevant images obtained by scanning electron microscopy of relevant products are shown in Fig. 2. Bone Morphogenetic Proteins The BMPs (BMP 1-7) are low-molecular-weight non-collagenous glycoproteins that belong to an expanding TGF superfamily of atleast 15 growth and differentiation factors. They make up only 0.1 % by weight of all bone proteins. Unlike DBM, which is a mixture of BMPs and immunogenic, noninductive proteins, the pure form of BMPs is non-immunogenic and non-species specific. Currently single BMPs are available through recombinant gene technology, and mixtures of BMPs are available as purified bone extracts for clinical studies. The recombinant human BMPs extensively studied are rh-OP-1 (osteogenic protein 1), rh-BMP-2 (Genetics Institute, Cambridge, MA), and rh-BMP-7 (Creative Biomolecules, Hopkinton, MA). In October 2001, approval was granted by the FDA for recombinant OP-1 implant, for use as an alternative to autograft in recalcitrant long bone nonunions. This is the first BMP approved for clinical use in the USA. The approved product, OP-1 implant, is a combination of rh-OP-1 and a bovine collagen carrier.

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Fig. 2 Relevant images obtained by scanning electron microscopy of relevant demineralized bone matrix: (a) Puros cortical; (b) Puros cancellous; (c) Puros cortical-cancellous mix

The rh-OP-1 is derived from a recombinant Chinese hamster ovary cell line, and the bovine collagen is derived from the diaphyseal bone and is primarily type I. It is a white lyophilized powder, which has to be reconstituted with two to three ml of saline, prior to use. It forms a paste which is then surgically implanted at the fracture site. The osteogenic activity of OP-1 has been proven in a validated critically sized fibular defect in human subjects [19]. In November 2001, the first two-year results of a clinical study of rh-BMP-2 were presented at the North American Spine Society meeting.

Other Growth Factors Besides the growth factors expressed from the extracellular matrix of the bone (DBM, BMP), there are other factors in the circulating blood, which play a role in bone healing. TGF is the most extensively studied growth factor in the field of bone biology. It comprises an entire family of molecules that includes the BMPs. In 1994, Genentech, Inc. (San Francisco, CA) was issued the patent for developing TGF through recombinant technology. This covered the nucleic acids, vectors, and host cells used for production of recombinant TGF. In an animal study, it was found that BMP, and not TGF, enhanced bone formation [20]. PDGF is another factor whose

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effect was studied on the bone healing of unilateral tibial osteotomies in rabbits. It was concluded that PDGF had a stimulatory effect on fracture healing [16]. Autologous growth factors (AGF) is an innovative concept. AGF gel is obtained from the buffy coat of the blood collected in the cell saver during surgery, through the process of centrifugation. It is rich in growth factors, especially TGF and PDGF. Approximately 20 ml of AGF is derived from 500 ml of blood in ten minutes and it is placed at the operating site. Bovine-derived bone morphogenetic protein extract is a cocktail of growth factors and is currently being evaluated for its role in human spine fusion and periodontal repair. It can be combined with either DBM or a coralline calcium carbonate carrier. Basic fibroblast growth factor (bFGF) is produced locally in the bone during the initial phase of fracture healing and is known to stimulate the cartilage and boneforming cells [13]. A combined product is a formulation of bFGF and hyaluronic acid (Hy). It is delivered as a single minimally invasive injection into the fracture site. Hy is a viscoelastic polymer found throughout the body that cushions and protects soft tissues. The synergistic combination of bFGF and Hy appears to accelerate the operating site healing process and underscores the importance of using an appropriate carrier not only for bFGF but also possibly for other growth factors.

Allogeneic Bone Graft The dry weight of the bone comprises 70 % inorganic materials and 30 % organics; 90 % of the organic material is type I collagen, while the other 10 % comprises proteins that induce mineralization and signal for regeneration. It was demonstrated in the 1960s that demineralized bone (bone exposed to acid to dissolve the inorganic component, leaving the organic matrix = demineralized bone matrix) could induce ectopic bone formation via the endochondral process (ectopic = any site that is not the normal physiological site). It was hypothesized that a diffusible factor was present in demineralized bone matrix. Characterization of the properties of demineralized bone matrix led to identification and cloning of a molecule (now called BMP-2) that could induce ectopic bone formation on its own. The BMP family has grown substantially and it is now recognized that these molecules also play important and essential roles in development. BMPs induce cell migration, proliferation, and differentiation, and it is not entirely clear yet which of these processes dominate in vivo. In order for BMPs to be effective, progenitor cells that can be induced to form bone must be present in the local area. This means that BMPs will not likely be effective in very large defects or defects which have been compromised by irradiation or large infection. Allogeneic bone graft has a documented 3–4-month period from integration to regeneration, but with a low capacity of preserving its initial volume at time of integration (Figs. 3 and 4). Xenogeneic Bone Graft Xenografts derived from natural bone sources have been extensively investigated in multiple experimental and clinical studies. In particular, cancellous bovine bone

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Fig. 3 Allogeneic bone chips of 0.5–1.00 mm mixed with sterile sodium chloride 0.9 % solution

Fig. 4 SEM image of the allogeneic bone chips of 500–1000 μm (5000)

has been used as a source for these bone substitute materials because of its close similarity to cancellous human bone. The organic component is removed by heat treatment, by a chemical extraction method, or by a combination of the two to eliminate the risk of immunologic reactions and disease transmission. Since the first reports of bovine spongiform encephalopathy, there has been a particular focus on the ability of these extraction methods to completely eliminate all protein from the bovine bone source. However, despite the hypothetical risk of organic remnants in bovine bone substitutes, there have been no reports of disease transmission from these materials. In contrast, a few cases of transmission of human immunodeficiency virus and hepatitis related to allogeneic materials have been reported. Deproteinized bovine bone minerals (DBBMs) are in general known to be biocompatible and osteoconductive, although the production methods have a strong impact on their biologic behavior. Two bovine bone substitutes derived from bovine cancellous bone, one deproteinized by high temperatures and the other mainly by chemical means.

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This is an osteoconductive and slow-resorbing material composed of an anorganic mineral matrix deprived of the organic scaffold in order to leave intercrystalline microtunnels and microcapillaries between the bovine apatite crystals. The high osteoconductivity is due to the natural microstrucure of the material, which demonstrates a large inner surface area and a system of intracrystalline spaces and microtunnels available for ingrowth of blood vessels and osteoblast migration. Long-term stability has been proved by many clinical studies. There must be no direct contact of the material with the implant surface for good implant osseointegration. The only early contact is between the clot and the matrix particles, and the angiogenesis and osteoblasts deposition occur from there. Integration is due to replacement of the bone substitute with newly formed bone. Histomorphometric analysis demonstrated that anorganic bovine bone increases the mineral portion in regenerated areas as compared to host bone areas [9–12]. Some of the material remains in the bone tissue and is slowly embedded in lamellar bone, resulting in denser bone, and this could explain the high survival rate of implants placed in areas augmented with it. Xenogeneic bone graft has a documented of 6–9 month period from integration to regeneration, but has a high capacity of preserving its initial volume at time of integration (Figs. 5 and 6).

Herbal Based Bone Filler Soybean is a natural material made of protein and carbohydrate fractions (approximately 40 % by weight for each fraction), of an oil fraction (approximately 18 %), and of minerals (approximately 2 %). Soybean also contains isoflavones, phytoestrogens with an ascertained action on eukaryotic cells. Isoflavones inhibit tumor cell proliferation and immunocompetent cell activation and seem to reduce scar formation in wound healing. Recently, a new class of biomaterials has been developed from defatted soybean curd and flour. The processing of these components by either thermo-setting or extraction allows the preparation of materials with different physico-chemical properties; by these processes membranes, films, granules and gels can be obtained. The bone regeneration potential of these biomaterials has been demonstrated by in vitro studies highlighting their inhibitory effect on monocytes/macro-phages and osteoclasts as well as their ability to induce osteoblast differentiation and bone nodule mineralization [21]. Soybean-based biomaterials clearly promote bone repair through a mechanism of action that is likely to involve both the scaffolding role of the biomaterial for osteoblasts and the induction of cell differentiation. Therefore, these biomaterials have a potential to become fillers alternative to osteoconductive products such as those based on either autologous bone or ceramics or PLA/PGA hydrogels. Indeed, in addition to their bone repair potential, the ductility of Soybean bone filler biomaterials brings advantages in the surgical practice when compared to the brittle and not malleable ceramics or to the relatively loose consistency of hydrogels. Soybean-based bone filler has a documented 4–6-month period from integration to regeneration, but has a high capacity of preserving its initial volume at the time of integration (Fig. 7).

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Fig. 5 Xenogeneic bone chips mixed with sterile sodium chloride 0.9 % solution

Fig. 6 SEM images of the xenogeneic bone chips at different magnification

Porous Titanium Granules The search for osteoinductive as well as osteoconductive materials has led to the novel idea of using titanium in bone augmentations of the alveolar crest. The indications for the use of porous titanium granules material in this area have so far been limited to recently reported sinus augmentations and for defects around dental implants, even though the idea was first tried in a pilot case in 1995 [22, 23]. The material has also been evaluated in a clinical study of sinus augmentations where the main part of the included study subjects (12 patients) had implants simultaneously installed and porous titanium granules material placed around them in the sinus floor in a one-stage procedure. Four patients had a delayed placement of implants due to insufficient primary stability at the time of augmentation. Three implant losses after an observation period of 12–36 months were seen and two of these were in the staged group. In the simultaneous placement group, one implant was lost after one year of loading (after a history of postoperative sinus

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Fig. 7 Soybean chips of 0.1–0.2 mm mixed with sterile sodium chloride 0.9 % solution

infection). The authors raised questions regarding the usage of the material in staged sinus lifts as well as further explore the risk of displacement of granules into the sinus during augmentation [23]. Porous titanium granules are not being replaced, but have a high capacity of preserving its initial volume at the time of integration (Figs. 8 and 9).

Osteoconductive Materials Osteoconduction is a three-dimensional process that is observed when porous structures are implanted into or adjacent to the bone. Porosity alone, however, is not adequate for bone ingrowth. Porosity with interconnectivity is the most essential prerequisite. This is based on the three-dimensional interconnections between the lacunae in the bone that provide intercellular communication. Although there are alternative views, the consensus of research indicates that the requisite pore size for bone ingrowth into porous implants is 100–500 μm, and the interconnections must be larger than 100 μm [24]. Synthetically produced bioceramics have the advantage of having no risk of cross infection but may still give rise to an antigenic response. Their physical properties can be manipulated to a great degree, and they may be used also in combination with bone-promoting molecules to enhance their effectiveness. They act as a framework for bone formation on their surface and are therefore osseoconductive. They include: • • • •

Calcium sulfate Calcium phosphate Bioactive glasses Calcium carbonate

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Fig. 8 Porous titanium granules of 0.5–1.00 mm mixed with sterile sodium chloride 0.9 % solution

Fig. 9 SEM images of porous titanium granules of 500–1000 μm (20,000)

Calcium Sulfate Gypsum, also referred to as plaster of Paris, owes its name to a village just north of Paris. Although its external use for creation of hard setting bandages dates back to the seventeenth century, the first internal use to fill bony defects was reported in 1892 by Dressmann [19]. The application of plaster of Paris as a bone void filler, and the use of antibiotic-laden plaster in the treatment of infected bony defects, has been supported by various studies [20–23]. Calcium sulfate (CaSO4) has long been used in its partially hydrated form. Medical-grade calcium sulfate is crystallized in highly controlled environments producing regularly shaped crystals of similar size and shape. It possesses a slower, more predictable solubility and reabsorption. This material typically dissolves in vivo within 30–60 days depending on the volume and location. The chief advantages are that it can be used in the presence of infection and it is comparatively cheaper. Since it is bioabsorbable, it has inherent advantages over other antibiotic carriers, such as polymethyl methacrylate, which become a nidus for further infection after elution of the antibiotics, thus requiring a separate operation for removal from the surgical site. When this is combined with the eradication of dead

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space and the acidic environment created during its resorption, the compound can be an extremely effective treatment for acute bony infections with bone loss.

Calcium Phosphate The earliest application of calcium phosphate salts was in the form of powders [25]. The most commonly used calcium phosphate ceramics are hydroxyapatite (coral based or synthetic) and tricalcium phosphate, used in the form of implant coatings and defect fillers. These materials require high temperature and high pressure processing to produce dense, highly crystalline, bioinert ceramics, which are not moldable intraoperatively and also have poor fatigue characteristics. Porous Coralline Ceramics Chiroff et al. [26] first recognized that corals made by marine invertebrates have skeletons with a structure similar to both the cortical and cancellous bone, with interconnecting porosity. There are two processes for manufacturing coralline implants. One approach is to use coral directly in calcium carbonate form. These materials are called natural corals. The trade name for natural coral is biocoral. The other process is replamineform process that converts calcium carbonate to hydroxyapatite [24]. Although there are hundreds of genera of stony corals, Porites and Goniopora are the only two genera meeting the required standards of pore diameter and interconnectivity [24]. The exoskeleton of the genus Porites is similar to the cortical bone, and the exoskeleton of the genus Goniopora has a microstructure similar to the cancellous bone. The products are trade named either Pro Osteon or InterPore porous hydroxyapatite. A version of hybrid, coralline product has also been developed. It is a composite of calcium carbonate and calcium phosphate. A calcium phosphate layer, largely hydroxyapatite, is formed on the calcium carbonate pores. The thickness of the hydroxyapatite layer is adjusted to alter the resorption rates. In December 2001, a synthetic porous-coated hydroxyapatite (PCH) bone substitute has been launched in Europe. It is a porous calcium phosphate scaffold with a biomimetic coating – first-generation tissue-engineered product. Its surface structure resembles that of the natural bone, which makes it osteoconductive. Tricalcium Phosphate Tricalcium phosphate (TCP) is a bioceramic with bioabsorbable and biocompatible character, but its inadequate porosity, comparatively small grain size and its rapid dissolution (six weeks), makes it a poor bone graft substitute. Biocompatible and resorbable calcium phosphate cement has been introduced for augmentation of fracture repair. The chemical composition and crystallinity of the material are similar to those of the mineral phase of the bone. It undergoes the same in vivo remodeling as normal bone to reestablish the bone morphology and strength. BSM – bone substitute material is a poor crystalline calcium phosphate cement with favorable absorption characteristics and easy intraoperative handling characteristics [27]. It is hydrated with saline to form a workable paste, which remains formable for hours at room temperature but hardens within 20 minutes at physiologic body temperature (Fig. 10).

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Fig. 10 Tricalcium phosphate of 0.2–0.5 mm powder granules

Collagen Type I collagen is the most abundant protein in the extracellular matrix of the bone. It has a structure that is conducive to promoting mineral deposition, and it binds the non-collagenous matrix proteins, which initiate and control mineralization by itself. Collagen functions poorly as a graft material, but when coupled with bone morphogenetic proteins, osteoprogenitor precursors, or hydroxyapatite, it enhances incorporation of grafts significantly. The fibrillar collagen is highly purified collagen obtained from bovine dermis. Autologous bone marrow aspirate can be added to these materials or it can be mixed with autologous bone as a bone graft extender. It does not offer structural support by itself and its movement may be difficult to control [28, 29]. A mineralized collagen sponge, launched in Europe for clinical use in 2000, is shown in Fig. 11. Each microscopic type I collagen fiber is coated with hydroxyapatite; these fibers are then fabricated and cross-linked into a three-dimensional, continuously porous, and stable final format (Fig. 12). It can be mixed with bone marrow aspirate to provide osteogenic and osteoinductive potential. Another novel bone-inducing protein, MP52, is integrated with mineralized collagen sponge bone graft substitute, to induce bone formation. MP52 is a member of the bone morphogenetic protein (BMP) family. Collagen, the major constituent of connective tissues and the major structural protein of each organ, is of particular interest as a natural polymer for obtaining drug delivery systems. It acts as a hemostatic and promotes the new tissue granulation and wound epithelialization, functioning as a dressing for different types of wounds. However, collagen itself cannot produce the healing of an infected wound because it is a protein in nature, and bacteria can use it as a substrate. But combined with suitable antibiotics highly efficient drug delivery systems for wound treatment can be obtained, due to a potential synergistic effect. An ideal drug delivery system must show a low allergization quota, stability at body temperature, tissue compatibility, bactericidal activity, high bacterial resistance, broad activity spectrum, and low resorption rate.

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Fig. 11 Collagen sponge 20  20 mm 19,2 mg from porcine origin and 12,8 mg disodium hydrogen phosphate

Fig. 12 SEM images of the collagen sponge obtained by lyophilization: (a) without conductive layer; (b) with conductive layer (1000)

Antibiotics’ local delivery has been studied successfully in clinics for aminoglycosides like gentamicin and tobramycin as well as for minocycline, tetracycline, teicoplanin, or sulbactam-cefoperazone. Although doxycycline is a bactericide for a broad spectrum of bacteria and inhibits the action of collagenase, so far it is found in very few drug delivery systems. In order to improve the biochemical and mechanical properties of support and control the release of drugs, the collagen has to be cross-linked, the usual crosslinking agent being glutaraldehyde [30] (Figs. 13 and 14). Tissue engineering is a rapidly evolving multidisciplinary field that applies the principles of biology and engineering in order to develop tissue substitutes to restore, repair, or improve the function of diseased or damaged human tissues. Cell-based therapies might therefore offer hope for a number of diseases, in particular those in which single pharmacological agents are not sufficient. The success of numerous therapies in regenerative medicine requires the ability to control the formation of stable vascular networks within tissues. The formation of new blood vessels, or neovascularization, is mediated, in part, by the interaction between endothelial cells

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Fig. 13 Collagen membrane 20  20 mm 5 mg of collagen bovine origin with 1 mg doxycycline and 0.001 mg glutaraldehyde

Fig. 14 Collagen membrane 20  30 mm from bovine origin

and insoluble factors of the extracellular microenvironment. These interactions are determined by the chemical, physical, and mechanical properties of the matrix. In tissue engineering, the role of the scaffold is comparable to the role of the extracellular matrix and consists in supporting the development of cells and tissue. Collagen is a significant constituent of the natural extracellular matrix (ECM) and plays an important role in the formation of tissues and organs, being involved in the functional expressions of cells. Collagen scaffolds have been used in a variety of applications due to a number of valuable characteristics like low antigenicity, high biocompatibility, and hemostatic and cell-binding properties. It is now evident that collagen and collagen-derived fragments control many cellular functions, including cell shape and differentiation, migration, and synthesis of numerous proteins. Several in vitro studies of cell-scaffold interactions and tissue synthesis as well as in vivo studies on induced tissue and organ revealed the excellent biological performances of collagen. To promote cell adhesion and growth, a biologically active scaffold must satisfy a number of features. The scaffold biomaterial has to

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be biocompatible and degrade in the body at a controlled rate; the average pore diameter must be large enough for cells to migrate through the pores and small enough to provide a critical total surface area for appropriate cell binding and has to preclude the risk of infection during applications. An essential element in graft procedure is the blood supply. The prevention of implant failure caused by hypoxia and following infection is still a challenge. However, in general, cell-based tissue engineering provides a successful treatment in wound healing disorders. It is known that endothelial cells that form the inner lining of blood vessels participate in important physiological processes including exchanges of molecules, coagulation, and wound healing. Also these cells are essential for vascularization of the new tissue during the wound healing and tissue formation processes. A requirement for promoting faster vascularization is the presence of large pore sizes into the scaffolds. The scaffolds made up of porous collagen matrices provide a three-dimensional (3D) structure which has a significant effect on cellular activity. Three-dimensional cell culture systems offer a milieu to study biosecretory, migratory, and proliferative functionality. Embedding of endothelial cells in threedimensional collagen-based matrices allows them to grow and attain confluence in a controlled environment. Such constructs permit endothelial cells to retain a quiescent state, the secretion of essential regulatory factors, and the associated potential for vaso-regulatory control, within matrices (vehicles) that can be stored, manipulated, functionally validated, and implanted at sites protected from environmental forces [31].

Future Directions Considerable interest has developed in creating osteoconductive matrices using nonbiologic materials. Degradable polymers, bioactive glasses, and various metals have been studied. The advantage of nonbiologic materials includes the ability to control all aspects of the matrix, avoidance of immunologic reaction, and excellent biocompatibility. Polylactic and polyglycolic acid polymers have been used extensively as suture materials and biodegradable fracture fixation implants. These materials have the advantage of being assembled in various forms and can be integrated with growth factors, drugs, and other compounds to create multiphase delivery systems. A synthetic bone graft scaffold, tissue engineered from amorphous D, L-polylactide-co-glycolide (PLG), is designed to resorb within 12–20 weeks following implantation. They provide a porous architecture for the ingrowth of new bone and then fully degrade. Hydroxyapatite coating of metal surfaces enhances ingrowth and direct bonding of the bone to porous surface [32, 33]. Essentially, these coatings can be used on implants with relatively simple surface geometry and use excessive high temperatures. This means that it is difficult to coat implants with complex surface geometry (e.g., porous surface) and that no biologically active agents can be added to the coating during the spraying process. A technology has been developed that allows the growth of a thin layer of bone-like ceramic over medical devices [34].

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The calcium phosphate coating is grown from an aqueous fluid at ambient temperatures. In contrast to conventional technologies, these “biomimetic” coatings can be applied on to surfaces with complex geometry, and active agents such as growth factors or antibiotics can be coprecipitated. This creates the possibility of using these coatings as slow release systems [35]. Alloplastic bone substitute has a documented 4–6-month period from integration to regeneration, but with a very low capacity of preserving its initial volume at the time of integration. Advances in tissue engineering and the integration of the biological, physical, and engineering sciences will create new carrier constructs that regenerate and restore functional state. These constructs are likely to encompass additional families of growth factors, evolving biological scaffolds, and incorporation of mesenchymal stem cells. Ultimately, the development of ex vivo bioreactors capable of bone manufacture with the appropriate biomechanical cues will provide tissue-engineered constructs for direct use in the skeletal system. There are some new bone tissue engineering products developed for application in dental implants, revision surgery, and spinal fusion. The bone marrow cells are harvested from the patient, then multiplied in culture, shaped in appropriate structure on a scaffold, and implanted into the patient. The process takes 4 weeks. Once considered a fantasy, there is a compelling evidence to support the utility of gene therapy for bone induction in humans [28]. Studies have successfully demonstrated several safe, effective strategies to form new bone via gene therapy in animals. Gene therapy involves the transfer of genetic information to cells. When a gene is transferred to a target cell, the cell synthesizes the protein encoded by the gene. The duration of protein production that is required and the anatomic location where the protein must be delivered determine the strategy employed. The gene therapy used for bone induction is short-term, regional therapy. The gene can be introduced directly to a specific anatomic site (in vivo technique), or specific cells can be harvested from the patient, expanded, and genetically manipulated in the tissue culture and then reimplanted (ex vivo technique). The vehicle for gene delivery can be either viral (adenovirus, retrovirus) or non-viral (liposomes, DNA-ligand complexes). The gene can be selectively transferred to a targeted cell osteoblast, fibroblasts) at the bone induction site.

Principle of Guided Bone Regeneration Guided bone regeneration (GBR) membranes were originally developed to promote new tissue growth within a protected volumetric defect for periodontal regeneration. The desire to promote new bone growth without resorting to grafting procedures led to the widespread use of this technique in implant surgery. The main aim is to allow ingress of bone cells to promote bone formation within the defect. The original membranes were expanded polytetrafluoroethylene (PTFE, Gore-Tex™). It is a non-resorbable material that requires removal at second-stage surgery. The need for removal led to the development of resorbable membranes made of synthetic polymers such as polylactate and polyglycolic acid, as well as collagen

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membranes. Resorbable materials should be functional between 3 and 6 months after insertion. However, they may be resorbed or lose their shape too quickly and limit the amount of bone regeneration achieved. The creation and maintenance of a volumetric defect are critical, and this may be improved by reinforcing PTFE membranes with titanium strips. Further enhancement of the space-maintaining properties is achieved by the use of fixation pins and screws that serve to “tent” the membrane. This can be achieved also by placing small bone chips within the defect, which will also act as osseoconductive/osseoinductive grafts. The cortical bone within the defect is perforated with surgical burs to promote osteogenic cells to occupy the space created. Guided bone regeneration is also commonly used at the time of implant placement to repair small fenestrations and dehiscences around implants. It should be remembered that any such bone created will not function to stabilize the implant at the time of placement and that initial implant stability remains the overriding priority. The amount of new bone created can be quite substantial but the degree to which it becomes osseointegrated is variable. The amount of the bone to implant contact of the newly generated bone is thought to change with time and loading. Guided bone regeneration is most predictable when attempting to increase the buccolingual dimension, but increasing the vertical dimension using this technique is very difficult and unpredictable. Although membranes are designed to allow the passage of nutrients, they nonetheless tend to compromise the blood supply to the overlying soft tissues. This can result in breakdown of the soft tissues, exposure of the membrane, and infection. This may result in a net loss of the bone in the surgical area or failure of implants to osseointegrate. It is important that all incisions are kept as remote from the grafted area as possible and that the wound is sutured hermetically and without undue tension. Supporting the tissues with sutures that take large bites and “sling,” the flaps can reduce tension at the wound edges. A variety of techniques and materials has been used to establish the structural base of osseous tissue for supporting dental implants. GBR is a surgical concept which has been in clinical use for well over two decades. It has undergone several developments and improvements and is nowadays considered a predictable treatment modality, once the previously described issues have been taken fully into account. Doubts have previously been raised regarding the quality and lasting capability of membrane-regenerated bone when being put into clinical function. Previous experimental studies have clearly shown the positive dynamics of this type of bone over time. Recently this has also been confirmed in several clinical studies. In a recent review study by Aghaloo and Moy [36], the GBR technique was compared to different types of grafting procedures such as autogenous onlay, veneer (OVG), interpositional inlay grafting (COG), distraction osteogenesis (DO), and ridge splitting (RS). The data originated from a database search which identified 526 articles. Finally 335 articles met the criteria. Implant survival rate was 95.5 % for GBR technique, 90.4 % for OVG, 94.7 % for DO, and 83.8 % for COG. Hence GBR technique performed better or equal to more advanced bone grafting procedures using autogenous bone. Another interesting clinical finding is that it seems slightly

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easier to augment the bone in the maxilla compared to the mandible. The use of provisional restoration during the healing period seems to improve the result. Early implant placement also seems to be preferable if possible, due to alveolar ridge preservation, more favorable defect morphologies, and probably a higher regenerative capacity of the adjacent bone [37]. Bone augmentation of the atrophic jaw bone and, particularly, in the esthetic zone in the maxilla is a delicate and techniquesensitive procedure. The principle of GBR offers an alternative that is less resource demanding and also results in less morbidity for the patients. Predictable results can be obtained if a thorough understanding of the biological principles is applied in the clinical setting. When bone reconstruction of the posterior atrophic maxilla is needed, different surgical procedures can be used. Following the loss of teeth, alveolar ridge resorption leads to a combined vertical and horizontal reduction of the bony support, at the same time increasing maxillary sinus pneumatization. Such a condition makes implant placement impossible, either because of insufficient vertical bone volume or an alteration of the intermaxillary relationship which is not compatible with prosthetically guided implantology. As for other areas, bone augmentation procedures can be performed prior to placement of the implant, concurrent with implant placement, or subsequent to it. Up to 10 years ago, autogenous bone was considered to be the gold standard in the reconstruction of atrophic areas of the jaws, due to its osteoconductive, osteoinductive, and regenerative properties; it should still be chosen as the most suitable material in severe atrophies (classes V and VI according to Cawood and Howell [38]). In the treatment of smaller defects (classes III and IV), bone substitutes of synthetic and xenogeneic origin have been playing an important role in implant surgery. All those materials, generally called “biomaterials,” are able to favor the adherence of cells and tissue regeneration, thanks to a variable degree of osteoconductive activity; after being tested for several years through randomized prospective or retrospective studies, they can be considered as a reliable way to rebuild the bone. Something fundamental is to remember that the difficulty in rebuilding a bone defect is more related to its extent than its depth. In other words, a very deep but localized defect is easier to treat than a superficial but extensive one. Careful preoperative analysis of the site to be regenerated is important instead of deciding at the time of surgery.

Biological Principles of Guided Bone Regeneration Over the last two decades, the development of the technique of guided bone regeneration (GBR) has had a significant impact on esthetic reconstruction in conjunction with implant therapy. This technique involves the use of physical barrier membranes during the healing phase in order to avoid ingrowth of undesired tissue types into a wound area [39–43]. During the 1980s the principle of guided tissue regeneration (GTR) was developed for regenerating periodontal tissues lost as a

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result of inflammatory periodontal disease. A series of studies documented the possibility of excluding undesirable cells from populating the wound area by means of membrane barriers. This favors the proliferation of defined tissue cells to produce a desired type of tissue. The principle of physical sealing of an anatomic site for improved healing of certain tissue types is by no means new. During the mid-1950s attempts were made for neural regeneration by the use of cellulose acetate filters [44]. GBR refers more precisely to the goal of the membrane application than guided tissue regeneration. This concept promotes bone formation by protection against an invasion of competing, non-osteogenic tissues. To this end, bone defects are tightly covered by a barrier membrane of defined permeability and excellent biocompatibility. Experimental studies have proven that certain tissues within the body possess the biologic potential for regeneration if the proper environment is provided during healing. The ultimate goal of GBR is to use a temporary device to provide the necessary environment so the body can use its natural healing potential and regenerate lost and absent tissues. The efficacy of barrier membranes in conjunction with bone healing and reconstructive therapy is probably the result of a combination of different mechanisms – mechanical, cellular, and molecular. Examples of these are: • • • • •

Prevention of fibroblast mass action Prevention of contact inhibition by heterotopic cell interaction Exclusion of cell-derived soluble inhibitory factors Local concentration of growth stimulatory factors Stimulatory properties of the membrane itself

The basic studies on GBR have shown that the sequence of healing occurring in regular fracture repair follows the same basic pattern that is found in osseous lesions during GBR therapy. Based on the scientific evidence available, it can be stated that certain conditions must be met for new bone formation to be predictably accomplished by GBR: 1. There must be a source of osteogenic cells. Viable bone must be present adjacent to the defect where regeneration is desired. 2. An adequate source of vascularity is essential. This supply originates mostly from the adjacent bone surface (Volkmann’s canals and marrow compartment). 3. The wound site must remain mechanically stable during healing. 4. An appropriate space must be created and maintained between the membrane and the parent bone surface. 5. Soft connective tissue cells must be excluded from the space created by the membrane barrier. The structure of the material used must be able to accomplish this. The general function of a membrane used for GBR therapy is to create an environment that will allow the normal healing process to form the bone in a defined

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region. Hence, the host tissue-biomaterial interaction should not interfere with bone formation and maintenance to a clinically significant degree. Biomaterial chemistry and structure should result in minimal foreign body response. Optimal bonebiomaterial interaction characteristics are also desirable. A GBR membrane that allows close adaptation of bone tissue will allow more complete fill of the space defined by the membrane and stabilization of the membrane within the overall system.

Biological Factors Influencing the Reconstruction of the Alveolar Bone An absolute prerequisite for implant treatment is the availability of sufficient alveolar bone to support and retain the endosseous implant [45]. Factors such as infection, cystic lesions, tooth-alveolar trauma, or congenital tooth agenesis cause a reduction of the alveolar ridge dimensions to a varying degree. With the increasing drive for optimal esthetic outcome of implant treatment, restoring both the hard and soft tissue levels is essential [46]. Tooth replacement in the anterior maxilla is a demanding treatment, since the absence of, or poor, preoperative planning or the choice of an inappropriate treatment approach can lead to everything from esthetic shortcomings to real disasters. Esthetic complications can be related to malpositioned implants and the choice of inappropriate prosthetic components. The most critical factors, however, are the anatomic causes that include bone deficiencies in the horizontal or vertical dimensions and often a combination of the two. This is not infrequently associated with soft tissue defects of the alveolar ridge. Alveolar atrophy and anatomic alterations will have a negative influence on the proper buccal-palatal position of the implant [46, 47]. Malposition of the implant may have effects on the shape, emergence profile, and interproximal contour. It is important for the clinician to understand that the anatomic contour of the ridge comprises the soft tissue and the underlying supporting bone tissue in all directions. Hence, the soft tissue contour is heavily influenced by the bone anatomy present. The concept of the so-called biological width has increased the knowledge and understanding of the interaction between the different tissue types and different biomaterial surfaces [48]. In brief, the soft tissue demonstrates relatively constant dimensions in thickness; the peri-implant soft tissue thickness is about 3–4 mm. It is slightly thinner on the buccal aspect and more pronounced at the interproximal areas. The soft tissue is also slightly thicker in the anterior maxillary area in contrast to the posterior region of the mandible, which demonstrates the thinnest portion in the oral cavity.

Tissue Integration The phenomena of ingrowth and surface bonding of tissue to a biomaterial are termed integration. Surface and microstructural characteristics are usually responsible

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for these events. The clinical benefits of GBR membranes that have the capacity to integrate with surrounding tissues are a result of a more mechanically stable (and therefore predictable) wound healing environment. While tissue integration appears to be necessary for optimal performance of a GBR membrane, chemical and structural properties that encourage tissue integration must be balanced with the overall functional needs for alveolar ridge augmentation.

Membrane Design Criteria and Material Selection The acceptance of membrane-assisted regeneration of osseous lesions in the oral cavity has introduced reconstructive dentistry to new therapeutic procedures and biomaterials. Clinicians are being exposed to an increasing number of membrane materials used, or proposed for use, in GBR. In order to select the best material for a specific clinical indication, it is imperative to understand the functional requirements demanded of membrane barriers for GBR procedures. If the only requirement of a membrane material used in GBR was to provide a barrier to the proliferation of fibrous connective tissue, any suitable biocompatible material in the form of a cell-occlusive film could be used in clinical practice. However, a membrane that is used for alveolar ridge augmentation must meet a number of requirements in addition to acting as a passive physical barrier: 1. The membrane must be constructed of acceptable biocompatible material. The interaction between the material and the tissue should not adversely affect the surrounding tissue, the intended healing result, or the overall safety of the patient. 2. The membrane should exhibit suitable occlusive properties to prevent fibrous connective tissue (scar) invasion of the space adjacent to the bone and provide some degree of protection from bacterial invasion should the membrane become exposed to the oral environment. 3. The membrane must be able to provide a suitable space into which osseous regeneration can occur. Space making provides necessary volume with specific geometry for functional reconstruction. 4. The membrane should be capable of integrating with or attaching to the surrounding tissue. Tissue integration helps to stabilize the healing wound. It helps to create a “seal” between the bone and the material and prevent fibrous connective tissue leakage into the defect and retards the migration of epithelium around the material should it become exposed. 5. The membrane must be clinically manageable. Two different types of membrane are the most commonly used on the market. Non-resorbable, e-PTFE (expanded polytetrafluoroethylene) membrane was the first material to be successfully applied for GBR. The use of this material requires a

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removal procedure once the healing is completed. Although this material demonstrates superior biological response, the extra steps in clinical handling led to the development of biodegradable membranes such as collagen or synthetic polymers. Furthermore, attempts have also been made using other types of barrier membranes such as lyophilized dura, calvarial bone, and peritoneal tissue. However, the results regarding these latter materials are still somewhat limited in applications related to implant treatment.

Biocompatibility When discussing the clinical outcome of membrane materials, biocompatibility is a fundamental requirement for acceptable function of any implantable medical device. Although this requirement is often taken for granted, tissue interactions involve many application-specific factors that are governed by complex mechanisms. A classical definition of biocompatibility by Williams [49] is “the state of affairs when a biomaterial exists within a physiological environment, without the material adversely and significantly affecting the body.” This should be interpreted with regard to biomaterial used, the indication, and the environment in which the material is placed and maintained. For example, degradable materials are clearly affected by the environment of the body; however, safe degradation is one of the primary intended functions of this class of biomaterials. Biocompatibility is a relative term. All implanted materials interact with the host tissue to some extent. Biomaterials with dissimilar chemical composition or biomaterials with the same chemical composition but with different macro- and microstructure will demonstrate different cellular or systemic responses.

Non-resorbable Membranes As previously described, non-resorbable membranes were the first materials successfully used in GBR. The best documented material is e-PTFE (expanded polytetrafluoroethylene) (Gore-Tex Augmentation Material, W.L. Gore & Ass. Inc., Flagstaff, AZ, USA). Originally used in medical applications such as synthetic vascular graft and heart patches, the material has an extensive documentation with regard to tissue response and safety. In the 1990s, barrier membranes with special characteristics for GBR were developed and became commercially available. Substantial experimental and clinical data are available for this type of membrane. The biological response is near to ideal, and the e-PTFE membrane is still considered the “gold standard” in membrane technology [50]. The material is made up of carbon and fluorine chains strongly bonded to each other. This creates a highly chemically stable and hydrophobic material which is ideal for biocompatible tissue interaction.

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The material has an open outer structure for early tissue integration and stabilization and an inner portion which is responsible for the occlusive properties of the device.

Biodegradable Barrier Membranes Collagen membranes are resorbed by enzymatic degradation, while synthetic polymers are resorbed via degradation into lactic acid and water. Bio-Gidetm was the first collagen barrier membrane designed for GBR and is by far the best documented product in the literature. It is made from native, noncross-linked collagen types I and III and consists of two functional layers. The compact layer is cell occlusive and fulfills barrier function, while the porous layer allows tissue integration. The membrane has hydrophilic properties, which enable self-adherence to the bone surface, thus providing easy clinical handling. Due to the lack of stiffness, collagen membranes are usually used in combination with bone chips or bone substitutes. Pure synthetic biodegradable membranes are also available on the market. These materials usually consist of a combination of PLA/PGA (polylactide and polyglycolide). An example of such membranes is RESOLUT (W.L. Gore & Ass. Inc., Flagstaff, AZ, USA) [51]. Although this demonstrates excellent biological behavior in experimental studies, the amount of clinical data is still somewhat limited.

Resorption Patterns of the Alveolar Ridge As the alveolar crest is wider in the premolar and molar area than in the anterior region, vertical resorption is slower in the posterior parts. On the other hand, molars and premolars are usually lost earlier than incisors. Thus when patients seek implant placement, the posterior regions are usually equally or more resorbed than the anterior. In the posterior molar area, resorption usually results in a wider arch and a wider crest as resorption reaches the oblique and mylohyoid lines. In the first molar and especially in the premolar area, two different resorption patterns are seen, although combinations of the two are common. Whether the resorption mode is due to the angulation of the alveolar crest or genetically determined is not known. The usual pattern is vertical resorption. This results in a flat and rather wide crest. If the resorption is moderate, leaving more than 10 mm height of bone superior to the mandibular canal, implant placement is usually uncomplicated. However, as resorption continues, the height above the mandibular canal is reduced, and in advanced cases, where resorption reaches the level of the mandibular canal, the foramen is found on the top of and even slightly lingual to the top of the crest with part of the alveolar nerve positioned under a thin layer of the bone or even directly under the alveolar mucosa. The other, less frequent, resorption pattern is lateral resorption. This results in a narrow crest, in advanced cases in a very high, thin ridge made up of cortical bone which is totally unsuitable for implant placement.

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Surgical Techniques The choice of procedure for reconstructing the posterior areas of the maxilla depends both on the depth and the extent of the defect. Starting from the post-extraction defects up to the most severe atrophies, the classification according to Cawood and Howell [36] clarifies, in a simple way, what kind of surgical technique is the most suitable for each specific bony defect. In accordance with that classification, the defects can be listed as follows: • • • • •

Post-extraction sites Horizontal defects (including dehiscences and fenestrations around implants) Vertical defects Combined (vertical and horizontal defects) Sinus elevation

Post-extraction Sites An alveolar bone loss of 23 % in the first 6 months after tooth extraction and of 11 % during the following 5 years was demonstrated by Carlsson [52]. Alveolar bone loss not only reduces the amount of bone available for adequate support of the prosthetic load but can also adversely affect the implant position, the peri-implant hard and soft tissue anatomy, and, consequently, the final esthetic and functional outcome. The immediate insertion of an implant in a post-extraction site would therefore preserve a greater amount of alveolar bone and also reduce the treatment time. Whenever an immediate implant placement cannot be carried out due to the impossibility of obtaining primary stability for the fixture, physiological resorption due to the remodeling processes must be avoided. This goal can be easily achieved by filling the alveolus with bone substitutes and a free gingival graft, and implant placement surgery can be carried out about 6 months later [53]. Where there is an infection in the extraction area, immediate implant insertion should be avoided and the surgery postponed until 40–60 days after the extraction.

Immediate Technique When the gap between implant and alveolus is less than 1 mm, tissue regeneration is entrusted to the clot. For larger gaps, bone substitutes can be used to fill the defect, and the area is covered with a collagen membrane and a free gingival graft, which improves the quantity and the quality of the soft tissues around the implant at the time of connection of the healing abutment (Fig. 15). The soft tissue graft is obviously useless in cases of immediate temporary crown placement.

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Fig. 15 Representative clinical case for immediate technique: (a) post-extraction site; (b) immediate implantation; (c) immediate augmentation

Delayed Technique In this situation, 6–8 weeks are needed to rebuild the gingival tissue after extraction, and implant placement is performed by means of a standard procedure.

Horizontal Defects These defects occur where there is adequate height but insufficient width of the ridge. They include implant dehiscences and fenestrations and can be treated with GBR procedures associated with autogenous bone chips and bone substitutes as well as ridge expansion techniques (at implant placement time) or autogenous bone blocks (prior to the implant placement).

Dehiscences and Fenestrations Dehiscences are exposures of the implant at the level of its head, occurring when the thickness of the ridge is insufficient in the most coronal area. Fenestrations are exposures of the middle third or the apical portion of the implant with or without involvement of the implant head.

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A bone regeneration technique using autogenous chips, bone substitutes, and resorbable or non-resorbable membranes is a possible treatment. Most dehiscences and fenestrations can easily be treated with a resorbable membrane acting as a stabilizer of the underlying material (granules or chips). When the exposed surface of the implant exceeds 1 mm outside the bone envelope and a large volume of bone has to be regenerated, non-resorbable e-PTFE membranes are indicated. In both cases, membranes have to be stabilized with pins or mini-screws, and periosteal releasing incisions of the buccal flap are mandatory to avoid tension at suturing. Vertical mattress and single sutures are required in all the techniques.

Autogenous Intraoral or Extraoral Blocks Harvesting bone blocks from inside the mouth (chin, mandibular body, or ramus) is indicated in the correction of a class IV atrophic ridge, particularly in the case of an extended narrow ridge. When a larger amount of bone is necessary (such as a resorbed maxillary ridge in edentulous patients), blocks can be harvested from the hip or the calvaria. In all these situations, attention must be paid to the block’s resorption during the 4 months’ healing. An average resorption of 20–30 % can be observed [54, 55] at implant placement time, in blocks harvested from the chin or the hip, due to their cancellous component, and 15 % resorption occurs in blocks harvested from the mandibular body and the ramus, which are basically made of cortical bone. An original technique to reduce the resorption of autogenous blocks was presented in 2005 [56], involving a layer of anorganic bovine bone placed on the top of the block and covered by a collagen membrane. Using this procedure it is possible to maintain the original size of the block; it can be successfully used for either intraoral or extraoral grafts [57].

Vertical Defects If the bone height is insufficient to guarantee long-term implant stability or the prosthetic rehabilitation would result in too long crowns, vertical ridge augmentation is mandatory. Vertical augmentation can be achieved by means of corticocancellous blocks harvested from the chin or the ramus or, alternatively, with cancellous bone chips, bone substitutes (Figs. 16 and 17), and a titanium-reinforced e-PTFE membrane [58, 59].

Vertical GBR with Membrane This procedure is very predictable, but it has to be carried out strictly according to the surgical protocol, in order to limit the risks of the membrane exposure. The most common flap design comprises a full-thickness mid-crestal incision within the keratinized mucosa of the edentulous ridge, extended mesially and

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Fig. 16 Vertical defect caused by an apical infection of the two upper central incisive

Fig. 17 Same vertical defect after 3 months treated with allogeneic and xenograft bone chips mixed in a ratio of 1:1

distally to at least one adjacent tooth. Vertical releasing cuts are performed at the mesial and distal line angles of the incision. A proper preparation of the recipient site is crucial for new bone formation. The buccal and palatal flaps are reflected and gently managed to avoid any perforation of the flap. Stainless steel mini-screws are used as “tent poles” to prevent collapse of the membrane and to predetermine either the width or the height of the future alveolar ridge. The mini-screws are placed and left to protrude out from the bone level to the expected height. The cortical plate is then drilled with a round bur to expose the cancellous bone and to provoke some bleeding. The titanium structure of the e-PTFE membrane is bent with pliers to adapt it to the ridge anatomy, and it is trimmed with scissors to extend at least 4–5 mm beyond the margins of the defect. Once placed over the surgical recipient site, the membrane is secured to the lingual and palatal aspect of the bone crest with fixation mini-screws or pins. The cancellous autogenous chips mixed with the bone substitute are placed to reconstruct the defect, and the buccal portion of the membrane is adapted to the vestibular bone plate and also secured with screws or pins. A releasing horizontal incision of the periosteum is now performed to give elasticity to the flap and obtain tension-free adaptation at closure.

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The two margins of the flap can be considered sufficiently released when they overlap by at least 7–10 mm. Closure is done with horizontal mattress sutures first and with interrupted sutures later. The membrane is usually removed 6 months after surgery, at implant placement time.

Sinus Elevation When atrophy of the posterior maxilla reduces the amount of bone suitable for implant lacement and a bone augmentation procedure is required, first of all one needs to evaluate if the sinus really has migrated from apical to coronal toward the margin of the alveolar ridge or if the sinus is in its previous position and the vertical height loss is due to vertical resorption. In the first case, a sinus elevation procedure is indicated, while in the second situation, vertical ridge augmentation should be performed without any sinus involvement. Once the necessity of performing a sinus elevation has been decided upon, the second step is the choice between a one- and a two-stage procedure. According to the recent literature [60], a residual alveolar ridge of 2 mm is the absolute minimum height for bone augmentation and simultaneous implant placement, and a two-stage technique must be done where there is a residual height of 0–2 mm (Fig. 18). The surgical procedure is the well-known Boyne and James technique [61], and autogenous bone or bone substitutes have been used for a long time to fill up the subantral cavity. At present, long-term prospective and retrospective studies confirm that bone substitutes are able to regenerate new bone without harvesting autogenous bone and that the implant survival rate in augmented sinuses with biomaterials is significantly higher than that obtained using autogenous bone chips [62–64]. In Cawood and Howell1 class VI defects, only autogenous bone is recommended, since the severe atrophy needs all the power of an osteoproliferative material [65]. This means that at least 80 % of sinus elevation procedures can be done by using biomaterials alone (Fig. 19). What kind of bone substitute should be chosen to get the best result? All the bone substitutes currently used in the sinus elevation procedures, either xenogeniec or alloplasts, offer osteoconductive properties only. The decision should be taken after considering the human hydroxyapatite structure: the more a granule of a biomaterial is similar to a human hydroxyapatite crystal, the more it is possible to get an osteoconductive effect. Another parameter to be taken into consideration is the biomaterial’s resorption time. A material which is resorbed too quickly does not allow the osteoblasts and the new vessels to promote formation of woven bone. A material which is resorbed too slowly, by delaying its total substitution with newly formed bone, inhibits bone–implant contact which is essential for osseointegration. A resorption time of 6–10 months can be considered reasonably ideal. In the author’s experience, anorganic bovine bone has given excellent results in over 15 years of sinus elevation surgeries, giving a new bone quality close to a class 2 native bone, [66] and a very good osteoconductive property, as verified from many histomorphometric studies [67–69].

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Fig. 18 Representative radiological images for sinus elevation: (a) before sinus elevation; (b) sinus floor dimension before sinus elevation; (c) the same site after 8 months

Nevertheless, the author has been using other bone substitutes, such as beta TCP, calcium sulfate, and DFDBA, whose clinical efficacy has been demonstrated in some studies, although with a lower predictability in terms of bone quality and implant survival rate. When the height of the residual ridge is 6–7 mm, the surgeon can decide whether to use short implants or elevate the sinus floor 2–3 mm with the Summer’s osteotome technique. The procedure, wrongly named “minor sinus elevation,” is a blind procedure and should be considered with care. In order to elevate the sinus membrane with this procedure, any of the biomaterials can be equally used, since the primary stability of the implant is guaranteed by the residual ridge [70].

Preoperative Antibiotics The same antibiotic prophylaxis is recommended as in conjunction with implant placement [71–76]. According to standard protocol, usually a 1-day regimen including 2 g of penicillin per day is enough. For advanced GBR cases, 1 week’s antibiotic coverage postoperatively is recommended.

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Fig. 19 Representative clinical case for sinus elevation. (1) Elevating the Schneider membrane; (2) filling the created space with xenograft 1.2 ~ 1.7 mm of bovine origin; (3) implant direction positioning; (4) collagen membrane placing; (5) tight stitching for good sealing

Flap Design A full-thickness crestal incision, slightly buccal but still within attached mucosa, is performed and extended mesially and distally to the adjacent teeth. Diverging releasing incisions are then performed buccally. Full-thickness mucoperiosteal flaps are elevated (Fig. 20).

Site Preparation Based on the available amount of host bone present, simultaneous implant placement can be performed. If optimal position and direction and satisfactory primary stability cannot be obtained, a two-stage procedure should be considered. The bone surface in

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Fig. 20 Relevant clinical images for flap design: (a) releasing incisions performed buccally; (b) full-thickness mucoperiosteal flaps Fig. 21 Implant placement for a removable prosthesis

the augmentation area should be carefully debrided in order to remove all remnants of the soft tissue. If implant placement is performed, it should be performed according to the protocol of the implant system used, aiming at a prosthetic-driven position (Fig. 21). Prior to placement of graft material and the barrier membrane, the buccal bone plate in the defect area must be perforated to create access for multipotent cells and blood vessels emanating from the marrow cavity. This can be performed either with a spiral or a round bur with a dimension of approximately 1 mm. This surgical procedure is believed to stimulate osteogenesis by activating a cascade effect of growth factors. Furthermore it allows the formation of an appropriate coagulum which will act as a matrix for the initial bone formation.

Graft Material Positioning The use of spacemaking materials underneath the membrane has been proven to provide a more predictable regenerative result and is today considered state of the art. Many different materials, including autogenous bone chips, freeze-drieddemineralized, deproteinized bovine bone, and synthetic graft materials such as

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Fig. 22 An allogeneic and xenograft bone chips mixed in a ratio of 1:1

TCPs, have been tested with varying results. The best documented filling material with predictable outcome is a combination of deproteinized bovine bone (Bio-Oss, Geistlich, Switzerland) and autogenous bone chips mixed in a ratio of 1:1. The addition of xenograft has been shown to minimize resorption of the newly regenerated bone (Fig. 22).

Membrane Selection and Positioning The anatomic shape of the defect to be regenerated dictates the choice of membrane material. The following protocols can be recommended.

Biodegradable Membranes Most defects can be treated with resorbable membranes together with autogenous bone chips alone or in combination with bone substitutes. Usually the autogenous bone chips are placed in contact with either the bone surface or the exposed parts of the implant and then covered with a layer of xenograft chips. Alternatively the two filling materials are mixed together in a ratio of approximately 1:1. The membrane must be cut and trimmed to adapt to the anatomy of the ridge and applied over the defect in order to cover the bone graft. Due to the hydrophilic properties of the collagen membrane, it will “stick” to the bone surface once wetted either with saline or blood (Fig. 23). Hence no fixation screws or tacks are needed for stabilization in most cases.

Non-resorbable Membranes When a large volume of bone (outside the bone envelope) must be regenerated, the use of non-resorbable e-PTFE membrane is indicated. The preparation of the augmentation site is identical to that described above. However, extra attention

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Fig. 23 Collagen membrane placing to cover a bone implant site

must be paid to the membrane adaptation and fixation. The membrane must be cut to extend at least 4–5 mm beyond the filling material to avoid interference with the surrounding soft tissue. It is important to avoid creating sharp edges since they can increase the risk for membrane perforation during healing. A critical note is to trim the membrane so a distance of 1–2 mm is maintained from the root surface of the neighboring teeth. This is to avoid contamination due to bacterial downgrowth along the root and also to enhance periodontal reattachment. Finally, the membrane should be fixated using either micro-screws or specially designed tacks. It is practical to start this procedure on the palatal side prior to the placement of the bone graft material. A critical technique is the adjustment of the flaps prior to suturing. A completely tension-free environment must be created by performing periosteal releasing incisions at the base of the buccal flap.

Suturing Suturing is recommended using non-resorbable sutures in a biocompatible material. A double suture layer should be created with a combination of horizontal mattress sutures (4/0) (on top of the crest) followed by single interrupted sutures (5/0 or 6/0) for mucosal closure.

Follow-Up Due to the compromised wound, the recommendation is to maintain the sutures in place for at least 14 days. The patient should receive systemic antibiotics (amoxicillin) for 5–10 days when they have undergone a more advanced GBR procedure. In addition, the patient should rinse with chlorhexidine solution for 3 weeks after placement of the GBR barrier. This could thereafter be switched to a 1 % chlorhexidine gel which is gently applied in the wound area only once daily.

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Fig. 24 Radiological image of a big defect of the maxilla

Fig. 25 Radiological image of the same area than Fig. 24 after 8 months from augmentation

The e-PTFE membranes are removed after 6–8 months (either at the time of fixture installation or abutment connection). During this healing period, the patients are checked once a month for plaque removal and any complications.

Temporary Dentures Implant treatment and related bone augmentative procedures are usually associated with a situation where the patient needs a temporary solution during the respective healing phases. Clinical studies have demonstrated a clear correlation between membrane exposure and pressure from temporary dentures in the wound area. Hence, strict rules apply for the design of the temporary solution in conjunction with GBR. Ideally, fixed solutions such as Maryland bridges or conventional temporary bridges are the first choices if possible. If a temporary removable denture is necessary, it should be designed in such a way that no contact is present between

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Fig. 26 Radiological image of a clinical case before extraction of infected alveolars

Fig. 27 Radiological image of the same area than Fig. 26 after 4 months from augmentation for bone preservation

the base of the denture and the soft tissue covering the GBR membrane. Furthermore, occlusal support of the denture is mandatory in order to prevent a pumping pressure when chewing.

Membrane Removal A non-resorbable barrier membrane is removed under local anesthesia. Technically the easiest way to approach the membrane is from the lateral aspect and to dissect it free from the covering soft tissue layer. Following this procedure, it is usually easy to remove the barrier from the underlying bone tissue. Great care should be taken to remove the entire membrane material. This is usually performed in conjunction with either implant placement or abutment connection. Biodegradable barrier membranes do not usually require this procedure. Most biodegradable membranes on the market are designed with a resorption pattern of less than 6 months.

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Summary A variety of techniques and materials has been used to establish the structural base of osseous tissue for supporting dental implants. GBR is a surgical concept which has been in clinical use for well over two decades. It has undergone several developments and improvements and is nowadays considered a predictable treatment modality, once the previously described issues have been taken fully into account. Doubts have previously been raised regarding the quality and lasting capability of membrane-regenerated bone when being put into clinical function. Previous experimental studies have clearly shown the positive dynamics of this type of bone over time (Figs. 24 and 25) Another interesting clinical finding is that it seems slightly easier to augment the bone in the maxilla compared to the mandible. The use of provisional restoration during the healing period seems to improve the result. Early implant placement also seems to be preferable if possible, due to alveolar ridge preservation, more favorable defect morphologies, and probably a higher regenerative capacity of the adjacent bone. Bone augmentation of the atrophic jaw bone and, particularly, in the esthetic zone in the maxilla is a delicate and techniquesensitive procedure. The principle of GBR offers an alternative that is less resource demanding and also results in less morbidity for the patients. Predictable results can be obtained if a thorough understanding of the biological principles is applied in the clinical setting (Figs. 26 and 27). Complications may arise due to inadequate coverage of the graft or flaps which are stretched too much, with compromised vascular circulation. Exposure of the graft may result in loss of part or all of the graft.

References 1. Fleming JE, Cornell CN, Muschler CF (2000) Bone cells and matrices in orthopedic tissue engineering. Orthop Clin North Am 31:357–374 2. Bauer TW, Muschler GF (2000) Bone graft materials. An overview of the basic science. Clin Orthop Relat Res 371:10–27 3. Urist MR (1965) Bone formation by auto-induction. Science 150:893–899 4. Schwartz Z, Mellonig JT, Carnes DL Jr et al (1996) Ability of commercial demineralized freezedried bone allograft to induce new bone formation. J Periodontol 67:918–926 5. Shigeyama A, D’Errico JA, Stone R et al (1995) Commercially prepared allograft material has biological activity in vitro. J Periodontol 66:478–487 6. Hammerle CH, Karring T (1998) Guided bone regeneration at oral implant sites. Periodontol 2000 17:151–175 7. Valentini P, Abensur D (1998) Histological evaluation of Bio-Oss in a two stage sinus floor elevation and implantation procedure. Clin Oral Implants Res 9:59–64 8. Wenz B (2003) Characteristics of Bio-Oss and Bio-Gide. In: Simion M (ed) Advanced Techniques for Bone Regeneration with Bio-Oss and Bio-Gide (Maiorana C. RC Books, Milano, pp 73–81 9. Joyce ME, Jingushi S, Bolander ME (1990) Transforming growth factor-? in the regulation of fracture repair. Orthop Clin North Am 21:199–209

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10. Bostrom MP, Lane JM, Berberian WS, Missri AA, Tomin E, Weiland A et al (1995) Immunolocalization and expression of bone morphogenetic protein 2 and 4 in fracture healing. J Orthop Res 13:357–367 11. Onishi T, Ishidou Y, Nagamine T, Yone K, Imamaru T, Kato M et al (1998) Distinct and overlapping patterns of localization of bone morphogenetic protein (BMP) family members and a BMP type II receptor during fracture healing in rats. Bone 22:605–612 12. Sakou T (1998) Bone morphogenetic proteins: from basic studies to clinical approaches. Bone 22:591–603 13. Bourque WT, Gross M, Hall BK (1993) Expression of four growth factors during fracture repair. Int J Dev Biol 37:573–579 14. Nakamura T, Hara Y, Tagawa M, Tamura M, Yuge T, Fukuda H et al (1998) Recombinant human basic fibroblast growth factor accelerates fracture healing by enhancing callus remodeling in experimental dog tibial fracture. J Bone Miner Res 13:942–949 15. Trippel SB (1998) Potential role of insulinlike growth factors in fracture healing. Clin Orthop 355S:301–313 16. Nash TJ, Howlett CR, Martin C, Steele J, Johnson KA, Kicklin DJ (1994) Effect of plateletderived growth factor on tibial osteotomies in rabbits. Bone 15:203–208 17. Tuli SM, Singh AD (1978) The osteoinductive property of decalcified bone matrix: an experimental study. J Bone Joint Surg Br 60:116–123 18. Adkisson HD, Strauss-Schoenberger J, Gillis M, Wilkins R, Jackson M, Hruska KA (2000) A rapid quantitative bioassay of osteoinduction. J Orthop Res 18:503–511 19. Geesink RGT, Hoefnagels NHM, Bulstra SK (1999) Osteogenic activity of OP-1 bone morphogenetic protein (BMP-7) in a human fibular defect. J Bone Joint Surg Br 81:710–718 20. Heckman JD, Ehler W, Brooks BP, Aufdemorte TB, Lohmann CH, Morgan T et al (1999) Bone morphogenetic protein but not transforming growth factor-? enhances bone formation in canine diaphyseal nonunions implanted with a biodegradable composite polymer. J Bone Joint Surg Am 81:1717–1739 21. Giavaresi G, Fini M, Salvage J, Nicoli Aldini N, Giardino R, Ambrosio L, Nicolais L, Santin M (2010) Bone regeneration potential of a soybean-based filler: experimental study in a rabbit cancellous bone defects. J Mater Sci: Mater Med (2010) 21:615–626 22. Holmberg L, Forsgren L, Kristerson L (2008) Porous titanium granules for implant stability and bone regeneration—a case followed for 12 years. Ups J Med Sci 113(2):217–220 23. Bystedt H, Rasmusson L (2009) Porous titanium granules used as osteoconductive material for sinus floor augmentation: a clinical pilot study. Clin Implant Dent Relat Res 11(2):101–105 24. White E, Shors EC (1986) Biomaterial aspects of Interpore-200 porous hydroxyapatite. Dent Clin North Am 30:49–67 25. Ferraro JW (1979) Experimental evaluation of ceramic calcium phosphate as a substitute for bone grafts. Plast Reconstr Surg 63:634–640 26. Chiroff RT, White EW, Weber KN, Roy DM (1975) Tissue ingrowth of replamineform implants. J Biomed Mater Res 6:29–45 27. Knaack D, Goad ME, Aiolova M, Rey C, Tofighi A, Chakravarthy P et al (1998) Resorbable calcium phosphate bone substitute. J Biomed Mater Res 43:399–409 28. Cornell CN, Lane JM, Chapman M, Merkow R, Seligson D, Henry S et al (1991) Multicenter trial of Collagraft as bone graft substitute. J Orthop Trauma 5:1–8 29. Chapman MW, Bucholz R, Cornell CN (1997) Treatment of acute fractures with a collagencalcium phosphate graft material: A randomized clinical trial. J Bone Joint Surg Am 79:495–502 30. Albu MG, Ghica MV, Leca M, Popa L, Borlescu C, Cremenescu E, Giurginca M, Trandafir V (2010) Doxycycline delivery from collagen matrices crosslinked with tannic acid. Mol Cryst Liq Cryst 523:97 = [669]–105 = [677] 31. Titorencu I, Albu MG, Giurginca M, Jinga V, Antoniac I, Trandafir V, Cotrut C, Miculescu F, Simionescu ANDM (2010) In vitro biocompatibility of human endothelial cells with collagendoxycycline matrices. Mol Cryst Liq Cryst 523:97 = [669]–105 = [677]

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53. Nevins M, Camelo M, De Paoli S et al (2006) A study of the fate of the buccal wall of extraction sockets of teeth with prominent roots. Int J Periodontics Restorative Dent 26:19–29 54. Nystrom E, Ahlqvist J, Gunne J et al (2004) Ten year follow-up of onlay bone grafts and implants in severely resorbed maxillae. Int J Oral Maxillofac Surg 33:258–262 55. Nystrom E, Ahlqvist J, Legrell PE (2002) Bone graft remodeling and implant success rate in the treatment of the severely resorbed maxilla: a 5 year longitudinal study. Int J Oral Maxillofac Surg 318:158–164 56. Maiorana C, Beretta M, Salina S et al (2005) Reduction of autogenous bone graft resorption by means of Bio-Oss coverage: a prospective study. Int J Periodontics Restorative Dent 1:19–24 57. Maiorana C, Sommariva L, Brivio P, Sigurta D, Santoro F (2003) Maxillary sinus augmentation with anorganic bovine bone (Bio-Oss) and autologous platelet-rich plasma: preliminary clinical and histologic evaluations. Int J Periodontics Restorative Dent 23:227–235 58. Simion M, Trisi P, Piattelli A (1994) Vertical ridge augmentation using a membrane technique associated with osseointegrated implants. Int J Periodontics Restorative Dent 14:496–511 59. Simion M, Jovanovic S, Tinti C et al (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 60. Maiorana C, Simion M (2003) Chapter 3. In: Advanced techniques for bone regeneration with Bio-Oss and Bio-Gide. RC Books, Milano, pp 41–50 61. Boyne PJ, James R (1980) Grafting of the maxillary sinus floor with autogenous bone marrow and bone. J Oral Surg 38:613–618 62. Valentini P, Abensur D, Wenz B et al (2000) Sinus grafting with porous bone mineral (Bio-Oss): a study on 15 patients. Int J Periodontics Restorative Dent 20:245–252 63. Valentini P, Abensur D (2003) Maxillary sinus grafting with anorganic bovine bone: a clinical report of long-term results. Int J Oral Maxillofac Implants 18:556–560 64. Maiorana C, Sigurta D, Mirandola A et al (2005) Bone resorption around implants placed in grafted sinuses: a clinical and radiologic follow-up after up to four years. Int J Oral Maxillofac Implants 2:261–265 65. Geurs NC, Wang JC, Schulman LB et al (2001) Retrospective radiographic analysis of sinus graft and implant placement procedures from the Academy of Osseointegration Consensus Conference on sinus graft. Int J Periodontics Restorative Dent 21:517–524 66. Haas R, Mailath G, Dortbudak O et al (1998) Bovine hydroxyapatite for maxillary sinus augmentation: analysis of interfacial bond strength of dental implants using pull-out tests. Clin Oral Implants Res 17:151–175 67. Maiorana C, Redemagni M, Rabagliati M et al (2000) Treatment of maxillary ridge resorption by sinus augmentation with iliac cancellous bone, anorganic bovine bone and implants: a clinical and histologic report. Int J Oral Maxillofac Implants 15:873–878 68. Wallace SS, Froum SJ, Cho SC et al (2005) Sinus augmentation utilizing anorganic bovine bone (Bio-Oss) with absorbable and non absorbable membranes placed over the lateral window: histomorphometric and clinical analyses. Int J Periodontics Restorative Dent 25:551–559 69. Hallman M, Sennerby L, Lundgren S (2002) A clinical and histologic evaluation of implant integration in the posterior maxilla after sinus floor augmentation with autogenous bone, bovine hydroxyapatite, or a 20:80 mixture. Int J Oral Maxillofac Implants 17:635–643 70. Santoro F, Maiorana C (2005) Chapter 5. In: Advanced osseointegration. RC Books, Milano, pp 117–124 71. Dressmann H (1892) Ueber Knochenplombierung bei Hohlenformigen Defekten des Knochens. Beitr Klin Chir 9:804–810 72. Mousset B, Benoit MA, Bouillet R, Gillard J (1993) Plaster of Paris: a carrier for antibiotics in treatment of bone infections. Acta Orthop Belg 59:239–248 73. Mousset B, Benoit MA, Delloye C, Bouillet R, Gillard J (1995) Biodegradable implants for potential use in bone infection: an in-vitro study of antibiotic - loaded calcium sulphate. Int Orthop 19:157–161 74. Peltier LF (1959) The use of plaster of Paris to fill large defects in bone. Am J Surg 97:311–315

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75. Sidqui M, Collin P, Vitte C, Forest N (1995) Osteoblast adherence and resorption activity of isolated osteoclasts on calcium sulphate hemihydrate. Biomaterials 16:1327–1332 76. Scaduto AA, Lieberman JR (1999) Gene therapy for osteoinduction. Orthop Clin North Am 30:625–633

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Diana Dudea, Camelia Alb, Bogdan Culic, and Florin Alb

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Composition . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Organic Matrix . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Inorganic Fillers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Coupling Agent . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Initiator-Accelerator . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Optical Modifiers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Indications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Classification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Flowable Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Condensable Composite . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conventional Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Microfilled Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Hybrid Composite Resins . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Nanofilled Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Direct Composite Resins Restorations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Anterior Direct Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Posterior Composite Restorations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Consideration Regarding the Composite Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Problems Generated by Direct Composites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Indirect Composite Restorations in Dentistry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Advantages and Disadvantages of Laboratory Processed Inlays . . . . . . . . . . . . . . . . . . . . . . . . . . Technological Alternatives for Indirectly Processed Composite Restorations . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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D. Dudea (*) • C. Alb • B. Culic Department of Prosthetic Dentistry and Dental Materials, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania e-mail: [email protected]; [email protected]; [email protected] F. Alb Department of Periodontology, University of Medicine and Pharmacy “Iuliu Hatieganu”, Cluj–Napoca, Romania e-mail: albfl[email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_53

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Abstract

Resin composites, also named resin-based composites, composite resins, or composites, contain four major components: an organic polymer, inorganic fillers, a coupling agent, and an initiator-accelerator system. Classification of the composite resins takes into account the consistency of the material (correlated with the filler/organic matrix ratio) and, further, the size of the fillers’ particles. The area of indications of the composite materials in dentistry covers a large range of domains: from direct and indirect restorative materials, veneers, provisional restorations, inlays, onlays, crowns, sealants, cements used in adhesive cementation of the composite or ceramic crowns and bridges, inlays, root canal posts, and composite teeth for dentures. The present chapter emphasizes the characteristics of composite resins such as direct and indirect restorations, advantages and disadvantages, clinical indications and clinical protocols used to process them, clinical considerations regarding the composites properties, and practical problems that can be encountered when used. Direct composites offer rapid, minimally invasive, single-session methods of treatment. Due to the large range of materials available on the market, the esthetic results may be excellent. The clinical choice of a direct composite is based on the priority that should be given to mechanical or esthetic characteristics: if the mechanical parameters are mostly important, the material showing the highest percentage of filler is selected; in the case of special esthetic needs, the particle size is the factor that influences the selection. The indirect processed composite resins are alternatives for large restorations, on several teeth in a quadrant, when used to replace functional cusps, in patients with bruxism or parafunctional habits. Their advantages over the ceramic restorations include: repair capabilities, resilience for comfort and shock absorption, adjustable in the mouth, and no wear of opposing structures in functional contact. Keywords

Direct composites • Indirect composites • Composite resins • Anterior composites • Posterior composites • Composite inlays • CAD-CAM processed composites • Optical properties • Color • Translucency • Inorganic fillers • Coupling agents • Initiator-accelerator system • Conventional composites • Hybrid composites • Microfilled composites • Nanofilled composites

Introduction Esthetic dentistry is a continuous developing area of dental treatments, due to the patients’ increased demands in respect to the appearance of the dental arches and, consequently, of their restorations. The materials currently available for the esthetic restoration of dentition are composite resins and dental ceramics; in contrast to the ceramics that can be processed only in the dental laboratory or by using in-office CAD-CAM technology,

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Fig. 1 Direct composite resins. (a) Initial situation. Old discolored restoration on the upper incisors. (b) Clinical appearance of the restored teeth with direct composites (Clinical case Dr. Bogdan Culic)

composite resins are materials that can be used either in direct reconstruction of the dental structure (dentist-processed filling materials) or as laboratory-made fixed appliances (inlays, onlays, crowns, posts, and cores). The introduction of composite resins in dentistry is closely related to the initiation of enamel acid etching (aimed to improve the adhesion to the dental substrate, and to bis-GMA monomer development [1] (Fig. 1a, b).

Composition Resin composites, also named resin-based composites, composite resins, or composites [2], contain four major components: an organic polymer, inorganic fillers, a coupling agent, and an initiator-accelerator system [3].

The Organic Matrix The organic matrix is represented most often by a mixture of aromatic and/or aliphatic dimethacrylate monomers – most often bisphenol glycidyl methacrylate (bis-GMA or the Bowen’s resin), urethane dimethacrylate (UDMA), or a combination between the two [4]. Bis-GMA constitutes around 20 % v/v of standard composite resins [5]. However, it is a high-viscosity material that explains the addition of a low-viscosity monomer, like TEGDMA – triethylene glycol dimethacrylate [4]. Another shortcoming of bis-GMA is the contraction – its proportion in the composite mass is inversely correlated with the shrinkage [1]. One of the most important directions to improve the organic matrix is oriented toward decreasing the contraction. In this respect, the epoxy-based silorane system (Filtek Silorane LS) 3M Espe provides verified lower shrinkage than typical dimethacrylate-based resins, due to the epoxide curing reaction that involves the opening of an oxirane ring [6, 7].

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The Inorganic Fillers The inorganic fillers can be represented by fine particles of radiopaque glass or quartz, microfine particles of zirconium oxide, aluminum oxide or colloidal silica, glass fibers, nanoparticles of silicon dioxide, and silica-zirconia nanoclusters. The dimension of the inorganic fillers is the bases for the composite classification. Filler is the portion of the composite responsible for the hardness, compressive and tensile strength, and abrasion resistance; for the reduction in polymerization shrinkage; and for the volumetric stability; it is also the component that stands for the reduction in water sorption and staining and for a viscous consistency that is easier to handle during insertion into cavity [4]. The inorganic particles such as strontium, barium, glass, and zirconium are also responsible for the radiopacity. The glass fillers are more sensitive to acidic attack, and as a consequence, the glassbased composites are more susceptible to abrasive wear than silica filler-based resins. The content of filler ranges between 30 and 70 % vol or 50 and 85 wt% of a composite [4]. Addition of new formulas of fillers is experimented; the remineralizing potential of the material by adding dicalcium or tetracalcium phosphate nanoparticles is a promising direction, even if the opacity of such composites limits their esthetic appearance [6, 8, 9]. In the same idea, other formulas, based on calcium fluoride, have been developed [6, 10]. In this case, remineralization may be promoted by the slow release of calcium and phosphate ions followed by the precipitation of new calcium phosphate mineral [6, 9, 10].

The Coupling Agent The coupling agent covers the inorganic particles and creates the bond between the previous two phases. This agent is a molecule with silane groups at one end (ion bond to Si2O) and methacrylates at the other end (providing covalent bond with the resin) [1]. A properly coupling agent prevents the water penetration at the resin-filler interface, preventing leaching [4].

The Initiator-Accelerator The initiator-accelerator system allows for the curing process, either by chemical reaction (self-curing, chemically activated composites), light activated (light curing), or dual activated (chemically and light activated). The light activation is enabled by a photoactivator (camphorquinone) which absorbs the 470 nm wavelength light, responsible for the curing process; the reaction is further activated by a tertiary organic amine containing a carbon ring double bond [3]. The light-curable composites are maintained in light-proof syringes. The two components of the activation system – the photosensitizer and the amine

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initiator – contained in the paste do not interact as long as they are not exposed to light [4]. In order to reduce the yellowish effect generated by the camphorquinone, other photoinitiators have been introduced lately (PPD-1-phenyl-1,2-propanedione) [7], Lucirin TPO (monoacylphosphine oxide), and Irgacure 819 (bisacylphosphine oxide) [6, 11]. Chemical curing is enabled by the reaction of an organic amine (contained in the catalyst paste) with an organic peroxide (catalyst paste) [3].

Optical Modifiers Optical modifiers are responsible for the translucency and color. They are represented by metallic oxides that act in terms of colorants or opacifiers (titanium dioxides, aluminum oxides).

Indications The area of indications of the composite materials in dentistry covers a large range of domains: – Direct restorative materials of the anterior and posterior teeth, when the dental structure is lost due to dental decay, fractures, erosions, and attritions; – Direct composite veneers when the change of the dental shape, position, and dimension is aimed or is needed to mask a discromic tooth or to close diastemas; – Indirect (laboratory processed) veneers, inlays, onlays, and crowns; – Direct or indirect provisional restorations, during the intermediate stages of fixed prosthetic treatments; – Cavity liners in deep cavities; – Fissures and pit sealants of the recently erupted teeth, in preventive dentistry; – Resin-based cements used in adhesive cementation of the composite or ceramic crowns and bridges, inlays, onlays, and veneers; – Root canal posts; – Composite teeth for dentures [1–6]

Classification Classification of the composite resins takes into account the consistency of the material (correlated with the filler/organic matrix ratio) and, further, the size of the fillers’ particles. The viscosity-based classification of composites includes: the conventional (universal) restorative materials indicated for restorations of anterior and posterior teeth, flowable materials, and packable condensable composites. Conventional composites are used for the restoration of the anterior and posterior teeth.

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Flowable composites are used as sealants, liners, or layers aimed to enhance the adaptation of the conventional composite to the cavity walls; the resin-based cements are, also, flowable composites. Packable composites are more viscous materials used only on the posterior teeth, due to their enhanced mechanical resistance and their ability to reproduce tight contact areas [6].

Flowable Composites Flowable composites have the advantage of a highly wettability of the dental surface that ensures their penetration into narrow spaces and their ability to form thin layers; these properties made them the materials of choice for preventive methods and minimally invasive techniques of dental decays. In the first instance, dental sealants are used to block the fissures and pits on the dental surface, in order to prevent the dental plaque accumulation; the infiltration with newer flowable composite resins of demineralized enamel, in the earliest stages of dental decay, is the philosophy of the minimally invasive treatment approach. Another characteristic, the high flexibility, explains the indication in the treatment of cervical lesions (dental decays or wear cervical lesion) so they are less likely to be displaced under an increased stress The most frequent area of indication is, still, their application as liners under high- or medium-viscosity composites; they provide a thin-layer material that fills the irregularities of the prepared dental surface and favor a better application of the restorative material. In order to enhance the adhesion and to reduce the clinical steps, flowable composites with acidic adhesive monomers, typically found in dentin bonding agents, have been developed. These self-adhesive flowable composite resins recommended as liners, sealants, or restorative materials for small lesions are controversial in respect to their longevity [1, 2, 4, 6].

Condensable Composite Condensable composites are used as alternatives to amalgam fillings or indirect restorations (composite or ceramic inlays) for the posterior teeth. Due to the characteristics of the filler portion, either modified in respect to the size distributions of the particles or by addition of fibers, they are designed to resist the occlusal forces in this segment of dentition [6, 12]. In addition, due to their higher consistency, they allow for a good contact area between adjacent teeth; in this respect, the main indication area are the restorations that involve the proximal surfaces (class II) [1].

Conventional Composites Conventional composites are classified, further, according to the dimensions of the filler particles in traditional (macrofilled), hybrid, microfilled, and nanofilled composite resins.

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The first traditional dental composites had large particle fillers, 1–50 μm. Sometimes, the term conventional is used in order to refer to this class. Even if these materials had increased strength, they were difficult to polish, since the large particles that were loosen in the process generated rough surfaces. In addition, in time, the organic matrix develops abrasion faster than the filler that results in protruding particles from the surface. This uneven surface is also susceptible to staining [1, 4, 6].

Microfilled Composites Microfilled composites have inorganic particles, more often amorphous spherical silica (colloidal silica), of 0.04 μm (one-tenth of the wavelength of visible light) and the loading rate low that explains the mechanical weakness of this group of materials. Their indication in the posterior zone is limited due to the fractures that occur mostly in class II cavities. The high percentage of resin (40–80 vol%) is responsible for the water sorption and of their increased staining resistance; however, the microfilled composites are the materials of choice for the anterior zone, in the exposed dental areas, due to their highly polishable surfaces. Their filler particles are smaller than the abrasive particles of the instruments used for finishing; in the process of polishing, the fillers are removed in addition to the resin matrix that leave a smooth surface. However, even in this area, they are not indicated in class IV cavities, when the incisal margins or angles are involved, due to the limited resistance to incisal forces; in this case, a combination with a palatal hybrid composite portion is indicated. The addition into the organic matrix to prepolymerized resin fillers (PPRF) increases the filler content that results in reduced shrinkage. The preparation of the PPRF involves adding 50 % vol of silane-treated colloidal silica to the monomer; the composite paste is further heat cured and ground into particles. These fillers are often referred to as organic filler; however, the term composite filler is preferred, since a high level of silica is contained, to lower the viscosity. When a more extended restoration in the anterior area is required, it is indicated to combine the microfilled composite with a hybrid composite; the latter, used for the reconstruction of the palatal portion, is responsible to stand the occlusal stress, while the microfilled resin provides a polished labial surface of the restoration [2–4].

Hybrid Composite Resins In order to increase the rate of loading, a combination of inorganic particles ranging 0.04–4 is used that enables increased resistance to stress but reduces the esthetic appearance of the labial surfaces. The composition consists mainly of colloidal silica and ground particles of glasses containing heavy metals, the latest with 0.4–1,0 μm.

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They are the material of choice for anterior (even for class IV restorations) and posterior restoration, in the case of less-demanding esthetic needs. They were obtained by further grinding of the particles that formed the macrofilled composites and by the addition of submicron-sized silica. Microhybrid and nanohybrid composites are hybrid resins with the average of the particle size less than 1 μm (0.4–1 μm). They are nowadays considered the universal composites, since they combine the mechanical and esthetic characteristics, so they are used in both anterior and posterior zone, with good results. They are also recognized to have better handling and lustering properties [1–4, 6].

Nanofilled Composites Nanofilled composites are the materials with nano-sized fillers, dispersed in the organic matrix either as 5–100 nm isolate particles or as fused aggregates of primary nanoparticles, with the cluster size significantly exceeding 100 nm. They were introduced in order to improve their mechanical and esthetic characteristics (Filtek Supreme XT, 3 M Espe). Currently, nanofill and nanohybrid composites are considered the state of the art in the domain of filler formulation [2, 6, 13–15]. Studies aimed to increase the clinically relevant properties of the composite resins were conducted lately toward reducing the filler size to provide more effectively polished and with greater resistance materials and toward improving the polymeric matrix, in order to reduce polymerization shrinkage and to develop self-adhesive composites [6].

Direct Composite Resins Restorations Direct restorations represent the most commonly used restorative method, in case of dental decays, erosions, fractures, and abrasions, when reconstruction of the original dental configuration is usually aimed. It is also indicated for direct recontouring, in situations of dental shape or dimension anomalies or to close diastemas. The composite resins are frequently used materials, when direct restoration of either anterior or posterior teeth is aimed. They offer rapid, minimally invasive methods of treatment with satisfactory esthetic results [15, 16, 17]. The reasons to indicate these methods originate in their ease of use, the reduced number of treatment sessions, the increased adhesion to dental structure, and the large range of materials available on the market. In order to satisfy the requirements of a proper restoration, the composite resin should have adequate mechanical, biological, and esthetic properties, according to the position of the tooth in the dental arch. The clinical choice of a composite is based on the priority that should be given to mechanical or esthetic characteristics – if the mechanical parameters are mostly important, the material showing the highest percentage of filler is selected; in the

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case of special esthetic needs, the particle size is the factor that influences the selection [1].

Anterior Direct Composites For the restorations placed in the esthetic zone, the importance of the optical properties is predominant, since a natural aspect of the restored tooth is aimed. From an optical point of view, the goal of an esthetic restoration is to match as closely as possible the color and translucency of the natural dentition.

Dental Optical Characteristics Dental optical characteristics should be thoroughly understood, in order to be able to replicate them correctly by using dental composites. In terms of dental anatomy, young teeth, recently erupted, have a convex surface and are well contoured, with lobes, grooves, and fossae that produce a diffuse reflection of incidental light, which accounts for the “textured” aspect. Both the convexity of the surface and the texture diminish with the attrition, due to aging [18, 19]. The enamel, the dentin, and the pulp, which form the tooth crown, are different from the point of view of their structure, composition, and optical properties: dental pulp influences less the overall optical aspect; in contrast, the dentin and enamel, through their properties and thickness in each tooth area, are responsible for the optical dental characters. It is considered that, due to its dominant volume, saturation, and opacity, the dentin generates the global color of the tooth. This is more easily assessed in the cervical area, where the enamel is thin and dentin is seen through the transparency. The dentin, in shades of yellow-orange-brown, has different degrees of saturation, which increases with age and may be reduced by “whitening” procedures that induce, in fact, a “desaturation.” The chromaticity of the enamel is less important; however, it varies in shade (white-gray, white-bluish). The optical behavior of the enamel is dominated by translucency; the thicker the enamel (as in young teeth), the lighter will be the general appearance. A thinner enamel layer, as in elderly persons, reveals the dentin, which accounts for the reduced lightness. Crystallographic studies indicate the possibility of correlating the dimensions of the hydroxyapatite crystals in the enamel with its lightness [18–21]. The dentin and enamel are differently distributed in the tooth configuration; therefore, the gradient in the optical properties – color and translucency – varies along the dental crown. In conclusion, the dentin presents a chromatic translucency, with increasingly stronger saturation, with age, due to extrinsic pigmentation; the enamel is either achromatic translucent or chromatic translucent, with a grayish-white or yellowish-white aspect [22]. Due to the unlimited possibilities regarding the available shades and opacities used to reproduce the optical properties of the dental structures, the initial aspect of the direct composite restorations may be excellent [23–29].

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Indications of the Anterior Composite Resins Indications of the composite restoration on the anterior teeth are represented by: – Loss of dental structure, from various causes (dental decays, fractures, erosion, abrasion, or attrition) that affect different portions of the tooth. The dental decays are initiated mainly on the proximal surfaces, and according to the invasion into the dental structure, either a cavity class III (without the involvement of the corresponding incisal angle) or a class IV (when the angle is to be restored) is prepared (Figs. 2a, b and 3a, b). Fractures can affect only the incisal portion, or the entire crown, when a more extensive restoration is needed; however, it is aimed, often, to maintain a minimally invasive treatment method, and the buildup of a composite is the method of choice (Fig. 4a, b). Erosion can be located either on the palatal surface, when the acidic cause has an intrinsic origin, or on the labial surface – in the case of acidic food intake. Abrasion affects the incisal edge but, in the case of parafunctional traumatic forces, cervical noncarious lesions are to be treated, as well.

Fig. 2 Proximal decays. (a) Preparations without incisal angle involvement. (b) Restorations with direct composite resins (Courtesy Dr. Mihai Varvara)

Fig. 3 Proximal dental decays. (a) Preparations with the incisal angle involvement. (b) Restorations with direct composite resins (Clinical case Dr. Bogdan Culic)

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Fig. 4 Fractures of the incisal margin. (a) Initial situation. (b) Restoration in composite resins (Clinical case Dr. Bogdan Culic)

– Spaces (diastema) between adjacent teeth; – Peg shaped or other abnormality in the anatomy of the frontal teeth; more often the maxillary lateral incisor is subjected to these malformations; – Discoloration of one or more anterior teeth, due to an intrinsic cause (pulp necrosis, tetracycline-induced discoloration). In these cases, a veneering is indicated, and, when needed, an individualized reshaping of the respective tooth. As in the case of posterior restorations, the only absolute contraindications are the allergic history to constituents of the composite materials. However, caution should be taken to situations of abnormal occlusal forces (bruxism), when crown lengthening is indicated for esthetic reasons. Other relative contraindications are poor oral hygiene and increased receptivity for caries, gingival inflammation, and other conditions that limit the isolation of the operatory field.

Clinical Protocol for Direct Anterior Composites The materials of choice for anterior composites are micro- and nanocomposites and hybrid composites with micronic and submicronic fillers. However, in the case of reconstruction of the incisal margin or edge, when the fracture resistance is challenging, a combination between hybrid composite (for palatal surface) and microfilled composite (for labial surface) is recommended. For cervical lesions, flowable resins are indicated, with an elasticity module comparable with that of the dentin. In order to restore the dental structure in composite resins, tooth preparation follows the protocols indicated for class III–IV or class V cavities, modified in the respect of minimal invasive tendencies. A good isolation is required, either with the rubber dam or with retraction cord inserted into the gingival sulcus. One of the most challenging missions regarding anterior composites is to restore the optical characteristics of the dentition to mimic the adjacent dental structures. In this respect, two steps are followed: color matching and reproducing the optical

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properties of the dentition by selection and stratification of the most appropriate combination of composite resins. A. Color matching is a clinical step that can be initiated either by visual or instrumental observation. When the global shade is taken with a measurement system, the visual observation is still mandatory, in order to specify the individualization of the tooth (translucency, opaque spots or pigmentation, fissures). Instrumental assessment of the dental color can be done either with a spectrophotometer (e.g., Spectrophotometer Vita Easyshade ® in global mode or tooth area mode, depending on the extension of the future restoration); the results of the recordings, expressed in Vita Classic coding system, are used as a “starting point” that is further completed by the visual observation. Another option is a spectrophotometric device that enables the color map of the tooth (SpectroShade) or a colorimetric system. Visual selection aims to match the tabs of a shade guide with the dental area. Vita classical shade guide or the shade guides available in the composite kits (if applicable) are used. It was stated that a common mistake in color selection step is to use ceramic shade guides, since they are fabricated from materials that differ completely from composite resins and, in addition, it is difficult to find a similarity between the ceramic shade guide and the correspondent composite resin. The custom-made shade guides are the best option, with different material thickness and separation for enamel and dentin [23]. Color selection involves the analysis of each tooth area; further, these segments are assembled in order to design the color and translucency map for each tooth. The most evident instrumental errors are generated in the incisal zone, in the case of translucent areas; the spectrophotometer interprets these areas as dark-colored portions. The corrections are done by visual observation of the tooth (Fig. 5a–d). Most composite kits have their own shade guides, either universal of differentiated in terms of dentin and enamel shades. There are composite kits that have stratified shade guides, to suggest the perceived color when increasing thicknesses of composite. Some shade guides have “recipes” of combinations of the available composites, necessary to get the selected shade. When the shade is analyzed, several rules should be respected: – The lighting should be uniform; either indirect natural lighting or illuminants with color-corrected spectrum should be used. The selection is then verified in several color sources, to avoid differences in color perception, due to metamerism; the nature of incidental light plays a major role in determining the amount of light transmission or reflection. The nature of the light source influences, by this, the way in which a restoration is perceived [30]. – The color matching should be planned at the beginning of the clinical session, before the rubber dam is fixed; however, all the intense colors from the adjacent area of operatory field should be removed. – The observation of the matching pairs is done at a distance of 25–35 cm, keeping the shade tab in the proximity of the reference tooth; this can be either the tooth to be restored or adjacent or contralateral teeth.

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Fig. 5 Non-vital, discolored tooth. It was decided to preserve the remaining dental structure and to provide a direct composite, for a short term; however, the color matching is difficult in this situation. (a) Defective restoration. Selected spectrophotometric shades C4, C4, A4; (b) acid etching. Visual correction of the shade: A4, A3, A3. (c) Adhesive (Adper Single Bond 2 TM – 3M Espe). (d) Final result. Opacities used: A4D, A3.5B, A3E,TY (Filtek Ultimate- 3M Espe) (Clinical case Dr. Diana Dudea)

B. Composite selection and buildup, in order to generate a restoration in concordance with the natural adjacent teeth. In these cases, the replacement of dental tissues with materials having optical properties similar to the enamel and dentin is aimed. In order to effectively mimic the dental structures, the restorative materials must have a refraction index as close as possible to the dental tissues [19, 20]. Moreover, within the layering techniques, it is therefore important not only to choose materials with optical properties similar to natural structures but also to build them in order to gain effects of chromatic “depth” that characterize natural dentition [19]. When the composite material is selected, from the optical point of view, two factors should be considered:

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– The color parameters that are selected usually by using the Vita Classic color coding (A, B, C, or D – most often A shades are recommended) – The translucency/opacity that is influenced by the composition of the composite material, mainly by the differences in the refractive index between filler and organic matrix, the amount of filler, and the filler size [16, 31] Translucency is intermediary between transparency and opacity – in nature, the translucency of the enamel varies among teeth and is subjective perceived; however, four factors should be considered when the translucency of the composite resin is decided: presence or absence of color, thickness of the enamel, degree of translucency, and surface texture [22]. The terminology used to indicate various classes of opacities is sometimes confusing, since there are different systems used by the producers of composite materials; in general, four categories can be differentiated: – Opaque composite material (opaque dentin, opaque) is the most opaque group of materials – they have an increased opacity in comparison with the human dentin [31] that is used to block the passage of the incidental light; they are used to mask discromic substrate or to avoid the effect generated by the darkness of the oral cavity in the restorations which involve both the lingual and the labial surface (labial-lingual class III or class IV); with no dentinal opacity, even if the composite’s color is a perfect match with the remaining dental structure, the restoration will appear too dark, because a relative translucent material is not able to mask the dark background of the oral cavity [31]. – Intermediate opacity (body dentin, body) that mimics the dentin used to provide the “basic” hue and saturation of the tooth. It is the opacity of choice when a single material is used for the restoration; “body” is also used to design composites used to substitute both the dentin and enamel [33], but there are also authors who design “artificial enamel” or “body” with the meaning of a material to replace the natural enamel [32]. – Translucent material (enamel) responsible for the translucent layer that covers the dental surface enables the diffraction of the light and controls the lightness. – Transparent material (incisal, transparent) aims to create a natural, opalescent zone in the incisal area. More recently, value-changing composite resins were developed to alter the luminosity of a restoration [34]. For the small-sized restorations, when the teeth are not discolored, the blending or the chameleonic effect is considered; this effect was defined as the composite adaptability, in terms of its color, to the surrounding area. It was, however, demonstrated that the blending effect is generated by the relative translucency of the composites, which enable the influence of the optical properties of the underlying and surrounding dentin and/or the enamel on the perceived aspect of the restored tooth [35, 36].

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Fig. 6 1.2. Labial lack of substance – dysplastic condition 2.1. Dental decay – (a) initial situation. (b) 1.2. Labial filling, 2.1. distal filling. Filtek Ultimate (3M Espe). Spectrophotometric color selection: 1.2. A2, A2, A1. 2.1. A2, A2, A1. Shade used – 1.2: A2B, A1E, 2.1. A3D, A3B, A2E (Clinical case Dr. Diana Dudea)

For the intermediate-sized restorations, two layers of composite are most frequently applied: body dentin/enamel – the former aimed to replace the dentin (several colors are occasionally used, the cervical more saturated than the incisal layers) and the last added for the value and translucency control (Fig. 6a, b). In larger defects, when an increased amount of dental structure should be restored, several opacities are needed. In these cases, the histological stratification or layering technique should be used: the more opaque materials (dentin shades) are required for masking the discolored dental structure and to generate a background in the larger class III and IV restorations, in order to prevent the grayish effect due to the darkness of the oral cavity. The more translucent materials, body and enamel, are layered superficially, in order to create the natural depth from within the restoration [27, 31]. When the stratification technique is used, a good understanding of the dental morphology and structure is required, since the complexity of the optical aspect of teeth is related to their particularities of shape and structure. In these cases of advanced tooth structure loss, a direct mock-up of wax-up, followed by indirect mock-up, should be performed, in order to provide a silicon index aimed to sustaining the composite material during the buildup process (Fig. 7a–d). For the stratification technique, several opacities should be used: – Palatal layer, incisal margin, and proximal walls are built in enamel shades. – Dentin and body dentin shades replace the dentinal tissue; during buildup, it is aimed to reconstruct the configuration of the dentinal lobes, in one or several chromatic tonalities. – In case of young dentition, in the incisal third, in close proximity to the incisal edge, small increments of transparent should be added, in order to mimic the natural opalescence. – The spaces between the dentinal lobes should be filled with body dentin and, incisally, with enamel. – The final outer layer is reconstructed in enamel shades.

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Fig. 7 (a) 2.1. Fracture of the incisal third – initial case. (b) Wax-up on the model necessary for the silicon index. (c) Silicon index in position used for stratification. (d) Direct composite resin restoration – stratification (Clinical case Dr. Diana Dudea)

When a discromic dental substrate is to be masked, it is recommended to start with an opaque composite (opaque dentin), or an opaque layer, that is used to cover the dark surface; however, these layers remove the natural dental translucency that is further reproduced artificially with a dentin shade of hybrid composite and, superficially, fine veneer of microfine composite in enamel shades [1]. Translucent shades are used to mimic the natural opalescence of the incisal margin. Finally, the texture is modeled during the grinding and polishing, depending on the clinical situation, more or less pronounced, according to the adjacent teeth.

Posterior Composite Restorations When a posterior restoration is envisioned, several treatment options should be taken into account: direct restoration, using amalgam or a tooth-colored material, or indirect restorations that require teamwork – dentist-dental technician (inlays, onlays, crowns) (Fig. 8).

Indications of Direct Composite Restorations in the Posterior Area In the last decades, the amalgam fillings are, more and more, replaced by direct or by indirect composite restorations, in laboratory processed inlays, onlays, or crowns.

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Tooth colored materials Direct restorations Fillings Dental amalgams

Posterior restorations

Metalic

Ceramic Indirect restorations Crowns, inlays, onlays Composite resins

Combination (Metal-ceramic Metal-composite)

Fig. 8 Restorations of the posterior teeth – technical possibilities and materials used

The larger use of composite resins in the posterior area is due to increasingly demands for a natural appearance and to the large range of available composite colors and translucencies that allows for a natural reconstruction of the teeth, in most of the cases. Secondly, concerns regarding the toxicity of the amalgam fillings, due to their content in mercury, are reasons that explains the preferences for composites, in comparison to the amalgam restorations. However, it is stated that the concerns regarding the biocompatibility of the amalgam restorations are not based on scientific evidences [37]. The design of the required cavity is also beneficial, in respect to the conservation of the dental structure: in the case of the composites, due to the chemical bonding, the undercuts are not needed, so the preparation minimally invades the sound dental structure. Currently, the design of preparation is limited to the removal of the tooth structure that involves carious dentin and fragile [38, 39]. The preparation is narrower; by this, in addition to the minimally invasive principle, the area of restorative material exposed to occlusal forces is reduced. A limited quantity of composite is desirable also in order to prevent the polymerization contraction and to limit the marginal leakage at the dentin-restoration interface. The composite materials are clinical solutions for patients experiencing thermal sensitivity generated by amalgam fillings.

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In comparison with the indirect restorations, posterior direct composites are provided by a limited number of sessions – the treatment requires a single, in-office session; the costs are reduced since the dental laboratory is not involved. The benefits of the adhesive restorations over the classical-cemented crowns, inlays, or onlays should be mentioned – not only the quality of the cementation but also the reinforcement of the remaining weakened dental structure is demonstrated in the case of chemical-bonded material. However, in spite of their excellent imitation of the dental structures, both from an optical and from mechanical point of view, their technique sensitivity during the insertion, modeling, and polymerization is recognized. A durable composite filling requires a good isolation of the operatory field; moreover, the clinical protocol, starting with etching, bonding, insertion into the cavity, modeling, curing, and polishing – all technical steps – should be followed according to the respective instructions. In practice, the posterior direct composite restorations are indicated as sealants and reconstruction materials of premolars and molars, with loss of substance in dental decays, fractures, abrasions or attritions, and erosions. Due to the performance of the currently available composites, there are no absolute contraindications of the direct composite restorations, with the exception of allergic history to materials’ components. However, several clinical situations should be regarded as relative contraindications and the treatment plan decided after a careful assessment of the general and local conditions: – Lack of oral hygiene and limitations in the patients’ ability to improve the status of the dental cleaning; – Gingival bleeding – considering the importance of a clean operatory field and of a good isolation, the composite filling placement should be planned after a rigorous professional cleaning, to avoid gingival inflammation due to bacterial deposits; – Difficult clinical accessibility, when the possibility of isolation is limited (distal cavities, malpositioned teeth, reduced opening of the oral cavity, difficult communication with the patient); – Large defects with extended enamel loss due to the reduced surface responsible for a durable bonding; – Defects extended below the gingival margin, unless a gingivectomy is performed, in order to place the gingival limit of the restoration above the free gingiva. If the cervical limit is located on the cementum or in the next proximity of the gingival margin, an open-sandwich restoration, with the gingival portion represented by an increment of resin-modified glass ionomer, could be the treatment of choice; – In the case of teeth subjected to traumatic forces – premature contacts and bruxism – the failures, in these cases, involve not only the composite resins (fracture, attrition, or debonding) but also the remaining dental structure that can be subjected to cracks or fractures; – Lack of substance of the occlusal surface that involves the reconstruction of one or more cusps (mainly supportive cusps), due to the large amount of force that is transmitted in this area.

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Clinical Steps for Direct Posterior Composite Restorations In the posterior area, depending on the extension of the substance loss, various clinical protocols of preparation and application could be followed: Dental sealants – basically, the procedure consists in the blocking of the shallow depressions on the dental surface with composite resins, in order to prevent the dental plaque accumulation and the consecutive demineralization of the dental structures; it is indicated on both primary and permanent posterior teeth, at a moment that follows closely the eruptive stage. Among the composites, the materials of choice are flowable composite resins, but also other groups of materials could be used: glass ionomers and resin-modified glass ionomers. The clinical protocol is initiated by the dental surface cleaning and, according to the type of resin, is followed by the conditioning of the area (acid etching and bonding). In the case of self-adhesive composites, the preliminary etching and bonding step is not required. The most important clinical limit of this procedure is the sealants’ debonding due to lack of retention; patients should be included in preventive programs and followed on a regular basis. Preventive restorations consist in minimally invasive cavities that with a limited extension, at the carious fissures and pits. The preparation is in contradiction with the classical principles of preventive extension and retention of cavities, and it is recommended due to the development of adhesive dentistry (dental surface etching, composite bonding). According to the cavity extension, either flowable composite resin (when only shallow fissures or pits are involved) or a combination is used – conventional composite to restore the occlusal morphology and flowable to fill the fissures and pits and to cover the entire restoration (when a more extended cavity is needed). Conventional restorations extended on one or more of the dental surfaces: Class I cavities, for the occlusal surface, is limited more often to preventive restorations. In the case of deep decay, a pulpal protection, with a liner (glass ionomer), is indicated, followed by the bonding procedures; the restoration involves the placement and model of the occlusal morphology using the incremental technique. Class II cavities are extended on a single or both proximal surfaces and on the occlusal surface; as in class I, the use of a liner is limited to the cases with close proximity of the dental pulp. In these cases, the restoration of the proximal walls requires a matrix band; this should be placed in close contact with the adjacent tooth, in order to obtain the contact area. The matrix is blocked with a wedge that allows to create proper space for the dental papilla. In the light of preventive dentistry, a minimally invasive approach is indicated in this location, as well: each proximal and occlusal cavity should be prepared and restored separately [38] (Fig. 9a–f). In deep cavities, or when the exposure of the dental pulp is suspected, a calcium hydroxide-based liner is indicated. Another option is the glass ionomer liners that have the benefit of bonding by the subjacent dental tissue and by the overlaying composite.

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Fig. 9 3.6, 3.7. Occlusal direct restorations. (a) Defective restorations. (b) Occlusal boxes, protective liner on the cavity floor. (c) Acid etching. (d) Application of the adhesive system. (e) Buildup of dental morphology, by using incremental technique. (f) Composite restorations. Occlusal contacts are verified with articulating paper (Clinical case Dr. Florin Alb)

Various formulas and corresponding protocols could be followed further in order to apply the bonding system: self-etching, total etching, in one, two, or three steps. Traditionally, the composites indicated for the posterior area are highly viscous materials that allow for a better condensation and modeling of the occlusal details and of the proximal contours. An option was the packable and condensable composites, based on elongated, fibrous fillers, with a textured surface [4], that resist flow and enable a better condensation that mimics the insertion of the amalgam fillings. However, viscous materials are difficult to adapt to the cavity walls – thicker consistency has significantly more voids than the medium or thinner consistency [38, 40]. It is recommended to inject the material by using composite capsules or placemen tips that allow a better adaptation and a more homogenous composite mass in comparison to the classical method of insertion with conventional spatulas; another method of decreasing the viscosity is to use heated composite to 60–68  C. This method facilitates the insertion of the composite into the cavity, improves the adaptation to the cavity walls, and reduces the voids into the composite mass [38, 40].

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In order to improve the marginal adaptation, a first increment of flowable resin could be used on the gingival and pulpal floor and for the proximal cavity. This improves the marginal adaptation, since the flowable consistency allows for the penetration in all the cavity details and also reduces total polymerization shrinkage. Another advantage is the easier handling of the next increments of viscous material that adheres easier to the cured flowable. The conventional hybrid composites are most often the materials of choice for posterior area. The insertion of the composite into the cavity is, more often, based on the incremental or layering technique. According to this technique, the composites are inserted into the cavity in layers that do not exceed 2 mm in thickness, in order to both ensure a complete cure of the material by the curing light and control polymerization shrinkage [6]. The layers are positioned in an oblique pattern that means to avoid the connection of two opposite walls by the same increment. This prevents cuspal flexure due to composite shrinkage and subsequent fissures or fractures that are more frequent in the horizontal layering technique. The incremental technique influences the ratio of bonded to unbonded restoration areas – the “C-factor”– that should be decreased, in order to lower the total bond strength of the composite filling [38]. The bulk placement technique is advocated lately, due to the newer class of composites with reduced polymerization shrinkage. Whatever the technique used, the final increments are placed according to the occlusal morphology. The objective of this procedure is not only to fill the cavity and to ensure a good marginal adaptation but also to rebuild the occlusal morphology: cusps, ridges, grooves, and fossae. At this point, the occlusal contacts are verified – it is intended to keep them outside the enamel-composite interface. For good occlusal adjustment, contacts are verified in maximum intercuspation and protrusive and laterotrusive movements; finishing and polishing are required to prevent plaque accumulation and abrasion of the opposite occlusal surface.

Clinical Consideration Regarding the Composite Properties When a composite restoration is envisioned, two main objectives should be considered: – To provide a long-term resistant substitute for the dental structure, by using a material that is comparable, in terms of the mechanical properties, with the tooth itself; – To enable a natural look of the restored tooth, in other words, to use a material or a combination of composite resins that mimic the optical properties of the tooth; The decision regarding the composite used in a certain clinical situation is based on the location of the tooth to be restored (anterior vs. posterior) and the dimension of the future restoration [1].

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In the case of large restoration, in the posterior area, either an indirect composite or a direct restoration with a radiopaque composite, with a high inorganic load is needed [1, 41]. The anterior composites should provide the highest polished surface, fluorescence, and the proper level of translucency; microfilled or nanofilled resins will be indicated in this case, either by themselves or in combination with a stronger, more flexural, and fracture-resistant composite, for the palatal surface, as was described previously [1]. In these cases, the stratification will take into account not only the optical characteristics but also the mechanical resistance. The properties of the direct composite resins are influenced by their composition and the polymerization process. Among the constituents, the inorganic filler and the coupling agents are responsible for the composite’s hardness, flexural resistance, translucency, and coefficient of thermal expansion, while the chromatic stability and the shrinkage are influenced by the organic matrix [16, 23, 24]. The factors that are directly correlated with the properties and, consequently, with the clinical applications are the average size of the filler, filler volume level, size distribution, index of refraction, radiopacity, and hardness of the filler [4].

Handling Characteristics Handling characteristics are regarded in respect to the composite viscosity and the tendency to stick to the instrument. Flowable material provides the easiest adaptation to the cavity walls. However, there is the risk of air trapping and the void formation, within these materials. Conventional softer materials adhere well to the margins and to the cavity walls; however, they could stick to the instruments and are difficult modeled. More viscous materials are less sticky, but they are more difficult to be inserted into the cavity and have to be modeled with special designed instruments. Optical Properties The optical properties of the composite resins are represented by the composite color (shade) and translucency/opacity; they should mimic the respective optical properties of the dental structures, in order to gain a natural appearance. Opalescence and fluorescence are characteristics that influence, in addition, the optical appearance of a restoration.

Color The available commercially composites are manufactured in a wide range of opacities and shades. They are designated most often according to the Vita Classic shade guide codification system. Among the four shades (A, B, C, and D), the most used are the A groups, since it is recognized that the majority of human teeth have A shades. For this reason, there are composite sets that include only A shades, in various saturations. According to the position into the dental arch, one or more shades can be used in order to reproduce the natural aspect of the tooth, and the map of the tooth is built.

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Translucency Translucency is the ability of a material to allow light to pass through and thus to allow the appearance of the underlying background to show through [16, 42, 43]. Translucency is influenced by the composition of the material, mainly by the differences in the refractive index between the filler and organic matrix, the amount of filler, and the filler size [16, 42]. For acceptable esthetics, the translucency of a composite restoration needs to be similar with the optical properties of the dental structure. The index of refraction of the filler needs to closely match that of the resin (around 1.50) [4]. A multitude of translucencies are provided in the composite kits, in order to offer support to mimic the natural aspect of the dentition. The corresponding terminology was discussed previously. However, when similar colors and translucencies of composite resins originating from different brands were evaluated, poor color compatibility of pairs of identical shade designation was observed. The best color match was recorded for A2 shade pairs, followed by C2 pairs, B2 pairs, and opaque A2 shade pairs [44]. In order to characterize the composite translucency, different parameters are in use: translucency parameter (TP), contrast ratio (CR), and relative translucency (T). The translucency parameter of a material refers to the difference in color between a uniform thickness of the material over a white background and the same thickness of the material over a black background and provides a value corresponding to the common visual perception of translucency [42]. Another method used to measure translucency is by calculating contrast ratio, as the ratio of illuminance (Y) of a certain material over a black background (Yb) to the illuminance of that material over a white background. Another important optical measure is the refractive index (the relationship between the speed of the light in vacuum and in a certain material); the refractive index of a composite material should be as close as possible to the refractive index of the dental structures (e.g., enamel) [22]. The specific factors that should be considered when a certain composite material is selected in respect to the translucency are the translucency of the material, itself, the thickness of the dental layer to be replaced, the chromaticity that should be imitated, and the surface texture of the labial surface. The thickness of composites is also important in the perception of the final esthetic outcome. When the thickness of translucent chromatic composite increases, the chroma increases and value decreases. The level of polish influences the perception of a shade. The more polished a surface, the more light is transmitted and less reflected – the value decreases as a consequence [22, 45]. Light curing creates changes of the translucency/opacity of a composite. Saturation, value, and translucency may decrease or increase after light-curing, and the variation is related to the nature of the composite. In microfilled composites, the translucency and chroma decrease after polymerization, while in hybrid composite resins, the same parameters increase as a result of curing [22]. Composite resins should have fluorescence similar to the natural teeth that emit blue fluorescence when exposed to ultraviolet light. Fluorescence, although

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minimally perceptible under normal light, becomes important when the tooth is viewed under UV illumination; due to fluorescence, the teeth and corresponding restorations are brighter in the natural daylight; however, external staining and sealant application on the composites’ surfaces decrease the fluorescence [46].

Radiopacity Radiopacity is required mostly in the case of posterior fillings, since here the secondary caries are more difficultly detected by clinical means; it is believed that the optimal radiopacity of a dental composite should be comparable to that of the enamel; an increased radiopacity can act as a barrier that masks the radiotranslucency of a marginal gap [4]. Polymerization Contraction Polymerization contraction ranges between 1.5 % and 5 % – as a result, gap formation, marginal leakage, cusp reflection, and corresponding fissures of the cavity walls are among the most undesirable events that affect the composite longevity. The polymerization contraction occurs toward the cavity wall that proves to be more bonding resistant; this is generally represented by enamel surfaces. Current research is oriented both to find new formulas of monomers with a lower contraction rate and to adopt technologies that provide less contraction. In order to reduce polymerization shrinkage, incremental technique, with the addition and polymerization of reduced portions of material at a time, is recommended. In addition, two-step polymerization, with increasing irradiance of the curing regiments, decreases the contraction rate [38, 47].

Clinical Problems Generated by Direct Composites Marginal Leakage Marginal leakage is due to polymerization shrinkage and to the quality of bonding between the composite and the dental substrate (enamel or dentin). A common situation is that of the proximal or cervical fillings, where the gingival wall is represented by dentin or cement and the other walls of the cavity by enamel; in this case, the material tends to pull away from the dentin or cement during curing, due to polymerization shrinkage, with gap formation [4].The formation of the marginal gap is followed by the progression of the cariogenic bacteria inside the space and, consequently, by the development of secondary caries. Postoperative Sensitivity Postoperative sensitivity is correlated with the flexure of opposing cavity walls, due to polymerization shrinkage, or with the application of the composites in deep cavities, without pulpal protection; in order to prevent the second shortcoming, a glass ionomer, resin-modified or flowable composite liner is used, in addition to the composite restoration, in such clinical situations [7]. Cuspal flexure is avoided by using oblique increments during stratification, so that every layer contacts directly a

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single wall. Another explanation for the postoperative sensitive is gap formation between the filling and the cavity wall; either as a result of thermal, chemical, or bacterial stimuli, the fluid within the dental tubules migrates and stimulates the odontoblastic processes that result in pain sensation [38, 48]. It is shown that postoperative sensitivity increases for larger restorations [38, 49].

Marginal Staining Marginal staining occurs as a result of microleakage, due to composite contraction, but also as a result of inappropriate bonding – when the excessive layers of adhesive will change color and form a dark zone around the restoration. Fracture Due to Inadequate Mechanical Properties Fracture may be due to inadequate mechanical properties of the composite, incorrect indication of the particular area of placement (anterior or posterior), but also as a result of poor design and preparation of the cavity, with limitations in the quantity and quality of the adjacent remaining dental tissues [6]. In order to prevent adhesive or cohesive fractures, large restorations should be avoided, cusp coverage is also indicated. For patients with bruxism, other methods of treatment are suggested. Microcracks Microcracks at the surface or in the volume of restoration can be correlated with the water sorption – water sorption is a function of the resin content of the material and the strength of the resin-filler interface. Extreme water sorption causes the expansion and plasticizing of the resin, which leads to reduced longevity of the composite resin and hydrolysis of saline, which in turn creates microcracks [16, 27, 50, 51, 52]. Wear of the Composite Surface Wear of the composite surface is dependent on the intrinsic mechanic composites but also on the occlusal contact with adjacent teeth, the material of the opposite restorations, and the position of the tooth in the dental arch. Posterior teeth, being subject of increased occlusal forces, are more exposed to attrition. It is considered that the resin composite wear occurs as a result of the combination between the chemical and the mechanical degradation of the restorative material. The mechanical loading is due to either the attrition or abrasive factors. In the first instance, the cause is the contact with the opposite dental surface [38]. In this respect, the quality of the occlusal surface of the opposite teeth or restoration is paramount. Ceramic materials, particularly when inappropriate glazed, induce the most important wear; loss of occlusal morphology, due to cracks followed by fracture and loss of composite mass, is experienced. The consecutive egression of the opposite teeth is expected in these clinical situations. In the case of abrasion, the food particles in mastication are responsible for the organic matrix wearing and for the exposure, as a result, of the inorganic filling. The average rates for posterior composite resin are between 7–12 μm/year and 0.1–0.2 mm more than enamel over 10 years [4, 37].

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The differences in the origin of the traumatic forces may explain the behavior of the composite resins. Among the groups of composites, the microfilled have the most critical rate, due to the lowest filler concentration (30–50 %) to the attrition, but they are more resistant to abrasion since they have finer particles and reduced interparticle spaces [38, 39]. Unfortunately, the occlusal wear occurs under normal loading, when the composite restoration is in contact with intact enamel (0.1–0.2 mm/10 years, more than the enamel) [4]. The indication of direct posterior composites is questionable for the clinical situation of bruxism, due to the higher occlusal stress.

Water Sorption Water sorption is due to the organic matrix of the composite; it is responsible for the composite restorations’ increase in volume. The internal structure may be compromised, due to reduction of bond strength [2]. Staining One of the most important shortcomings of the composite resins is their discoloration over time; the color changing can be caused by internal or external factors. Internally induced discolorations are permanent and are related to polymer quality, type, and amount of the inorganic filler, as well as the synergist added to the photo initiator system. Components like unconverted camphorquinone and tertiary aromatic or aliphatic amines are responsible for a yellowish or brownish discoloration [16, 43, 53–55]. The intrinsic discoloration can be assessed by artificial accelerated aging tests. It was demonstrated that these methods are suggestive when the characteristics of the composition that interfere the optical properties are assessed. The extrinsic discoloration is due to exposure to food colorants, UV radiation, temperature change, and water absorption [16, 23–29]. In order to assess the extrinsic discoloration, various accelerated staining protocols have been developed. Staining solutions and immersion time are significant factors that affect color stability of composite resins. The samples of composite materials are immersed in various natural colorants (tea, coffee, red wine, fruit extracts or beverages, chocolate) and artificial staining solutions [16]. There are studies that show that, in spite of their excellent esthetic appearance, nanocomposites undergo a greater color change than microhybrid composites after staining [16, 53, 56].

Indirect Composite Restorations in Dentistry Our patients require today highly esthetic solutions also in the posterior teeth, that are at the same time functional and long-lasting. While direct composite restorations have adequate properties to be used in most areas of the mouth, concern still exists when these materials are placed in large preparations, on several teeth in a quadrant, and mostly when used to replace functional cusps, in patients with bruxism or parafunctional habits [6]. The concerns are related to the fracture of the restoration,

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its wear – and also the ability to restore a correct contact area in class II direct restorations. Microleakage at the interface composite filling – tooth structure with the risk for secondary caries – represents today the most frequent cause for direct composite replacement. If the gingival floor of the proximal box is placed subgingivally, the glass ionomer (GI) cements or resin-modified glass ionomer (RMGI) cements used to cement an inlay will provide better adhesion then a direct composite [57]. Esthetic inlays and onlays – made in composite and ceramics – offer a viable alternative for patients who desire a more esthetic option than gold, but a more functional alternative than direct composite fillings. While ceramic inlays/onlays provide the patients with a lot of benefits – high esthetics and stability over time – composite resins present some major advantages, including repair capabilities, resilience for comfort and shock absorption, adjustability in the mouth, less chance for differential wear at the luting agent-restoration interface, and no wear of opposing structures in functional contact [58] (Fig. 10a–d).

Advantages and Disadvantages of Laboratory Processed Inlays The advantages of laboratory-fabricated composite restorations over the direct composite restorations are: – Restoration of the contact area at the right position and shape with the use of removable dyes, giving the technician the ability to reproduce the complex convex-concave morphology of proximal surfaces. – Better reproduction of the occlusal morphology in the lab with the use of an articulator – while direct techniques the patient is under rubber dam isolation so the dentist cannot use the opposing teeth to check occlusal contact only when the restoration is finished; Both proximal and occlusal details can be achieved especially with the use of CAD-CAM technology in composite inlays (Vita Enamic, Vita or Lava Ultimate, 3M). – Excellent esthetic results that can be achieved with shade matching, a better color stability over time because of superior polymerization obtained in the laboratory; – Excellent mechanical properties similar to those of the natural tooth (the flexural strength and the elasticity modulus are close to those of the dentin); – Improved resistance to wear, because of the composition and structure of these materials that have a higher filler content than the in office composites; – Ability to restore the vertical dimension of occlusion (VDO); – Less marginal leakage than the direct techniques. With the lab composite, polymerization shrinkage took place in the lab stage and any marginal gap of the inlay will be filled with a film layer of dual-cure composite or GI cements; – Ability to be finished and polished in the laboratory easier and better than in the mouth;

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Fig. 10 (a–d) Four types of esthetic inlays: (a) composite inlays; (b) feldspathic inlays; (c) pressed ceramic eMax inlays; (d) CAD-CAM-produced ceramic inlays

– The triple polymerization performed in the lab, under light, heat, and pressure, will generate less unreacted monomers so better mechanical, biological, and esthetic properties overall; – Reduced chair time in cases with more inlays on the same quadrant; the preparation and impression are faster than the direct technique [18, 59].

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Disadvantages would include: – Inlay preparation and cementation is a sensitive technique that is very much dependent on the skills of the dentist, but this is also true for direct composites; – Longer processing time and additional costs.

Technological Alternatives for Indirectly Processed Composite Restorations There are two methods available for the fabrication of indirectly processed restorations: conventional and CAD-CAM methods.

Conventional Technique The conventional technique includes additional laboratory cost and a second appointment needed for inlay cementation as compared to the direct restorations. These can be successfully eliminated by the use of CAD-CAM in-office technology that is able to produce both ceramic and composite inlays, onlays, and crowns, using systems such as CEREC (Sirona), E4D Dentist (E4D Technologies), and Lava system (3M). Results are impressive, because certain clinical steps are eliminated, including the conventional impression, long-distance transportation to the dental lab, and additional clinical appointments. As the inlay can be completed in several hours, the patient no longer needs to be recalled a second time in the office; the cementation can be performed 3–4 h later, which means a lot of time saved for the dentist and the patient. The precision of these inlays is extremely high, and the materials from which they are milled, have optimal properties [60, 61] (Fig. 11a–d).

Fig. 11 (a–d) Clinical case (a) initial aspect of teeth; (b) rubber dam in place before preparation; (c) preparation; (d) ready for PVS impression (Clinical case Dr. Florin Alb)

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The clinical case of a 28-year-old patient TC who presented after orthodontic treatment to replace old restorations with secondary caries on teeth 4.4, 4.5, 4.6, 4.7 clearly illustrates the technique and pro’s and con’s for these indirect restorations. Teeth 4.5 and 4.6 are non-vital teeth endodontically treated and restored with fiber post and core (4.6). Due to subgingival extension of the cavities, all four being on one quadrant, the decision was made to apply four composite inlays in one single arch impression. Preparation was performed under the rubber dam isolation. After the preparation was finished, the rubber dam was removed for the impression (polivinylsiloxane impression material in a double-mix impression technique) The inlay was fabricated in the lab using Vita LC composite (Vita), following the Vita 3D-Master shade guide, which allows the dental technician to use a logical approach in layering the desired shades. Shade matching was done before the rubber dam was placed and also after the impression was taken for the core color of the inlay using the Vita Dentin shade guide. A cast is created in the lab and it is mounted in an articulator. Next, any slight undercuts were filled with wax (Classic Opaque Sculpting Wax, Renfert) on the model. The first layer of indirect composite (Vita LC, Vita) base dentin shade was placed and cured. Then, the tooth was built up incrementally with dentin and enamel shades layered to create a more lifelike appearance. The inlays and onlays were placed in the Ivoclar curing unit (under light and heat) for the recommended time limit. Once the restorations were cured, they were finished and polished, then protected with a layer of wax before micro-sandblasting the inner surfaces. Some practitioners also apply on the inner surface a composite-activating primer in order to reactivate the surface of the heavily polymerized composite (over 95 % degree of monomer conversion). Before cementation, the rubber dam was applied. The internal surfaces of the preparations were cleaned with orthophosphoric acid 37.5 % for 10 s, rinsed, and dried. The composite inlays/onlays were tried in for fit – this being a major advantage for the composite restorations as compared to the ceramic ones, they can be easily fitted and adjusted and then refinished at the chair side. The RMGI cement (G-Cem, GC) was applied in the cavity and on the inner surface of the inlay. The composite restorations are applied using low pressure and ultrasounds with a silicone insert. Pre-curing for 1 s, about 1 cm away from the margins, is very beneficial as it allows for easy removal of the excess. Other materials frequently used for inlay cementation are the self-adhesive dual-cure composite cements like Panavia, Kuraray, and Rely-X, 3M. Composite inlays on teeth 4.5 and 4.7 show very good fit after cementation (Fig. 12). There is also a very conservative “trend” in dentistry nowadays, so dentists are also looking for an alternative to crowns and bridge retainers, frequently favoring an inlay or an onlay over a full-crown preparation. During the last 5 years, in the United States alone, over 200 million crowns were placed and many of those crowned teeth could have benefited from a more conservative treatment [62]. Recent literature studies indicate a 10-year success rate of composite and ceramic inlays above 90 %, which is close to the gold standard – the noble alloy inlays that are among the most successful and long-lasting indirect restorations [63, 64].

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Fig. 12 (a–d) Cementation of the composite inlays. (a) Rubber dam in place for adhesively luting the indirect composite restorations; (b) inlays cemented on teeth 4.7 and 4.5; (c) inlays cemented on all teeth; (d) final aspect of composite inlays in the mouth (Clinical case Dr. Florin Alb)

An important step in the use of indirect composites for larger restorations than inlays and onlays was the discovery and applications of reinforcing fibers – dating back to the 1960s, it was first proposed for reinforcing denture base acrylic and is now most commonly used in prefabricated endodontic fiber posts. With the fiberreinforced composite (FRC), we are now able to fabricate orthodontic retainers, periodontal splints, and fiber-reinforced bridges as short- to medium-term provisional and temporization during the healing phase of implant treatment [65]. They offer some advantages over PFM technology: no preparation or limited preparation required, lower cost, less chairside time, and good medium-term survival (75 % at 4 years) [66, 67]. The fiber-reinforced composite bridges (FRCB) can be fabricated directly in the patient’s mouth or by indirect technique (Fig. 13). Composite resins used in indirect techniques have a wide variety of applications in dentistry today: from inlays, onlays, overlays, and table tops, to provisional crowns and bridges, adhesive bridges, to extended splinting systems fabricated in the lab for patients with periodontal disease, to orthodontic devices used for short period of time, and to long-term resin crowns and bridges. The applications of composite resins – in direct and indirect procedures, both in the dental office and in the dental laboratory – have increased exponentially during the last 60 years since their introduction in the 1950s, and it is most likely that their uses will continue to grow both in frequency and application due to their versatility [65].

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Fig. 13 (a, b) Fiber-reinforced composite bridges (FRCB). (a) Initial case, missing lateral incisors. (b) FRCB was fabricated in the lab using Construct fibers (Kerr) and Vita VM LC composite (Vita) (provisional treatment for esthetic rehabilitation during the osteointegration phase) (Clinical case Dr. Florin Alb)

CAD/CAM Processed Composites CAD/CAM Systems In the last years, CAD/CAM technologies (computer-aided design/computer-aided manufacturing) became an important component of dental medicine. The technique is orientated to the future and focused on high efficiency and standardization of prosthodontics treatment [68]. The advantages in CAD/CAM technologies consist in the quality of the materials, industrial processing, reproducibility, and low cost/unit. CAD/CAM systems available are “in-office” systems, or laboratory systems. All CAD/CAM systems have three main components (Fig. 14):

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Fig. 14 (a) CEREC Omnicam (Sirona, Germany). (b) Restoration design. (c) Final restoration (Vita Enamic)

• A scanner – for image acquisition, transforming images into data • An image software – for data processing, restoration 3D design • A milling unit – for transforming the virtual design into a finite product (inlays, onlays, crowns, bridges) For direct clinical application, intraoral scanners were developed (Fig. 14a). Acquiring optical images of the prepared teeth directly with the intraoral camera eliminates the need for conventional impression procedures and improves patient comfort [69, 70]. With the aid of the chairside software, the restoration is designed (Fig. 14b) and then milled in the office, within a single-visit restoration (Fig. 14c). Single-visit restoration eliminates the need for provisional restorations, increases durability of adhesion to dental tissues, and also reduces postoperative sensitivity.

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Advantages of in-office CAD/CAM systems: • The possibility to obtaining tooth crowns, inlays, onlays, veneers, or bridges in a single visit; • Eliminates the temporary restorations; • Esthetic restorations are obtained; • When inlays or onlays are fabricated, the longevity of the restorations is higher, when compared to direct composites. Disadvantages • For a ceramic restoration glazing has to be performed; for composite, external staining is necessary in order to obtain esthetic restorations; • CAD/CAM system cost [71]. The major concerns about chairside CAD-CAM restorations are the accuracy of intraoral digital impressions and the resulting internal and marginal fit discrepancies. However, recent studies demonstrated that digital impression systems allow the fabrication of fixed prosthetic restorations with similar accuracy as conventional impression methods. According to the results of previous studies, marginal gap of the feldspathic crowns fabricated with CEREC Omnicam (Sirona, Germany) system ranged from 18 to 40 μm and is in clinically acceptable limits [72]. Composite Materials for CAD/CAM Technology Innovative materials were developed in order to fulfill researchers, clinicians, and patient demands. A great variety of dental ceramics and a large selection of composite resin materials are nowadays on the market. Advantages of ceramics are a high flexural strength and great color stability, while disadvantages are high antagonistic tooth wear and loss of tooth structure due to a minimum thickness of 1.5–2.0 mm. These two parameters are better for composite resins, but the wear of the material itself is higher. While ceramics are stiffer and harder than natural tooth structure, composite resins show lower values [73]. For many years, composites were used for provisional crowns and bridges (Vita CAD-TEMP, Vita, Germany; Telio CAD, Ivoclar Vivadent, Lichtenstein) (Figs. 15 and 16) with a maximum use of 12 months. VITA CAD-Temp composite blocks consist of a fiber-free, homogeneous, high-molecular, and cross-linked acrylate polymer with microfiller. CAD-Temp is used for the fabrication of multi-unit, temporary bridge restorations with up to two pontics and temporary crowns. Telio CAD consists of acrylate polymer (PMMA) blocks for the fabrication of long-term temporaries using CAD/CAM technology. Recent materials like Lava Ultimate (3M Espe, USA) (Fig. 16) and Vita Enamic (Vita, Germany) represent high improvements in CAD/CAM dentistry material science. Material indications are for permanent restorations in the anterior and posterior areas, veneers, inlay, and onlays. The options for

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Fig. 15 Vita CAD-TEMP

Fig. 16 Telio CAD

Fig. 17 Lava Ultimate

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composite block color are large, in order to cover different clinically situations, monochromatic and polychromatic blocks (with three different shades in one block) being available [74]. Lava Ultimate is a resin nanoceramic material which blends approximately 80 % nanoceramic particles with a highly cured resin matrix. With a flexural strength of 200 MPa, the material is less brittle than glass ceramic; it resists chipping and cracking when milled (Fig. 17). It enables a simple, efficient, no-firing process for making CAD/CAM restorations. Staining and manual polishing is required. Vita Enamic combines the properties of ceramic and polymer. It consists of a hybrid structure with two interpenetrating networks of ceramic and polymer, a so-called double-network hybrid. Due to the fine structure of feldspar ceramic and the acrylate polymer network, this material has a similar abrasion, high flexural strength, and elasticity close to dentin. The Vickers hardness was evaluated with values between those for dentin and enamel. Indications are inlays, onlays, veneers, and crowns on the posterior area. No-firing process is required [75]. Using light-cured staining materials, external characterizations can be made for maximum esthetics (Fig. 18a–f).

Fig. 18 (a) Vita Enamic Block. (b) Initial situation. (c) Tooth preparation inlays on 26, 24, and onlay 25. (d) Virtual model and restoration design. (e) Milled restorations on site. (f) Stained and polished restorations – final result (Clinical case dr Bogdan Culic)

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50. Imazato S, Tarumi H, Kobayashi K, Hiraguri H, Oda K, Tsuchitani Y (1995) Relationship between the degree of conversion and internal discoloration of light-activated composites. Dent Mater 14(1):23–30 51. Ferracane JL, Berge XH, Condor JR (1998) In vitro aging of dental composites in water effect the degree of conversion, filler volume and filler/matrix coupling. J Biomed Mater Res 42:465–472 52. Bayne SC, Taylor DF, Heymann HO (1992) Protection hypothesis for composite wear. Dent Mater 8:305–309 53. Al Kheraif AA, Qasim SS, Ramakrishnaiah R, Rehman I (2013) Effect of different beverages on the color stability and degree of conversion of nano and microhybrid composites. Dent Mater J 32(2):326–331 54. Ren YF, Feng L, Serban D, Malmstrom HS (2012) Effects of common beverage colorants on color stability of dental composite resins: the utility of thermocycling stain challenge model in vitro. J Dent 40(Suppl 1):e48–e56 55. Lee BS, Huang SH, Chiang YC, Chien YS, Mou CY, Lin CP (2008) Development of in vitro tooth staining model and usage of catalysts to elevate the effectiveness of tooth bleaching. Dent Mater 24:57–66 56. Yazici AR, Celik C, Dayangac B, Ozgunaltay G (2007) The effect of curing units and staining solutions on the color stability of resin composites. Oper Dent 32:616–622 57. One-visit biomimetic composite resin inlays/onlays. www.dentistrytoday.com. June 2010 58. Kois JC (1996) The restorative periodontal interface: biological parameters. Periodontology 2000(11):29–38 59. Roberson TM, Heymann HO, Swift EJ (2002) Sturdevant’s art and science of operative dentistry, 4th edn. Mosby, St Louis 60. Schlichting LH, Maia HP, Baratieri LN, Magne P (2011) Novel-design ultra-thin CAD/CAM composite resin and ceramic occlusal veneers for the treatment of severe dental erosion. J Prosthet Dent 105:217–226 61. Rekow ED (2006) Dental CAD/CAM systems – a 20-year success story. J Am Dent Assoc 137:5s–6s 62. Christensen GJ (2008) Considering tooth-colored inlays and onlays versus crowns. J Am Dent Assoc 139:617–620 63. Thordrup M, Isidor F, Hörsted-Bindslev P (2006) A prospective clinical study of indirect and direct composite and ceramic inlays: ten-year results. Quintessence Int 37(2):139–144 64. Peumans M, Voet M, De Munck J, Van Landuyt K, Van Ende A, Van Meerbeek B (2013) Fouryear clinical evaluation of a self-adhesive luting agent for ceramic inlays. Clin Oral Invest 17:739–750 65. Belvedere P, Turner WE (2002) Direct fiber-reinforcement composite bridge. Dent Today 21:88–94 66. Van Heumen CM, van Dijken WV, Tanner J, Pikaar R, Lassila VJ (2009) 5 year survival of 3 unit fiber-reinforced composite fixed partial denture in the anterior area. Dent Mater 25:820–827 67. Khatavkar RA, Hegde VS (2010) A conservative treatment option for a single missing premolar using a partial veneered restoration with the SR Adoro system. J Conserv Dent 13(2):102–105. doi:10.4103/0972-0707.66722 68. Beuer F, Schweiger J, Edelhoff D (2008) Digital dentistry: an overview of recent developments for CAD/CAM generated restorations. Br Dent J 204(9):505–511 69. Birnbaum NS, Aaronson HB (2008) Dental impressions using 3D digital scanners: virtual becomes reality. Compedium 29(8):494–505 70. Feuerstein P, Puri S (2009) An overview of CAD/CAM and digital impressions. Oral Health 99 (9):65–71 71. Davidowitz G, Kotick PG (2011) The use of CAD/CAM in dentistry. Dent Clin N Am 55 (3):559–570

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Clinical Evaluation of Disilicate and Zirconium in Dentistry

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Domenico Baldi, Jacopo Colombo, and Uli Hauschild

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Zirconium . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Criteria . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Laboratory Techniques . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Kinds of Operations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Crowns and Bridges . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Abutments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Complex Direct Screwing Cases . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Lithium Disilicate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Crowns and Inlays in Lithium Disilicate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Cases with Mixed Use Between the Two Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Case . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

In recent years there has been a huge leap forward in the prosthetic field, primarily linked to the ability of the clinician to exploit new technologies and new materials. In fact, the application of CAD/CAM techniques which are already well known, for example, in engineering, has allowed us to spread the use of new ceramic materials, making the spread much easier. Principally two materials have greatly changed dental practice: the zirconium and lithium disilicate.

D. Baldi (*) • J. Colombo • U. Hauschild Department of Fixed Prosthodontics, University of Genova, Genova, Italy e-mail: [email protected]; [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_50

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The purpose of this publication will be to investigate the characteristics of these two materials and their clinical applications, trying to rationalize the selection criteria and operational protocols. In fact, if, on one hand, therapeutic results that were unthinkable until a few years ago are possible, on the other hand, these materials require a thorough understanding of their characteristics in order not to run into avoidable failures. Keywords

Prosthodontics • Dental materials • CAD/CAM • Metal-free • Prosthetic protocols • Dental ceramic

Introduction In recent years, the changing socioeconomic conditions of the population and the increased attention to everything that concerns the sphere of personal image and aesthetics have had a profound effect on society. Furthermore, relatively new concepts in medicine have been established. These concepts include minor invasion and biological saving, that is, the tendency to choose therapeutic approaches that are minimally invasive and more respectful of the anatomical structures of the patient, where this is obviously possible [1]. For this reason, over the years dental research has been decisively oriented toward the development of new materials, especially in the prosthetics and restorative field, that would provide an improvement in aesthetic performance and a less invasive approach concerning dental structures. In conservative dentistry, one merely has to think about the current predictability of dental enamel adhesives and composite resins that consent infinitely more aesthetic and more conservative approaches than the use of traditionally reliable materials, such as silver amalgam. In the prosthetics field, innovations have been even more overwhelming. Simply think that up to not more than 15 years ago, the state of the art of complete tooth restoration was represented by gold artifacts and ceramics or, in situations limited to frontal areas, by integral ceramic articles. At present two major revolutions have been made. The first is represented by the introduction of a range of new materials which, in a relatively short period of time, are proving to be able to ensure functional performance comparable to those of materials with a much more experienced background. I am particularly thinking about the zirconium, in a monolithic form or layered with ceramic, and lithium disilicate [2]. The other big novelty has been the development and the perfectioning of CAD-CAM systems for the processing of these materials, which have ensured their accessibility for most dentists. Indeed materials such as zirconium and lithium disilicate were introduced because the simultaneous development of CAD-CAM systems allowed to obtain

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products by milling blocks of material which were previously prepared in industrial form. This system consents the manufacturing of articles starting from a digital print of a preparation or of a stone model. The CAD (Computer-Aided Design) part allows to project the scanned image on a computer where, by means of a software, the operator can establish the parameters of the artifact. At this point, the file is sent to the CAM (Computer-Aided Manufacturing) system to realize the reconstruction with a micro milling machine. A block of the selected material, which has been industrially produced, is milled to the exact dimensions established by the operator. This is very important because the process of the realization of industrial and standardized blocks ensures a controlled production that compensates the elevated contraction of highly resistant ceramic materials during the sintering process [3]. In a very short lapse of time, all this has led to a range of innovations that require time and application in order to be included in the daily routine of a dental practice, as well as thorough knowledge, which can frequently be distorted by an abundance of often scientifically invalid information from companies. The purpose of this chapter is to provide a key to be able to understand the fundamental features of these new materials and their applications, in order to introduce them to daily practice without the risk of having to handle unexpected complications. This discussion will specifically examine two materials in particular, zirconium and lithium disilicate, which were introduced to dental practice in relatively recent times. There will be a brief analysis of their characteristics and above all of their clinical use.

Zirconium Zirconium commonly refers to zirconium oxide, which is a metal frequently found in nature combined with silica oxide. Zirconium is an extremely interesting material that has been valued in dentistry for some outstanding features, namely, its mechanical strength, its dimensional stability, and an elastic form similar to that of steel. From the chemical point of view, zirconium can take on three forms that vary in function of the temperature: the cubic phase over 2370  C, the tetragonal phase between 2370 and 1170  C, and the monoclinic phase below 1170  C. These three phases obviously correspond to different mechanical behaviors of the material. Without going into detail, what happens is that with decreasing temperature and thus in the transition from the cubic phase to the monoclinic, there is a worsening of the mechanical properties and a decrease in the density of the material that virtually translates into a reduction in the dimensions. For this reason, to stabilize the zirconium at room temperature and to control the transformations, metal oxides such as yttrium or ceria are added to the crystal structure. This leads to the formation of multiphase materials known as partially stabilized zirconium [4–7].

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An extremely important mechanical characteristic in dentistry use is the ability of the zirconium to stop the propagation of fracture lines. This characteristic, which is unique of zirconium oxide, is explained by the fact that the area immediately preceding the fracture is in a state of compression. This is due to the fact that when a fracture starts to propagate, it develops a state of stress at high energy that determines the transformation of the zirconium crystal from tetragonal to monoclinic. This determines a difference of volume between the area of the fracture and the following one, which leads to a compression that stops the fracture line. This extremely important feature does not obviously exclude the possibility that the material will fracture. It has however been highlighted how failures of this type are always attributable to either design errors of the article or clinical errors [8–10]. From the clinical point of view, zirconium can be used in two different forms: pure in the form of tetragonal zirconium polycrystals stabilized with yttrium or infiltrated in the interior of glass ceramics [11]. In the first case, we are dealing with a material, which has all the main characteristics of zirconium, thus mostly resistant to fracture and flections combined with high biocompatibility. The main drawback is the lack of translucency that makes this material less efficient in areas with high aesthetic value compared to integral ceramic systems. This material can be machined with CAD-CAM technology and used in a wide range of clinical situations. The main uses are related to a fixed prosthesis, in particular as a substructure for anterior and posterior crowns and bridges of limited extensions, or to realize abutments for dental implants. This material can also have other less frequent but still viable uses, such as for example, the realization of endodontic posts, onlays and overlays, and Maryland bridges [12–17].

Criteria What mainly interests the clinician are the criteria for the choice of this material that we are going to analyze. 1. Aesthetics– Substructures in zirconium have a great aesthetic advantage linked to the fact that, especially in the presence of periodontal phenotypes of type II, they allow to minimize the so-called umbrella effect and consequently the gray halo that shines below the portion of the free gum, which is fairly common in traditional metal-ceramic prosthesis. However, in general, the use of a white substructure in any case allows to minimize the effects of failures due to gum recession even in a medium-long term [18]. 2. Preparations – From the design point of view of the abutment, the choice of zirconium-ceramic does not determine upheaval in preparation techniques. Substantially the same types of preparation and the same space required for the

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modeling of a manufactured metal ceramic are anticipated, which is about 2 mm occlusal and 1, 5 mm circumferential. There are situations, especially in posterior occlusal or anterior palatal areas, in which one may decide not to coat the zirconium with ceramic, leaving the zirconium itself in occlusion, thus earning between 0.5 and 1 mm. This choice must be carefully weighed considering the occlusion and the type of antagonist, because zirconium has an elevated hardness and a different abrasion form from a natural tooth [19]. 3. Impressions – The impression technique and the material decided to be use do not vary when choosing artifacts in zirconium ceramic, metal ceramic. What may vary can be the retraction technique of periodontal tissues. In fact, the high aesthetic qualities of zirconium can, in certain situations, favor a juxtagingival positioning of the margin. 4. Luting – The luting of zirconium is always of traditional type. The systems are the same as those of the metal ceramic, and the choice mainly falls on glass ionomeric or zinc oxyphosphate cement. In fact, because zirconium is not a silicabased material, it does not turn out to be etchable, and thus an adhesive-type luting is not practicable. The only possibility of taking this path is to use silica-coating techniques of the substrate by means of high-speed silica powder jets [20]. This technique is recommended whenever one wants to realize articles, such as Maryland Bridge or overlays, in which the adhesive luting technique plays a key role.

Laboratory Techniques Zirconium is delivered to the laboratory in blocks of different shapes and colors ready to be milled. There are essentially two types of production of the blocks: uniaxial cold pressing and isostatic cold pressing. Uniaxial pressing provides for the application of a uniaxial pressure to zirconium powder contained in a mould. This determines a high resistance of the material due to the deformation and interlocking of the particles, but at the same time determines different degrees of density within the block due to friction between the particles and between the particles and the mould, which can compromise the integrity of the structure, even after the milling. Isostatic pressing involves the insertion of zirconium powder in a deformable container subjected to external isostatic pressure that is uniform in all directions. Regardless of the method used to obtain these zirconium blocks, they are referred to as unrefined. At this point they are stabilized and thickened by a sintering process in a furnace at zero pressure. In dental systems, materials are normally used in an unrefined phase because zirconium is easier to work and wears less on the milling machines, even if at this

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stage the blocks should be milled to an enlarged size of 20–25 % to compensate for the shrinkage during sintering. With regard to this fact, the unrefined zirconium blocks have a label that identifies the density, so that the technician can exactly calibrate the milling machine to obtain a correctly oversized article [10]. Alternatively the mechanical properties of zirconium can be further improved through a process called hot isostatic post-compaction (HIP) that eliminates all porosity and increases the hardness and density of the material, making it unnecessary to compensate for the sinter shrinkage [21, 22]. At this point, the dental technician, regardless of whether he/she chooses unrefined zirconium blocks or HIP, has two techniques for the construction of the building available: CAD-CAM or MAD-MAM. In both cases, the work is subdivided into three phases: scan, design, and milling. The CAD-CAM, which we have already mentioned earlier, involves the scanning of the abutment, either in the mouth or on the model, and the creation of an image on the computer, on which a user can draw the artifact and send data to a milling machine that realizes it, via software. Milling is a subtractive process starting from a block and is the most widely used method. However, in CAD-CAM systems, there are also alternative methods to obtain the artifact. The additive method and the selective fusion mainly also exist. The additive method involves the construction of the article by adding material to an abutment. For an article in zirconium, an oversized metallic abutment must be realized to compensate for the sintering shrinkage. Compact zirconium powder will be added under isostatic pressure and then sintering will be carried out. On the other hand, selective melting is an experimental method for zirconium, but widely used for metals, which involves the creation of the article-collecting CAD data and then proceeding to melt thin layers of powder, which is susceptible to the heat of a laser beam. CAD-CAM systems, which are currently the most widely used, can be closed or open. That is, the dental laboratory can realize the design on a type of file that may or may not be read by machines of different brands from the one that scanned the file. Closed systems were the first to be introduced to the market and have a major setback of forcing the technician, after having made a significant financial investment, to work only with that particular manufacturer. The advantage is linked to the fact that these large milling centers use much more expensive machines that guarantee the milling also of HIP blocks. In open systems instead, which are becoming increasingly more popular, there is the opportunity of interacting with different CAD and CAM systems using universal file formats. The advantage is related to greater versatility and of the laboratory being able to work with different milling centers [23, 24]. The MAD-MAM system also consists of two phases: the first MAD (ManualAided Design) consists of transferring information. Generally a wax copy of the artifact to be obtained is made and it is then position in a pantograph. At this point the MAM (Manual-Aided Manufacturing) system component comes into action. The “tactile arm” of pantograph reads the wax model and transfers the information to the milling machine arm, where a system of drills grinds the zirconium block.

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Also in this case, if unrefined zirconium is being used, the final dimensions of the article after milling have to be oversized by 20–25 % to compensate for sinter shrinkage. These systems exploit pantograph technology, which has been known for hundreds of years, and their diffusion has been supplanted by the adoption of CAD-CAM systems. However, manual systems, with respect to softwares, have the great advantage of being able to intervene and correct any errors while making the wax model [24, 25].

Kinds of Operations Crowns and Bridges Zirconium gives the structure stability and supports the layered ceramic. This consents an increase in the transparency of the color and the uniqueness of the color with stratification by hand. The product combines all advantages to give aesthetic, functional, and mechanical yields. It is very suitable in cases where there are missing teeth to construct bridges between natural teeth and over gaps created by extracted teeth. Zirconium exists in different qualities, depending on the sintering of more or less transparent materials. The more opaque zirconium consents the covering of devitalized dental parts or metal posts. The semitranslucent zirconium, as the name suggests, lets more light in and however gives the same stability and increases the aesthetics due to the play of light. The monolithic zirconium is also known as “Prettau,” which is a name that comes from the country of its first user, Enrico Steger, who has made it a strong point of his company as well as marketing it worldwide. Until a short time ago, only one type existed. It had to be colored before sintering to render the raw element more natural and aesthetic. Through this technique, the technician could choose a color according to the needs of the case. The monolithic zirconium, against a less aesthetic yield, allows to obtain greater stability and resistance to chewing forces, eliminating the possibility of any “chipping.” Then there is a third relatively new variant of zirconium, commercially known as is zirconium anterior, which is created via a different pressure and temperature of sintering. A greater transparency that however maintains a higher mechanical strength of lithium disilicate is able to be reached, thanks to this variant. It is not suitable for structures that are too long.

Abutments Besides using zirconium for the construction of crowns and bridges, it best demonstrates its properties even in the manufacturing of abutments for implants.

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The exceptional hardness of this material allows the direct screwing on of the platform, which is a progress that is not however particularly suitable because of the different resistance between zirconium and titanium. The latter is less hard and therefore subject to wear over time if in contact with the first. The state of the art in oral implantology suggests a bonding connection between a titanium abutment and an individualized body in zirconium. This allows to take advantage of a perfect connection between titanium and titanium, a proper support to the soft tissues, in addition to a more pleasant aesthetics and a maximum biocompatibility, thanks to the characteristics of zirconium.

Complex Direct Screwing Cases In complex cases it may be necessary to use a body in translucent or monolithic zirconium, which has to be layered successively. The latter has a less aesthetic yield but guarantees greater resistance to fracture and chipping. If we work on implants, it is possible to screw them directly to the structure after a well-calculated study of size and morphology. Nowadays, thanks to the expansion that guided computer surgery is having, it is possible to have definite implant positions with the positive result of an access to the screw through the central axis of the manufactured teeth. This can be achieved, thanks to a positive implant design following prosthetic rules. In this type of structure, it is however advisable to cement titan base abutments, only after having finalized and polished the artifact. Only after a final inspection, we can proceed with bonding just like with individual abutments. As we can see, the zirconium has excellent translucency, which the value may vary depending on the choice of the specific material, needs, and abilities of the technician.

Lithium Disilicate Lithium disilicate (LIS2) is a material that belongs to the glass-ceramics family. The main component is, as in most of the ceramics, silicon oxide, which is supplemented with materials so-called smelters that serve to lower the melting temperature and others so-called “stabilizers” which give the glassy component greater stability and better chemical-physical characteristics. These materials are produced by compacting powders that are then heated to such temperatures, which bind the particles together. They differ from traditional ceramics because they also contain a glassy phase which is able to interact with the remaining solid refractory material and while cooling it may solidify and bind the particles that have not yet merged together. The glassy phase gives the ceramics some peculiar properties including a good hardness and mechanical strength and excellent corrosion resistance.

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Lithium disilicate has been developed to improve the mechanical properties of glass ceramics and expand the use of integral ceramic in posterior sectors. It presents a flexural resistance of up to 450 MPa, and the fracture toughness is three times more in comparison, for example, to leucite. In fact, in this material, the main glass phase, which is precisely a lithium disilicate, covers approximately 70 % of the total volume and forms an unusual microstructure. This microstructure consists of many small crystalline plates, which are connected to each other and oriented in a confused way. These plates seem to play a very important role in increasing the resistance of the material. This arrangement also appears to have the capacity to restrain fractures, causing them to branch and deviate. Lithium orthophosphate is a secondary vitreous part, which is also present in this material, but it is much less important in relation to the total volume. Lithium disilicate is obtained by a hot-pressing system at 920  C [26–28]. From the clinical point of view, lithium disilicate can be treated as an integral ceramic, so the choice to perform a restoration with this material is mainly due to aesthetic reasons. Disilicate guarantees an excellent aesthetic appeal combined with excellent mechanical properties, definitely higher than those of feldspathic ceramics. The mechanical capacity of the material makes it possible to work with reduced thickness of material. In particular, especially in the anterior, we can reach thicknesses of restoration of less than 1 mm. Obviously being a highly translucent material that provides a great passage of light, it is very important to be careful, especially in the front areas of the substrate on which the restoration is carried out. The latter can greatly influence the aesthetic characteristics of the material. Moreover, the other great advantage of the disilicate is the ability to perform adhesive luting, thereby increasing the mechanical characteristics of the material. If this direction is chosen, the adhesive approach must be the same as that used with ceramic materials. Thus protocol anticipates well-defined steps on both the substrate and on the restoration. With regard to the substrate, the use of the rubber dam is recommended, followed by etching with 37 % phosphoric acid for 20 s and rinsing with water, and then the application of a primer. If there are areas of exposed dentin, air should be gently blown and thus the application of a bonding without polymerize. Instead with regard to restoration, acid etching with hydrofluoric acid at 9 % is recommended, followed by rinsing with water, silane application and then drying with preferably hot air. At this point a bonding without polymerize is applied. After the completion of the surface treatment the choice of the color of the photo or autopolymerizing resin cement, depending on whether you want to saturate or desaturate the final color. Then, the work is heated using special ovens and the color is applied inside the restoration with a spatula, proceeding to inserting the latter in the mouth. At this point excesses are removed and any necessary occlusal retouches are made [29, 30].

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Crowns and Inlays in Lithium Disilicate When we are faced with cases of single teeth, particularly the anterior or we have to complete a portion of the teeth even if partially filed, the best choice is an element in lithium disilicate, so as to mimic nature as much as possible. These elements can be modeled manually and pressed with appropriate ovens or milled by a CAD-CAM system. We can manufacture the anatomical shape or a “core” by later layering with ceramic, so as to add a touch of individuality and to increase reflection of the light. Disilicate inlays are more delicate and must be prepared by the clinic according to their mechanical properties. When luting it is necessary that the thicknesses are uniform and have a homogeneous support. It is recommended to use a microscope in order to have a better-sealed circumferential closure.

Cases with Mixed Use Between the Two Materials The clinician is free to choose the most suitable materials according to the complexity of the case and the requests of the patient. The processes and materials can be mixed with each other, but for what concerns the majority of our cases, the appearance of the frontal implants is obtain by using lithium disilicate, while in the area of the molars is layered with the zirconium ceramic, always taking into detailed account the criteria of protecting the zirconium where it is necessary to avoid any future cracks in the ceramic due to the chewing load.

Clinical Case The case presented is that of a young patient who has a fracture of the element 1.1 (Figs. 1 and 2). It is decided to remove the element and immediately insert a fixture in the postextraction position. A connective tissue graft taken from a part of the palate is simultaneously performed (Fig. 3). So, a transfer imprint with open tray technique is performed after a temporary period of about 3 months with a two tab Maryland Bridge (Fig. 4) and the realization of a customized abutment in zirconium, on which a second provisional structure is positioned to guide the maturation of tissues (Figs. 5 and 6). After other 3 months, the case is finalized with a lithium disilicate crown (Figs. 7 and 8). This case is an example of how the combined and aware use of these materials can guarantee excellent results.

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Fig. 1 Baseline of the patient: frontal view

Fig. 2 Baseline of the patient: palatal view

Fig. 3 Implant inserted and connective tissue graft

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Fig. 5 Zirconium abutment in situ after 3 months from the surgery

Fig. 6 Second provisional structure is positioned to guide the maturation of tissues

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Fig. 7 Final PFZ restoration: frontal view

Fig. 8 Final PFZ restoration: lateral view

References 1. Touati B, Miara P, Nathanson D (1999) Esthetic dentistry and ceramic restorations. Martin Dunitz, London 2. Zillio A (2013) Zirconia-The power of light. Teamwork Media, Brescia 3. Sturdevant JR, Bayne SC, Heymann HO (1999) Margin gap size of ceramic inlays using second generation CAD/CAM equipment. J Esthet Dent 11(4):206–214 4. Piconi C, Maccauro G (1999) Zirconia as a ceramic biomaterial. Biomaterials 20:1–25 5. Von Clausburch C (2003) Zirkon and zirkonium. Dent Lab 51:1137–1142

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6. Frieman S (1991) Introduction to ceramics and glasses. In: ASM engineering materials handbook, vol 4. ASM International, Philadelphia, pp 1–40 7. Duran P, Moure C (1984) Sintering at near theoretical density and properties of PZT ceramics chemically prepared. J Mater Sci 20(3):827–833 8. McLaren EA, Giordano RA (2005) Zirconia based ceramics: material properties, esthetic and layering techniques of new veneering porcelain. Quintessence Dent Technol 28:100 9. Helvey GA (2006) Press to zirconia: a case study utilizing cad/cam technology and the wax injection method. Pract Proced Aesthet Dent 18(9):547–553 10. Raidgrodski AJ (2004) Contemporary all ceramic fixed partial dentures: a review. Dent Clin N Am 48:531–544 11. Christel P, Meunier A, Heller M, Torre JP (1989) Mechanical properties and short term in vivo evaluation of yttrium oxide partially stabilized zirconia. J Biomed Res 23(1):45–61 12. Meyenberg KH, Luthy H, Scharer P (1995) Zirconia posts: a new all-ceramic concept for nonvital abutment teeth. J Esthet Dent 7(2):73–80 13. Luthardt RG, Sandkhul O, Reitz B (1999) Zirconia-TZP and alumina-advanced technologies for the manufacturing single crowns. Eur J Prosthodont Restor Dent 7(4):113–119 14. Glauser R, Sailer I, Wohlwend A, Studer S (2004) Experimental zirconia abutments for implant supported single tooth restorations in esthetically demanded regions: 4-years results of a prospective clinical study. Int J Prosthodont 17(3):285–290 15. Blatz MB (2002) Long term clinical success of all ceramic posterior restorations. Quintessence Int 33(6):415–426 16. Hayashi M, Tsuchitani Y, Miura M, Takhsige F (1998) 6-years evaluation of fired ceramic inlays. Operat Dent 23(6):318–326 17. Sailer I, Feher A, Filser F, Luthy H (2006) Prospective clinical study of zirconia posterior fixed partial dentures: 3-years follow up. Quintessence Int 37:685–693 18. Gamborena I, Blatz MB (2006) A clinical guide to predictable esthetics with zirconium oxide ceramic restorations. Quintessence Dent Technol 29:11–23 19. Robertson T, Heymann H, Swift E (2002) Sturdevant’s art and science of operative dentistry, 4th edn. Mosby, St. Louis 20. Blatz MB, Sadan A, Blatz U (2003) The effect of silica coating on the resin bond to the intaglio surface of Procera AllCeram restorations. Quintessence Int 34(7):548–555 21. Rogers J, Weber W (2007) Ceramic materials are not all the same. Spect Dialogue 6:76–80 22. Li J, Liao H, Hermansson L (1996) Sintering of partially stabilized zirconia hydroxyapatite composites by hot isostatic pressing and pressureless sintering. Biomaterials 17(18):1787–1790 23. Reichert A, Herkommer D, Muller W (2007) Copy milling of zirconia. Spect Dialogue 6:40–56 24. Tinschert J, Natt G, Hassenpflug S, Spiekermann H (2004) Status of current CAD/CAM technology in dental medicine. Int J Comput Dent 7(1):25–45 25. Liu PR (2005) A panorama of dental CAD/CAM restorative dentistry. Compend Contin Educ Dent 26(7):507–512 26. Blatz MB, Sadan A, Kern M (2004) Ceramic restorations. Compend Contin Educ Dent 25 (6):412–416 27. Tinschert J, Natt G, Mautsch W, Aughtun M (2001) Fracture resistance of lithium disilicate, alumina and zirconia based three unit fixed partial dentures: a laboratory study. Int J Prosthodont 14(3):231–238 28. Margeas RC (2007) Material and clinical considerations for full coverage all ceramic restorations. Funct Esthet Restor Dent 1(3):20–24 29. Blatz MB, Oppes S, Chiche G, Holst S (2008) Influence of cementation technique on fracture strength and leakage of alumina all ceramic crowns after cycling loading. Quintessence Int 39 (1):23–32 30. Blatz MB, Sadan A, Kern M (2003) Resin ceramic bonding: a review of the literature. J Prosthet Dent 89(3):268–274

Ceramic Veneers in Dental Esthetic Treatments

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Dan Pătroi, Teodor Trăistaru, and Sergiu-Alexandru Rădulescu

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Aesthetic Criteria . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Lips . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Attached Gingiva Area . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Teeth . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Perception of Shape Depending on the Position of the Teeth . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Materials Used for Veneering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Acrylic Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Composite Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Ceramic Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Indications of Full-Ceramic Veneers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Contraindications of Full-Ceramic Veneers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Full-Ceramic Veneers Advantages . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Disadvantages of Full-Ceramic Veneers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Longevity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Therapeutic Protocol . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The First Treatment Session . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Second Treatment Session . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Third Treatment Session . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Fourth Treatment Session . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Aesthetic treatment of the anterior teeth has always been a challenge in clinical practice. Along with the evolution of dental materials, the number of therapeutic options increased: in addition to aesthetic fillings or direct composite resin D. Pătroi (*) • T. Trăistaru • S.-A. Rădulescu UMF Carol Davila, Department of Fixed prosthodontics and Occlusology, Bucharest, Romania e-mail: [email protected]; proteticafi[email protected]; [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_55

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veneering and all-ceramic crowns, ceramic veneers and inlays are now available. In these circumstances, dentists and their patients can choose the best therapeutic option from a range of possibilities outnumbered 15–20 years ago. This choice is based both on the study of aesthetic criteria and the patient desires in accordance with therapeutic needs that he has. The commonly used material for aesthetic restorations is dental ceramics, due to its color stability, biocompatibility, mechanical properties, and excellent aesthetic results. Minimally invasive dental restoration idea is gaining more ground in current dental practice, and thus full-ceramic veneers are increasingly used. There are three types of ceramics used for manufacturing these veneers: feldspathic ceramics, leucite-reinforced glass ceramics, and lithium disilicatereinforced glass ceramics. All these have special optical and aesthetic properties, miming natural tooth appearance, but the mechanical properties are different. As full-ceramic veneers are single-tooth restorations with very low thicknesses, mechanical resistance of the material is important both for their handling during manufacturing phases or intraoral cementation and especially for preserving the integrity of the restoration in the oral cavity. Teeth preparation techniques for full-ceramic veneers differ depending on the aesthetic needs of the patient, the type of veneers to be made, and their indications and contraindications. Operator protocol includes four treatment sessions at the end of which the patient has the veneers cemented on the natural teeth. There are many advantages of natural teeth veneering and some disadvantages, which are small compared to the benefits. The survival rate of full-ceramic veneers is very high, with some authors communicating a percentage of 96 %  2 % at 21 years and others of 95.6 % at 10 years. Keywords

Full-ceramic veneers • Preparation techniques • Advantages • Disadvantages • Survival rate • Success rate • Veneers longevity • Veneers indications • Veneers contraindications • Aesthetic criteria • Ceramic materials • Composite materials • Acrylic materials • Feldspathic ceramics • Leucite-reinforced glass ceramic • Lithium disilicate-reinforced glass ceramic • CAD/CAM technology • Therapeutic protocol

Introduction The success of an aesthetic restoration can be cumulated under the concept: “Beauty lies in the eyes of a concerned”; thus the dentist becomes an artist, and his work of art is the patient himself, both being involved equally, emotionally, and subjectively during treatment plan. Beautiful art always catches the attention of those who look, and smile itself being the first contact with the speaker exerts a powerful influence on first impression. It is

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said that “a smile is worth a thousand words.” Smile is given more importance in modern society, so with an attractive and healthy smile the patient can integrate more easily into society. Dentistry involves a synergistic combination of biofunctional-aesthetic-mechanical principles and constantly changing in a direction always influenced by social, economic, and cultural factors. Thus emerged in recent decades an increase in concerns for dental aesthetics and sometimes cosmetic dentistry. The main purpose of this domain is the analysis, design, and implementation of the “perfect smile,” based on rules, located on the border between art on the one hand and science on the other. In the designing and implementation of a treatment plan, compliance with these aesthetic criteria plays an important role in getting the best and lasting therapeutic outcome, at whose touch contribute, at least to the same extent other principles mentioned requirements. Aesthetic treatment of anterior teeth has always been a challenge in current clinical practice. With the development of dental materials, the number of treatment options increased; now available, in addition to aesthetic restorations or direct composite resin veneers and all-ceramic crowns, veneers, and ceramic inlays. In these circumstances, dentists and their patients can choose the best therapeutic solution from a range of possibilities, wider than 15–20 years ago.

Aesthetic Criteria Aesthetic rehabilitation of the anterior region must take into account the general criteria of beauty, which specialists in cosmetic dentistry have standardized as features of “perfect smile” with all components: lips, fixed gum area, and teeth. In turn, each of the three main elements has several variables that must be considered: Lips – shape, color, texture, symmetry, tonicity, and smile line Attached gingiva area – color and texture, shape and volume, symmetry, contour and gingival zenith, and gingival papillae Teeth – shape, dimensions, central incisor ratio and golden proportion, mark (decreasing vertical dimension of lateral teeth), midlines, axial tilt and location of contact points, the free incisal edge, lower lip line and buccal corridor, arch shape, teeth texture, and teeth color

Lips Overall artistic impression created by a picture can be enhanced or diminished by the frame. Lips can have the same effect on smile appearance. It is analyzed by: • • • •

Size Symmetry Shape Color

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Tonicity Texture Corners line (parallel to the pupillary line) Positioning ratio compared to the teeth

Average length of the upper lip, measured at rest from the center bottom of the nose to its inferior margin, has the following average values: 20–22 mm in young women and 22–24 mm in young men [1]. Shorter lips allow visibility of teeth at rest, and hyperactive lips reveal a great amount of attached gingiva during smiling, the so-called gummy smile. The analysis of these parameters is made clinically by asking the patient to pronounce phonemes “M” and “E.” Through repeated pronunciation of the letter “M” on average, a portion 2–4 mm from the free lateral incisor is visible, and the pronunciation of the letter “E” determines the maximum extension of the lips highlighting the smile line which can’t exceed more than 2–3 mm to the central incisor gingival margin [1]. The amplitude of vertical movement of the upper lip from the rest position to the highest position in laughter is typically 6–8 mm; in the case of hyperactive lip, movement is about two times larger [1]. Lip appearance varies from person to person and obviously changes with age. Position and volume of frontal dentures can significantly alter the appearance of the lips. For this reason, some descriptive parts (smile line, midline, canine line) can be sent to the dental laboratory in the idea that the denture must be performed in harmony with the surrounding elements.

Attached Gingiva Area Pleasant appearance of smile can be affected by some deviation from the normal attached gingiva area. Shape and volume – the volume is physiologically variable. In the literature three gingival patterns have been described, normal, thin, and thick, and these three types are differentiated by different thicknesses of alveolar bone margin, the appearance of the gingival free margin, and gingival contour. Thin and thick gingival types evolve differently if periodontal damage is present: thin type evolves toward bone resorption and gingival retraction, whereas in patients with thick type, it is associated with true or false gingival pockets and gingival hypertrophy [2]. In gingival hypertrophies, constitutional or pathological, the visible teeth shape changes; they become small and wide. Symmetry, contour, and gingival zenith – gingival contour follows the free gingival margin so that the free gingival margin of lateral incisors must be above the central incisors and canine line. A detailed analysis of gingival contour introduces a third element, namely, gingival zenith: the highest point of the gingival curvature of each tooth. Due to the mesial tilt of dental vertical axis, this point is not always located midway mesial-distal but slightly distal [3].

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Gingival papillae – another element that influences the aspect during smiling is the gingival embrasures: triangular shape spaces delimited by proximal faces and edges of the teeth, from the free gingival margin to the contact point. These spaces are occupied, normally, by the gingival papillae. When great gingival loss is present, the papilla volume is smaller than the newly created space, and black triangles appear, which are unaesthetic.

Teeth Teeth shape – the natural look of teeth varies according to age, sex, constitutional type, and patient personality; moreover, some theories argue that the shape of upper central incisor corresponds to the inverted contour of the face. The shape is determined by the ratio between the height and width of the tooth, the free edges, proximal faces, and free gingival margin. Visual perception of teeth is influenced by their inclination to the reference axes in the three plans. Teeth dimension and “golden proportion” are closely related to the shape and size of the teeth. The size will vary from patient to patient, and it is considered normal if they are proportional to each other and at the same time as the size of the face, the proportionality in which we find “the golden proportion.” “The golden proportion” or “the golden ratio,” related to 1/1,618, is a permanent retrieved constant in all shapes considered beautiful or proportionate. Starting from the elements of nature and to analyze the human body, this “golden rule” itself defines beauty as perceived by the human eye. This ratio can be found in people deem “beautiful” [4]. Restricting the theory of divine proportion to the perfect smile, it is considered that assigning a unit value to the mesial-distal width of the lateral incisor visible from the standard frontal view and then the mesial-distal length of the central incisor will be 1,618 times greater and the canine 0.618 times less. As well, the actual average size of the frontal teeth group varies according to the sex. Speaking of shape and size, it must be specified the role of visual illusion in the perception of an image. The optical illusion is the phenomenon by which the eye perceives a different image of reality. By understanding this phenomenon, it can become a useful tool in practice. When creating an optical illusion, an important role is played by the light and the white/black alternation [5]. Based on this idea and analyzing the degree of convexity of the facial surface, which changes depending on the amount of light reflected, the notion of “apparent surface” of a tooth appeared. This is represented by the flat area of the facial surface, which may be closer or further as size from the contour of the tooth. Basically, tooth size perceived by the human eye is given by this “apparent surface” that reflects light toward the observer, while the curved areas, which continue to the edges of the tooth, reflect the light sideways, appearing darker and narrower than in reality. Thus, when the size of the two areas is approaching the values, the tooth is flat and “seems” wider, and vice versa.

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Perception of Shape Depending on the Position of the Teeth Midlines – symmetry is an essential element of the concept of “beauty.” Numerous studies have shown, however, that in reality there is not absolute symmetry. Human faces, even the most perfect, have small asymmetries compared to the median or to the horizontal plane. One of the most used techniques for the analysis of facial symmetry is using the midline as landmark. Face image is divided into two halves which are then replicated and placed in the mirror. The two resulting images are similar to each other and to the original. As the differences are smaller, the more elements respect the symmetry. Thus, midline becomes an important milestone in facial and dental aesthetics. The normal situation implies that the two incisive lines should correspond one to another, and both should correspond to the facial midline. Another element that is analyzed is the buccal corridor. It is the term used to describe the space between vestibular surfaces of posterior teeth and corners of the mouth or the inner wall of the cheek. A narrow arch will allow a very large buccal corridor, with dark look and, conversely, a wide arch will narrow the buccal corridor, creating a feeling of “mouth full of teeth.” Axial inclination and location of contact points – another criterion used in the analysis of aesthetic smile is slight inclination of the vertical axes of teeth toward mesial. This tilt causes, along with tooth shape, the position of contact points on a vertical axis ranging from central incisors to canine and further premolars. In turn, the position of the contact point determines the shape and size of the gingival and incisal embrasures. The lack of embrasures or excessive size causes deviations from normal that give a bad look. Incisal edge, lower labial line, and buccal corridor – the free edge of upper frontal teeth is a critical element of the perfect smile. Due to the different sizes of teeth in this area, incisal edges lie on a line with two curves: concave at the level of central incisors, continuing convex toward the lateral incisors and returning to its original shape toward canines. As against the lower lip line, central incisors and canines have to be in slight contact and the lateral incisors at a distance of approximately 0.5–1 mm. Also, the symmetry of these landmarks as against the midline is analyzed. In the overall smile analysis, it is considered that the incisal line is a continuous curve, concave upper, and must be parallel with the lower lip line. Due to the normal curvature of the dental arch and remoteness of the point of observation, the apparent vertical and mesiodistal dimensions of teeth, from canine to second molar, should be decreasing distally. The ratio between these apparent sizes, the width of the entire arch visible when smiling, the wideness of the frontal group, and the width of central incisor respect, ideally, the gold proportion. Dr. Levin created, in 1975, a teeth proportionality analysis grille for wide smile. The set designed by Dr. Levin consists of 20 different sizes of grids that follow the same pattern, made of transparent material, so that they can be used in the dental office by direct overlap [6]. Arch form – in all three normal forms of arches, parabola, ellipse, and “U” frontal teeth are aligned on a curved segment. In practice, the analysis of arch form is drawing a line through the central canines, which must pass through the retroincisive papilla.

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A line located anteriorly shows a stretched arch, and the line situated posterior from the incisive papilla occurs with narrow arches. Generally, these abnormalities of dental arch form are associated by default with dentoalveolar spacing or crowding, modifying obvious negative aesthetic. Dental surface texture – differs according to person. Facial surfaces can have small concavities, convexities, fissures, grooves, and striations. Dental surface geography has a very important role in light reflection and color perception; that is why special attention is required when restoring these surfaces with dental remodeling techniques, partial or total, with direct or indirect methods. Teeth color – color is the property of a substance to reflect or absorb some of the visible white light and perceptual ability of the human eye. The color cannot exist without light, object, and observer. Its perception is directly influenced by the characteristics of these three elements. All these characteristics vary in the population from one race to another. The criteria presented above are generally available in the Caucasian racial group. Thus, for example, African or Asian population, dental characteristics differ from the Caucasian population. Although dental aesthetics can be considered a form of art, there are principles of dental aesthetics, defined in literature, that can be used as a guide for real therapeutic success.

Materials Used for Veneering Dental veneers (sometimes called porcelain veneers or laminated porcelain veneers) are very thin-fixed partial dentures, covering the front of the teeth (facial) in order to improve the patient’s appearance. They can be molded into any shape, size, length, and color and be bonded (cemented) adhesive to the surface of teeth. Veneers were invented in 1928 by dentist Charles Pincus from California to be used in order to temporarily change the appearance of actors’ teeth for various movies made there [7]. Later, in 1937 this dentist fabricated acrylic veneers that were cemented temporarily because adherence of dental cements available at that time to dental tissue was very poor. Buonocore, in 1959, introduced etching in order to adhesively cement porcelain veneers to the etched enamel. Simonsen and Calamia [8] in 1982 showed that even porcelain can be etched with hydrofluoric acid and thus obtain a bond strength between composite resin and porcelain, supposed to be able to maintain veneers on a permanent tooth surface. This was confirmed by Calamia [9] in an article describing a technique for fabrication and cementation of acid etched all-ceramic veneers using the refractory die, and Horn [10] describes the method of manufacture of veneers on platinum foil. Dental veneers can be made of ceramic, composite resin, or acrylic material. Thus, the final aesthetic degree of veneer depends directly on the aesthetic possibilities of the material used. Ceramic veneers have the highest degree of aesthetic, which basically mimic the appearance of natural teeth due to the optical properties of ceramic materials: translucency, fluorescent, transparency or opalescent, and color

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Table 1 Proprieties of the materials used for veneers Optical proprieties Aesthetic proprieties Biologic proprieties Mechanical proprieties Survival rate Success rate Cost Cementation Treatment session

Acrylic Low Low Gingival irritation Poor Low Low Low Provisional or absent 1

Composite Good Medium-good Good Medium Medium-high Medium-high Medium Adhesive 1–3

Ceramic Excellent Excellent Biologically inert High High High High Adhesive 3–5

stability in time. Composite veneers have a lower degree of aesthetics due to the lack of translucency of the material and color stability in time. In a clinical trial, Gresnigt, Kalk, and Ozcan discuss the emergence of a porous surface with composite veneers and marginal discoloration [11] which leads to decreased aesthetic properties of these types of veneers. Acrylic veneers have the lowest aesthetic degree due to limitations imposed by the material, which are generally used for temporary veneers made in the dental office. A comparison between the three types of materials used in veneering is shown in Table 1.

Acrylic Materials Although the aesthetic qualities of this class of materials are low, they are commonly used to obtain the provisional veneers during dental treatments or for temporarily changing the appearance of teeth (e.g., Snap-On Smile™, DenMat Laboratories, USA). When it comes to dental treatment through veneering, after tooth preparation, the patient should recover both physiognomy and jaw functionality. Thus, through direct methods of provisional prosthesis, temporary veneers are made to be worn by the patient until the final restorations will be done. The materials used are entirely provisional prosthetic materials: Telio CS C & B (Ivoclar Vivadent), Access Crown (Centrix), Protemp (3 M ESPE), Luxatemp (DMG), etc. In current dental practice, these restorations are made directly in the dental office, by copying natural teeth before preparing them or by coping a wax-up model. Later, after the preparation of the teeth, a temporary prosthetic material is applied in this impression, and the impression is inserted in the oral cavity. After material setting the impression is removed from the oral cavity, the temporary veneer is removed from the impression, and it is finished and then temporarily cemented to the dental preparation. The patient wears it until the end of the treatment. Snap-On Smile™ (DenMat Laboratories, USA) is a reversible treatment procedure, completely noninvasive. The manufactured restoration can be easily removed

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and reapplied by the patient, because it is not cemented to the natural teeth. This therapeutic procedure involves a noninvasive approach, fully reversible and easily accessible regarding the price of a restorative and cosmetic dentistry procedure. Clinical protocol is quite simple. The first step is to determine whether the patient is a suitable candidate and if he/she can understand both the advantages and limitations of this therapeutic option. Then, the shape and position of teeth that are intended to be corrected are analyzed and recorded. No preparation of the natural teeth is necessary. Full-arch impression using silicone material are made and an occlusal record is necessary. It is important that dental impressions to be accurate and precise to record all the convexities and the full contour of natural teeth, and at the free gingival margin, the impression should not be torn or modified. The impressions are then sent to the dental laboratory with a form detailing the changes, and there the provisional will be made following aesthetic and functional criteria. Patients can wear these restorations whenever they want and how long they are comfortable. They can apply and remove them from the oral cavity without the need for a dentist intervention.

Composite Materials Composite materials have better optical and aesthetic properties than acrylic materials but lower comparing with ceramic materials. Facial veneering concept became possible after the discovery of adhesive cementing techniques that provide aesthetic and functional needs of maintaining adequate bond. Meanwhile, originally used acrylic veneers have become aesthetically unacceptable due to early loss of color and low resistance in the oral environment. These disadvantages are offset to a large extent when using composites for veneering and are absent in ceramic materials [12]. Porcelain veneers have the highest rate of survival time, followed by composite veneers, and then the acrylic ones due to both physical and biological properties of these materials and methods of manufacturing the veneers (Table 1). If we refer only to methods of producing indirect composite veneers, then their survival rate is comparable to ceramic veneers [13]. To achieve composite veneers, it can be used both direct and indirect methods. Direct techniques aimed at achieving veneers in the dental office, by the dentist, using composites for crown restorations. The indirect techniques involve the dental laboratory, manufactured by a dental technician and then cemented in the mouth by the dentist. When direct technique is used, the dentist starts with ultrasonic scaler, airflow, and brushing in order to clean the teeth for both color composite materials election and prepare their surface for an adhesive technique if teeth preparation is not necessary. When the preparation is necessary, this should be achieved with tapered diamonds with smooth tip, especially for creating a feather edge margin, with continuous water cooling, removing between 0.5 and 1 mm of the enamel, taking care to keep the preparation solely at the enamel level. Incisal edges will be prepared with embrasures according to dentine mamelons. At the proximal surfaces the

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contact point will be removed only where caries will require or when we want to close the interdental spaces. Before starting the actual direct veneering process, whether or not the tooth has been prepared, it is necessary to isolate the teeth involved. This can be done with the OptraGate (Ivoclar Vivadent) device to avoid the change of clinician perception of hue and translucency of teeth generated by the rubber dam isolation. When using this device, gingival eviction is needed to achieve isolation in the gingival sulcus. It will use one or two nonimpregnated braided cords, depending on the depth of the gingival sulcus. Depending on clinician experience and indications of the composite material total etch, no etch or self etch adhesive systems will be used. Of these, the total etch systems involving etching of the enamel with 37.5 % phosphoric acid provide the best adhesive bond of the composite to the tooth [14]. Composite resin will be added in layers, using shades of the dentin and enamel for final color-rendering effects for the veneer and shades (clear, transparent, opalescent) for achieving transparencies from the incisal edge or proximal surface, according to final aesthetic needs. Final stratification will be about 2 mm, and the application of composite resin will start from the center of the tooth extending to the proximal and facial areas. For application and modeling layers of the composite, it can be used as dental spatulas and special tools for modeling, and to restore the contact point and the proximal areas, celluloid matrix fixed with wood wedges with triangular profile are used. Composite layer polymerization is made with a LED lamp using a light intensity and curing time according to the manufacturer. Finishing will be done with diamond or carbide burs along with abrasive discs and tapes, and the final polish will be achieved with rubber cups, which incorporate particles of aluminum oxide or diamond, and brushes with polishing paste. Direct veneering technique versus indirect technique has the main advantages of short working time, one session of treatment required, and low-cost price. Composite veneers, as well as the ceramic ones, can be achieved through indirect techniques involving dental laboratory. In this case they will be made by the dental technician, and the dentist in advance must polish the tooth/teeth involved, make an impression of the prepared teeth, and send the impression to the laboratory with a lab sheet that contains data about the color and the final shape of the teeth. Dental technician is the one who will manufacture the veneers with composites for dental laboratory use: Sinfony (3 M ESPE), SR Nexco (Ivoclar Vivadent), etc. Later they will be cemented by adhesive intraoral dental composite cements: Variolink II (Ivoclar Vivadent), Nexus NX3 (Kerr), Duo-Link (Bisco), etc. As a result of development of composite materials, prefabricated composite veneers were recently introduced in the dental practice. These prefabricated veneers were designed as an alternative to full-ceramic veneers, to be more accessible for patients in need of an aesthetic solution regarding the price compared with that of a full-ceramic veneer. They are available in different shapes, colors, and sizes, in kits which include a flowable composite material for intraoral cementation. Such a system is Componeer (Colthene). These prefabricated veneers will be customized

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directly intraoral by the dentist depending on the patient’s needs and then be adhesively cemented to the frontal teeth of the patient in order to restore the patient’s aesthetic appearance. The necessity of prior preparation of the teeth is minimal or absent even in case of using these veneers and is necessary only one appointment at the dentist for their application, as in the case of direct veneering. This new technique has the potential to be used routinely to lengthen the anterior teeth, to correct the teeth with minor malpositions, to mask some dental stains, and to close the treme or diasteme. The technique can also be used to restore extended caries and dental fractures, especially when other treatment options are refused by the patient for financial reasons. However, it is extremely important to perform controlled clinical trials with this restoration technique, before recommending it without restriction in general practice [15].

Ceramic Materials Since the introduction of indirect veneering technique more than two decades ago, restoration of teeth with the acid-etched ceramic veneers has proved to be a sustainable and aesthetic treatment. Clinical success of this technique can be attributed to higher attention on the details of a set of procedures: • • • • •

Case planning with accurate indications Conservative tooth preparation Correct selection of the ceramic used Appropriate selection of materials and methods of adhesive cementation Proper planning for continued maintenance of these restorations

This treatment method is used because of color stability of ceramic, biocompatibility, mechanical properties, and excellent aesthetic results. The idea of minimally invasive dental restorations is gaining more and more ground in current dental practice, and so, ceramic veneers are increasingly used. There are three types of ceramics used for manufacturing these veneers: • Feldspathic ceramics • Glass-based ceramic reinforced with leucite crystals • Lithium disilicate glass ceramic All of these have special optical and aesthetic proprieties, miming natural tooth appearance, but the mechanical properties are different. As ceramic veneers have very low thicknesses for single-tooth restorations, mechanical resistance of the material is important both for their handling during manufacture and intraoral cementation especially to preserve the integrity of the restoration in the oral cavity. Feldspathic veneers are made by firing ceramic powder on metal foils (e.g., platinum) or directly on the refractory die. Mechanical strength of these veneers has low values of 30–40 MPa, which requires their handling very carefully. After the

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adhesive cementation on dental hard tissues, their strength reaches proper functionality; in the literature there are studies that communicate a success rate of about 95 % at 10 years [16]. Glass-ceramic veneers have a higher resistance compared to the feldspathic ones, reaching values of 170–180 MPa for glass-ceramic reinforced with leucite crystals and 400 MPa for the lithium disilicate one. Glass-ceramic material has a polycrystalline structure produced by controlled crystallization of glass. This class of materials has many common properties such as glass and ceramics. They have an amorphous phase and one or more crystalline phases produced by a “controlled crystallization” as opposed to spontaneous crystallization, which usually is not desirable in the manufacture of glass. Properties resulting from the process of “controlled crystallization” include the absence of porosity, high strength, hardness, translucency or opacity, color, opalescent, very low thermal expansion, or even absent, high-temperature stability, fluorescence, biocompatibility, and thermal and electrical insulation capacity. These properties can be adjusted as needed by controlling the composition of the base glass and the type of crystals (leucite or lithium disilicate) introduced into the ceramic structure. To obtain full-ceramic veneers, these materials are either hot injected into a pattern that is shaped like the future veneer or CAD/CAM milled from ceramic bars. Hot injection is performed at high temperatures of 1100  C for system IPS Empress (Ivoclar Vivadent) containing ceramic reinforced with leucite or at 920  C for system IPS e.max Press (Ivoclar Vivadent) containing ceramic reinforced with lithium disilicate, either at low temperatures of 160  C for Cerestore system. The pattern is made in the dental laboratory from refractory materials using wax elimination technique. Veneers thus obtained are then processed and individualized by the dental technician for obtaining final restoration which will be adhesively cemented on the prepared teeth. CAD/CAM technology is the newest method introduced in prosthodontics, and its target is automated manufacturing of dental restorations. The design of the restoration is made using a computer program after prior intraoral scanning of the prepared tooth (optical impression) or of the working cast in the dental laboratory, and the finite prosthetic veneer is obtained by a milling machine from a ceramic block. This technology produces dental restorations that have superior marginal adaptation to the preparation [17] compared with conventional techniques. Theoretically, it would result in a better survival rate for these restorations, but studies published in the literature communicate survival rates similar to hot-pressing technique, when evaluating prosthetic restorations[18]. Regardless of the technology used, the preparation of teeth that are about to be veneered is the same. The techniques of preparation are divided into the following categories: • • • •

Classic preparation technique Minimum preparation Partial preparation No preparation

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The classic preparation technique involves facial surface reduction, proximal surfaces reduction within the contact point, and the reduction of the incisal edge. Facial reduction needed is about 0.5–0.7 mm for maxillary teeth and about 0.3 mm for smaller teeth such as mandibular incisors. Sometimes it is necessary to reduce small portions of dentin, with the amendment that dentin reduction should be up to 50 %, but the terminal limit must be situated in the enamel. Average tooth tissue reduction is 0.5 mm, but the preparation is reduced toward cervical (0.3 mm) and more pronounced toward incisal (0.75 mm), the reason being the cervical enamel is thinner, and when possible the entire preparation must be located in the enamel. Preparation must respect the anatomy of the tooth, so it is performed in the different planes as the primary morphology of the tooth. Depth orientation grooves are placed in the facial surface between 0.5 and 0.7 mm with a cylindrical diamond bur. Orientation grooves are connected to achieve a uniform reduction. Chamfer margin is used and placed closely to the gingival crest in the enamel. Proximal preparation must be performed with care not to damage adjacent teeth and also should not neutralize the contact point, the preparation margins being placed at this level. There are some clinical situations involving reduction of proximal surface entirely: proximal cavities or fillings, diastema, or interdental space closure which involve changing of the gingival emergence profile of the tooth concerned. Usually it is necessary that the enamel on the margins of the preparation has to be removed using a chisel to avoid sharp edges that often result when using the bur. Contact area will be precisely reproduced on the definitive cast, which will facilitate its separation in mobile abutments without affecting the preparation. Incisal edge preparation can be done in four different ways (Fig. 1): • Window – the preparation finishes at 0.75 mm before the incisal edge; this preparation has the advantage that preserves natural enamel in the incisal edge but at the expense of a thin enamel at this level. This type of preparation is indicated for teeth with thicker incisal edge. Loosening or fracturing risks when biting or in end-to-end relationship are zero. The veneer edge will become visible if the incisal edge will wear off. • Feather – preparation is extended to the incisal edge but it is not reduced. This type of preparation is recommended for incisors with thin or narrow incisal edges, but there is a risk of fracturing or loosening in protrusion movements or in end-toend relationship. • Bevel – the inclination is facial palatal preparing the entire width of the incisal edge. The technique offers a very good control of incisal edge aesthetics, and it is indicated especially when there are small fractured incisal angles or edges. The type II occlusal contacts must be considered because they are not allowed to interfere with the junction between the tooth and veneer but must be placed at least 0.5 mm from the junction. • Overlap – involves preparation of the incisal edge, and then it is extended on the palatal surface and sometimes can extend all the way to the cingulum.

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Fig. 1 Incisal edge preparation techniques: (a) window, (b) feather, (c) bevel, and (d) overlap [39]

Of these, the latter two provide the greatest resistance to fracture under occlusal forces [19]. Window-type preparation is rarely used in indirect veneering, being most common in the direct methods of preparation. Minimal preparation and partial preparation targeting removal of trace amounts of dental tissue. In this case, before the clinician starts preparing the tooth, along with the dental technician, they must perform a diagnostic wax-up of the final outcome. Wax-up cast has to be in office before preparing the teeth. Through it’s impression, a silicone index is obtained which will then be used during preparation. A second index, of transparent silicone, is used to transfer the diagnostic wax-up to the oral cavity (mock-up). First the transparent index is used to create the mock-up on which one can analyze the position of future veneers, integrated in the dental arch. The transparent silicone index is filled with fluid composite and then positioned on the arch; after curing it is removed carefully and the mock-up is ready. Remove excess from gingival mucosa. To obtain the mock-up two techniques are used: 1. We can fill the index of transparent silicone with fluid material and then apply in the mouth and light cure it. This technique does not require a long time to yield good results.

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2. Fill one third of the transparent silicone index with fluid composite, apply it to the prosthetic field, then ligt cure it. Carefully remove the index and then the rest of the surface is filled with composite material; by this technique the clinician has more control over the final result but requires more time. Dr. Galip Gurel [20] calls this mock-up APT (aesthetic pre-evaluative temporaries), but it results from the impression of the wax-up, to provide a silicone key that will be used intraorally to the patient. It evaluates functional movements and static occlusion in order to see if any occlusal obstacles are present or premature contacts. It also analyzes the phonetics to exclude such problem in the future, related to pronunciation of certain words or phonemes. The technique consists in preparing the teeth with the mock-up so that the result is similar to natural teeth. It starts with a guide bur for creating guiding grooves both the facial surface and the incisal edge. Facial surface is prepared according to morphological anatomy of that tooth with a diamond bur with rounded tip, taking into account the fact that the preparation should not exceed the enamel-dentin junction. For incisal edge one of the classical preparation techniques can be used. In the end, finishing burs are used to smooth the preparation margins and remove sharp edges and also to finish the gingival and proximal margins. Once the preparation is ended, the depth is verified with the silicone index. Depending on how much the tooth surface was prepared, one can define if the preparation is minimal or partial. No-prep technique does not imply preparation of natural teeth. Decision for making veneers without preparation must be well instrumented by the clinician, both in accordance with the indications and contraindications of no-prep veneers and depending on the patient’s expectations. In general, the thickness of this type of restoration is very low, which makes solving dental discoloration or major change in tooth color not possible.

Indications of Full-Ceramic Veneers Full-ceramic veneers offer conservative solutions from biological point of view and a high degree of aesthetics for the following clinical situations: • Dental discoloration: tetracycline-discolored teeth, dental fluorosis, or discoloration due to age. • Defects of the enamel: different types of enamel hypoplasia and dysplasia, with birth defects, can be masked or corrected. • Diastema and interproximal spaces: closes interdental spaces, to achieve an aesthetically pleasing appearance. • Teeth with malpositions: it should be noted here that severe dental malposition is a contraindication for veneers, but minor malpositions such as dental rotations and mild crowding or spacing may be resolved by veneers, if orthodontic treatment is not desired by the patient. Otherwise, the orthodontic treatment is

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the best choice in these situations. In these situations, veneers can create an “optical illusion” of shape changing, tooth position modification, dimension, and tooth surface alteration. Malocclusion: configuration of palatal surface of the upper front teeth can be modified by veneering this surface to restore lost anterior guidance and type II occlusal contacts, for example, if a patient is periodontally compromised. Veneering will be used but after solving periodontal problems and if orthodontic treatment is not desired by the patient. Direct restorations failure: frontal teeth with multiple or repeated superficial fillings which are no longer suitable from aesthetic point of view. Incisal edge wear: different patients due to professional vicious habits (carpenter, tailor) or developed independently (pen or nails chewing) that produce incisal edge changes. Fractured teeth: they can be restored to normal morphology with partial or fullceramic veneers. Small teeth: frequently, veneering of small lateral incisors is the solution of choice for rendering patient’s smile; often in these cases veneering the entire frontal teeth group is necessary to preserve their relative size. Lateral incisor agenesis: often when lateral incisor is absent, the canine erupts in its place and thus creates an unpleasant aspect of smile. Remodeling the canine into a lateral incisor using a veneer reduces the ugly smile, but often in these cases intervention on central incisors and first premolars is necessary in order to obtain a high aesthetic result.

“Permanent” dental whitening: full-ceramic veneers are a good alternative when conventional teeth whitening procedures are contraindicated or when the patient wants a permanent change of tooth color. Masking gingival retraction: if some minimal tooth root is exposed, this defect can be masked by veneers.

Contraindications of Full-Ceramic Veneers Full-ceramic veneers are not indicated in following situations: • Massive crown loss: when a great amount of dental crown tissue is lost either by caries or trauma. • Diastema or interproximal space with mesial-distal enlarged space: when intraoral or on diagnostic cast measurements are suggesting that the space is too big. • Severe dental malpositions: ectopic teeth that needs orthodontic treatment. • Endodontically treated teeth: represents a relative contraindication for veneers from the dental tissue resistance point of view and because of the low color stability in time. In this case, a full-ceramic crown is the best solution. • Poor dental hygiene: patients who do not take care of their natural teeth and end up having a high caries level are not candidates for treatment with all-ceramic

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veneers in order to restore the aesthetics of natural teeth, because on the uncovered tooth surfaces, after veneering new caries may appear and compromise the veneer. • Bruxism: is a relative contraindication. Patients with bruxism, if they do not wear the mouth guard, destroy their natural teeth, the same thing happening with any kind of dental restoration, but in terms of a compliant patient (wearing mouth guard and respects the dentist indications) veneering is possible.

Full-Ceramic Veneers Advantages Compared with the composites or acrylics, ceramic materials through their proprieties are the optimal material for dental hard tissue replacement for the following reasons: • Color stability: ceramic materials offer a double advantage – on the one hand ability to reproduce any color or transparency of the natural teeth and on the other hand ability to maintain color in time. • Chemical adhesion to the enamel: after adhesive cementation by both veneer etching and dental tissue, by applying silanization agents on ceramic material and bonding agents on dental tissue, and, finally, by interposing between the two conditioned components a composite cement, a stable chemical bond of the veneer to tooth surface is obtained. • Periodontal health: the ceramic surface, after glazing, is very glossy and thus prevents the accumulation of dental plaque. Also, due to special ceramic aesthetics, the preparation margin can be placed above the free gingival margin (can be placed at 0.5 mm above the gum), and thus artificial cleaning and auto-pruning are possible. In case of no-prep veneers, however, the cervical margin will be finished until thickness becomes sometimes 0.1 mm, and in this case there is a potential risk of dental plaque retention at that level. • Abrasion resistance: resistance to wear and abrasion of ceramic materials is extremely high compared with composite resins and acrylic materials. Regarding opposing natural tooth abrasion, new ceramic materials, reinforced with lithium disilicate (IPS e.max – Ivoclar Vivadent) has a comparable hardness to that of the enamel and thus does not cause excessive wear of opposing teeth [21]. • Tear resistance: full-ceramic veneer itself is quite fragile, but once it is adhesively cemented to the enamel, the restoration develops a high resistance to tensile and shear forces. For example, ceramic materials reinforced with lithium disilicate have a fracture resistance of 400 MPa and after adhesive cementation may reach values of 1000 MPa [22]. • Resistance to fluid absorption: after glazing ceramic materials are inert to fluid absorption compared to composite resin or acrylic materials, where fluid absorption in the oral environment is a well-documented disadvantage. • Aesthetics: regarding ceramic materials, dental technician has much greater possibilities of obtaining the desired color and texture of the restoration surface

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compared with other materials. Porcelain can be colored both internally and externally (superficial), resulting in a restoration that mimics natural tooth appearance. Both texture and microtexture can be made on the veneer surface in order to simulate the appearance of adjacent teeth and can be maintained indefinitely after intraoral cementation. Light transmission: dental ceramic is the only translucent, fluorescent, and opalescent material that is able to mimic all the optical properties of natural teeth. Preservation of dental hard tissue: most of the time the preparation for fullceramic veneers is made exclusively in the enamel, with a reduction varied between 0.3 and 0.7 mm, with/without involving incisal edge. Thus dental veneers are a more biological method for restoring dental stain or dental fractures compared to crowns. Local anesthesia: if the preparation is strictly limited to the enamel, local anesthesia is not required, although there are more emotive patients requiring administration of anesthesia. Interim restoration: not required for no-prep veneers.

Disadvantages of Full-Ceramic Veneers • Time: indirect veneering is a sensible technique and, hence, time consuming. Any kind of negligence in the manufacture protocol, either in dental laboratory or in the office, can have a disastrous effect on the final restoration. • Fracture repair: after adhesive cementation, any kind of veneer repair actually means replacing it. • Color: it is both an advantage and a disadvantage. After adhesive cementation in the oral cavity, any color change is impossible. Considering that while natural teeth change color and ceramic materials not, practically patients with such restorations become addicted to whitening treatments if not all of the teeth in aesthetic area are restored. But this is true for any fixed or removable prosthetic restoration. • Fragility: veneers are difficult to handle due to their resistance to tear, especially those made of feldspathic ceramics. • The price: the price of a full-ceramic veneer can be equal to or even higher than that for a crown. In general, full-ceramic veneers failure is associated with one of the following factors: inobservance of this type of restoration by the dentist, poor preparation of the involved teeth, interim restoration, inobservance of manufacturing process by the dental technician, veneer handling, choice of cement and cementing technique, and patient communication. Full-ceramic veneer remains the prosthetic restoration that best completes mechanical, biological, and aesthetic proprieties today. This type of restoration avoids the use of metal structures or zirconium oxide and thus achieves excellent aesthetic qualities. It is also the most conservative type, capable of preserving

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significant amount of the natural tooth enamel. The high rate of treatment success by veneering indicates this technique as the solution of choice in dental aesthetics.

Longevity In a recent study (year 2012), Layton and Walton [23] watched for 21 years a number of 499 full-ceramic veneers manufactured with feldspathic ceramic, adhesively cemented to a total number of 155 patients. For case selection, endodontically treated teeth, molars, and teeth with periodontal problems were excluded. Before the adhesive cementation, all veneers were etched with hydrofluoric acid and then silanized. Teeth were prepared so that a minimum of 80 % of their surface would be covered with enamel. After an observation period of 21 years, they announce a survival rate of 96 %  2 %. Also, they noted that the loss of veneers is not random, but the same patient has lost more veneers, which implies the patient’s compliance both during treatment and especially after treatment, in the follow-up period. A literature review on survival rate of feldspathic veneers, published in 2013 by Walt and Conway [24], shows high survival rates of 95.7 % at 5 years and 95.6 % at 10 years. They have considered 11 published studies, both prospective and retrospective, and have concluded that feldspathic veneers have a high survival rate when they are adhesively cemented to the enamel. Taking into account the necessity of redoing all kind of full-ceramic restorations (inlay/onlay, veneers, and crowns), both before cementation and after intraoral cementation, Hekland, Riise, and Berg [25] published in 2003 a study which included 2069 feldspathic ceramic restorations (ColorLogic) and 1136 leucitereinforced ceramic restorations (IPS Empress 1 and 2). A 4.4 % percentage of the total restorations included in this study needed adjustments before cementation, and full-ceramic veneers were the most frequently adjusted restorations. After 2 years from cementation, the reconstruction rate was 1 % which indicates a general survival rate of the full-ceramic restorations of 99 %. The highest survival rate within the group was acquired by inlay/onlay restorations (99.8 %) and the lowest one by the full-ceramic crowns (98.4 %). Full-ceramic veneers had a survival rate of about 99 % at 2 years. Full-ceramic veneers can be used as a treatment method for patients with parafunctions (bruxism). Regarding this, Beier, Kapferer, Burtscher, and Dumfahrt published in 2012 a study realized on 84 patients, half of which (42 patients) had bruxism. The authors realized a total number of 318 full-ceramic veneers on maxillary and mandibular frontal teeth. The survival rate of the veneers from this group was 94.4 % at 5 years, 93.5 % at 10 years, and 82.93 % at 20 years. The authors’ conclusion was that full-ceramic veneers are an aesthetic, predictable, and successful treatment option for restoring frontal teeth. The fact that half of the patients had bruxism produced only a minor decrease of survival rate at 10 years comparable with the other studies published in the literature [26]. Using leucite-reinforced ceramic material for manufacture of full-ceramic veneers, namely, IPS Empress system (Ivoclar Vivadent), Fradeani presents a

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survival rate of 98.8 % at 6 years [27]. Same author together with Redemagni and Corrado finds a survival rate of 94.4 % at 12 years on a research group of 46 patients which received 182 cemented full-ceramic veneers. This survival rate is reached only when full-ceramic veneers are correctly adhesive cemented [28]. Lithium disilicate-reinforced ceramics have the highest fracture resistance (400 MPa). This is why the indications of this ceramic type include besides singletooth restorations (inlay/onlay, veneers, frontal, and lateral coverage crowns) even three-unit bridge in the frontal or bicuspid area, as well as implant restorations. Once lithium disilicate ceramic material appeared (IPS e.max Press – Ivoclar Vivadent) in 2002, realizing high aesthetic restorations becomes possible, which mimic natural teeth both in lateral area and on implants. A recent clinical study, published this year (April 2014), consists of a total number of 860 restorations like inlay/onlay, veneers, crowns, and implant restorations adhesively cemented in frontal and lateral area and reports a success rate between 95.39 % and 100 %, along with a survival rate between 95.46 % and 100 %. Restorations from this study were checked for a 3–6-year period. This study included a total of 312 patients, among whom were patients with parafunctions (bruxism), but periodontally compromised patients were excluded. Of the 860 fullceramic restorations, a total of 26 have developed complications after cementation: 17 partial fractures of the ceramic, five total fractures, and four unluted restorations. Conclusion of this study: all-ceramic restorations made of lithium disilicate-reinforced ceramic prove to be efficient and reliable both on short and medium term [29].

Therapeutic Protocol Dental aesthetic philosophy comprises a therapeutic algorithm to be covered by any dentist with the patient or when aesthetics is the main reason why the patient comes to the doctor. Depending on the expectations of the patient, the dentist may suggest different clinical procedures that have indications of choice in the case as follows: • Changes in color: the first option is the external coronal tooth whitening and/or internal. • Dental malpositions: the option of choice is the orthodontic treatment. • Shape/contour tooth modification: if there are small shape changes, they should be evaluated if they cannot be made by selective preparation followed by fluoridation. Only when the above options are excluded, the dentist will consider restoring teeth with aesthetic veneers or crowns.

The First Treatment Session Treatment plan for veneering, as in other prosthetic treatments, will start by checking indications and contraindications for the type of restoration, e.g., full-ceramic veneers. It is also very important to check patient’s buco-maxillary system’s

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Fig. 2 (a) Initial view of frontal maxillary teeth. (b) Initial view of dental arches

functionality: static occlusion, especially the anterior occlusal contacts, and protrusive and laterotrusive movements. Subsequently key elements in dental aesthetics must be evaluated: contour, position, and color of teeth. All these things will be examined and discussed in the first treatment session, along with patient’s examination. Following clinical examination,if the treatment solution is veneering the teeth, then, in the first appointment, impressions for study casts will be taken, along with occlusal relationship recordings which are necessary for mounting the casts in the articulator. All these records are sent to the dental technician for mounting casts in the articulator, programming it to reproduce mandibular movements of the patient, and producing a diagnostic modeling (wax-up). In addition, the dental technician receives a communication sheet from the dentist which contains all necessary changes that were established after discussion with the patient and eventually photos of dental arches (Fig. 2a, b) and the patient’s face and profile. Thus, the dental technician can achieve by wax addition on the cast of the final design of patient smile. Diagnostic wax-up comprises several design phases and encompasses different anatomic shapes of teeth according to facial aesthetic details but with patient wishes regarding personal appearance, without which diagnostic wax-up is doomed to failure [30]. Stages of wax-up for teeth in frontal area are: 1. Diagnostic wax-up should start with maxillary central incisors modeling. The process begins by creating a flat surface up to the new incisal edge and to the new length determined by the clinician. From the two maxillary incisors, a single tooth so-called unitooth is created, by adding wax to the facial area. This step is very important because once created, the new incisal edges parallel to a flat surface; thus the perpendicular interincisive line can be easily positioned. Positioning of the interincisive line can be checked by placing both maxillary

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and mandibular casts in the articulator, such that the maxillary interincisive line to correspond with the mandibular one. Once the interincisive line was placed, it should create an angle of 90 with the incisive edge. Subsequently check if interincisive line coincides with the center line of the face. Registration of the center line is realized in the first treatment session, and then it is transferred to the study casts and reported to the technician (T – reference). The next step includes central incisors anatomical modeling taking into account previous measurements. It is important that the shape of the central incisors be identical as a mirror image of each. Keeping mesial edge and the mesiovestibular angle of the facial area of central incisors slightly rounded, the overall look of central incisors will be a square, but if we round these limits, teeth will get a round look. In this stage dental technician can help himself with special measuring tools, such as special rulers. During the diagnostic wax-up process, it is very important that the technician evaluates the progress after each step. For a successful diagnostic wax-up, the dental technician must follow the rule “apply wax-stop-look-check-modify.” After the anatomical shape of central incisors is established, the technician goes to create the maxillary lateral incisors. The key point is that the technician determines the correct width of the lateral incisor in relation to central incisor. Unlike central incisors, lateral incisors should not be perfectly identical. Anatomical variations of lateral incisors occur frequently and are easily tolerated by those who observe. The technician must define the final shape of lateral incisors such that they are approximately symmetric. Wax is added to the canine to establish anterior guidance, lateral guidance, and aesthetics. Canines are generally symmetrical and light asymmetries at this level can be seen easily. If the technician places on a flat surface the wax-up, he can observe the incisal edges level and frontal teeth symmetry. At this point it is good for the technician to let aside the cast, and when he returns, the technician can see things that he could miss. Next steps include future occlusal relations, such as new guidance surfaces, which the technician can check by mounting the casts in articulator. Finishing details to the canine’s form; making the canine’s facial surface with a rounded shape unlike flat facial surfaces of incisors [31].

Once completed, diagnostic wax-up is sent to the dentist office (Fig. 3), where with a silicone key it is transferred to the patient’s mouth.

Second Treatment Session This operation is called mock-up and it is realized during the second appointment. So you get a preview of the final result and also the patient’s consent on the final shape of the teeth (Fig. 4a, b). Without any preparation of the patient’s teeth, the silicon index is filled with composite material used for provisional prosthesis and is applied intraorally. After setting of the provisional prosthetic material, the silicone index is

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Fig. 3 Diagnostic wax-up: final aspect

Fig. 4 (a) Mock-up dental arches view. (b) Maxillary mock-up view

removed, and both patient and physician can examine the final shape of future veneers. Transferring the diagnostic wax-up into the patient’s mouth implies growth of the predictability of final aesthetic result because it offers a preview of it. Any change desired by the patient or dictated by functionality of the stomatognathic system are possible to realize directly on the mock-up through finishing by the dentist. During this second treatment session, both functionality of stomatognathic system and aesthetic facial landmarks related with smile line, inferior lip line and superior lip line are examined, along with the patient comfort and speech with new teeth shape. Once all these conditions are met and the patient agrees to the final aesthetics of future restorations, the third treatment session is scheduled.

Third Treatment Session The objective of the third patient scheduling is to achieve the preparations needed for future veneers (Fig. 5a, b). The preparation type is selected according to clinical and aesthetic needs of each tooth. In general the dentist will remove 0.5–0.8 mm from the

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Fig. 5 (a) Prepared teeth facial view. (b) Prepared teeth occlusal view

enamel’s thickness in order to create space for the full-ceramic veneer and to obtain a good adhesion of the veneer to the tooth after adhesive cementation [32]. Care must be taken not to cut more than 0.5–0.8 mm especially from the proximal and cervical areas where enamel thickness is less, so the preparation should be done exclusively in the enamel as possible. Extending the preparation into dentine leads to a decrease of adhesion of the veneer to the tooth in the area, although dentine adhesives have experienced a significant evolution in the last period of time [33]. After completing the preparation, in the same treatment session, the dentist will make an impression of the preparations, followed by manufacture of composite or acrylic temporary veneers. The easiest method to make temporary veneers is direct technique of copying and transfer the wax-up to the prepared teeth, as in the case of mock-up. For a superior aesthetic result of provisional restorations, prefabricated polycarbonate veneers can be used along with composite or acrylic provisional material [34].

Fourth Treatment Session Next, the dental technician is the one who manufactures the full-ceramic veneers, in the laboratory. They will arrive in the dental office and so the fourth treatment session begins. This is the last session, when the adhesive cementation of all-ceramic veneers is realized. To obtain a predictable and sustainable result over time, special attention must be paid to the cementation session, because it involves both processing dental surface and the internal surface of the veneer in order to obtain an adhesive bond between ceramic and dental tissue.

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Adhesive cementation is responsible for the success or failure of the entire treatment, and thus the latter is more important in clinical stage. Depending on the number of veneers to be cemented, the best would be to start out cementation from the midline to distal, preferably the two central incisors should be simultaneously cemented. Once the veneers arrive in the dental office, they should be checked in the mouth. For this maneuver, glycerine gel can be used in order to maintain the veneers on the tooth. Some manufacturers of adhesive cements provide a glycerin gel, color coded as the final cement, which gives you a preview of the final cementation (e.g., Variolink Veneer – Ivoclar Vivadent). The final color of ceramic restorations is influenced by underlying dental tissue color and by default of the cement’s color, as ceramic materials are translucent. Thus, this glycerine test gel (try-in gel) becomes mandatory to use before cementation when you wish to obtain a maximum aesthetics of the final result. Their application on the dental tissue surface before adhesive cementation brings no harm to the final veneer adhesion if their removal protocol is followed, namely, water wash and drying the tooth because the gels are water soluble [35]. Test phase should include phonetic and functional tests and patient’s views on the aesthetic outcome of the result. After the try-in phase has ended and glycerin gel was cleaned, cementation of veneers is realized. It starts with conditioning the internal surface of the ceramic veneer using hydrofluoric acid. Time and concentration of hydrofluoric acid vary depending on the type of ceramic used (e.g., hydrofluoric acid used for IPS e.max Press ceramic has a concentration of 4.9 % and the recommended action time is 20 s). Hydrofluoric acid is known to selectively dissolve the crystals and the glass component causing an uneven and porous surface. The microporosity of the ceramic surface increases the adhesion area and offers a micromechanic adhesion to the resin component of the adhesive cement system. Density of leucite crystals found on the surface influences the formation of microporosity in the etching stage with hydrofluoric acid. Leucite crystals are better dissolved unlike the glass component, in hydrofluoric acid. Microporosity formed after acid application increases the adhesion surface and thus increases the success rate of the adhesive bond [36]. After etching the internal surface of the ceramic veneer, a hydrofluoric acid neutralizing agent can be used, which will be removed from the veneer together with the acid by washing in running water or in an ultrasonic bath using a mixture of alcohol and water. Ultrasonic cleaning is the most effective method of removing debris, resulting in the internal surface of veneer after etching [37]. After etching the ceramic material will be silanized. The silanization process has the ability to decrease the superficial tension of the internal surface of the veneer. Silane is a coupling agent that links inorganic particles on the surface of ceramic and organic matrix from the adhesive cementation system. After the silanization agent was applied, wait one minute while evaporating ethanol/alcohol used as a solvent in silane agent, and optionally one can apply a bonding – Heliobond (Ivoclar Vivadent). This adhesive protocol is completed for the ceramic restoration. Adhesive protocol for dental surfaces begins with their wash. For this stage, use solution containing chlorhexidine and/or benzene, followed by washing under

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running water and drying to remove debris from the surface of the tooth. Full adhesive protocol for dental surfaces runs under isolation. Isolation can be done with a rubber dam or special devices (OptraGate) whose purpose is to prevent the saliva to reach the tooth surface. Once isolation is done by these methods, soft parts such as lips and cheeks are automatically removed, facilitating intraoral handling of the veneer. If a good control of gingival crevice fluid is desired, we can use gingival displacement cord placed before the application of the rubber dam, positioned carefully in order to avoid the gingival sulcus bleeding. Before cementing veneers, interdental celluloid strip are positioned between teeth, for a better control over the excess of the adhesive system and cement that can suppress the neighboring teeth. The next stage is the demineralization. Orthophosphoric acid of 37.5 % is applied for 40 s onto the enamel and 5–15 s onto dentin. Remove with a water jet and air followed by air drying, avoiding desiccating dental tissue. After etching adhesive system of the cement is applied, used according to the manufacturer. Often it is the successive application of a primer and bonding with dual cure. In the gingival third of the tooth surface, there is a risk that the preparation would penetrate the dentin. In this case it is recommended to apply a sealer to the exposed dentin area, which protects the dentin area and decrease sensitivity after cementation. According to the literature, areas of exposed dentin onto preparations surface cause a decrease in adhesive bond strength between ceramic and dental tissues, which leads to a decrease in the success rate of veneers in such cases [38]. Silane agent is applied, both the tooth surface and the internal surface of the veneer, using an applicator or brush for adhesives, and the excess is removed with the airflow. Next apply cement on both the tooth surface and the veneer, taking care not to form bubbles in the cement mass. Then position the veneer onto preparation and insert it with moderate pressure so as to avoid accidents like fracture. The most important factor in this situation is the cement viscosity: as the viscosity is lower (cement is fluid), the lower the risk of fracture. After the veneer is positioned on the tooth, it is light cured for 5 s, and then the excess cement is removed. Final light curing is about 20 s from each side of the tooth; thus the veneer is finally fixed on the tooth (Fig. 6a–c). If there is excess cement after light curing especially in the gingival area, it can be removed with: dental floss, hand instruments, or ultrasonic instruments. Following the steps described above, in just four sessions at the dentist, a high aesthetic result of teeth can be achieved, with high predictability. After final cementation the patient is enrolled in a program of dispensarization, with regular checkups. With the advent of newer conservative treatments such as vital and unvital teeth whitening, with orthodontic treatments that provide a high degree of dental aesthetic during treatment (In Visa Line), indication for veneering could be reduced. Porcelain veneers are still commonly used as a way to make cosmetic changes for teeth that are discolored, worn out, or fractured or to align teeth. With the advent of new ceramic materials and developing adhesive techniques, porcelain veneers are considered to be a reliable option, with high aesthetics and a good long-term prognosis.

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Fig. 6 (a) Intraoral view of the full-ceramic veneers. (b) Full-ceramic veneers occlusal view. (c) Final view of the veneered teeth

Summary With the development of dental materials, treatment options are wide: in addition to aesthetic fillings or direct composite resin veneers, now are available full-ceramic crowns, ceramic veneers, and ceramic inlays. The commonly used material for aesthetic restorations is dental ceramics, due to its color stability, biocompatibility, mechanical properties, and excellent aesthetic results. There are three types of ceramics used for manufacturing these veneers: feldspathic ceramic, leucite reinforced ceramic, and glass ceramic reinforced with lithium disilicate. These materials have particular aesthetic properties, mimicking natural tooth look, but their mechanical properties are different. Because fullceramic veneers have very low thickness, mechanical strength of the material is important for their handling during manufacturing or cementation process and especially for preserving the integrity of the restoration. Teeth preparation techniques for ceramic veneers varies, depending on the aesthetic needs of the patient, type of veneers to be made, and their indications or

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contraindications. The operative protocol includes four treatment sessions at the end of which the patient has the veneers cemented over his teeth. There are many advantages for veneering natural teeth and disadvantages are low compared to their benefits.

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Cena Dimova, Biljana Evrosimovska, Katerina Zlatanovska, and Julija Zarkova

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Aim of Alveolar Augmentation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Quality and Quantity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Quality After Tooth Extraction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Techniques to Preserve the Bone After Tooth Extraction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Grafts and Donor Location . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Autogenous Bone Graft (Autografts) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Types of Autogenous Bone Graft . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Allogenic Bone Grafts (Allografts) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Xenogenic Bone Grafts (Xenografts) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Donor Location (Site) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Augmentation Techniques and Material (Horizontal, Vertical) . . . . . . . . . . . . . . . . . . . . . . . . Clinical and Radiological Assessments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Augmentation Procedures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Substitutes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Calcium Phosphates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Beta Tricalcium Phosphate (βTCP) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Nanosized HAP . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Synthetic Hydroxyapatite (HA) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Coralline Hydroxyapatite . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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C. Dimova (*) • K. Zlatanovska • J. Zarkova Faculty of Medical Sciences, Dental Medicine, Macedonia FYR, University “Goce Delcev” – Stip, Stip, FYR Macedonia e-mail: [email protected]; [email protected]; [email protected]; [email protected] B. Evrosimovska Faculty of Dentistry, Macedonia, FYR, University “Sts. Cyril and Methody” Skopje, Skopje, FYR Macedonia e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_51

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Calcium Sulfate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Glasses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Glass Ionomers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Combined Synthetic Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Mechanism of Bone Regeneration After Ridge Augmentation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . The Terms “‘Guided Bone Regeneration” (GBR) and “Guided Tissue Regeneration” (GTR) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Guided Tissue Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Guided Bone Regeneration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Membrane . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Nonresorptible Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Resorptible Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Collagen Membranes . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Sinus Floor Augmentation, Bone Splitting, Distraction Osteogenesis . . . . . . . . . . . . . . . . . . . . . . . . Maxillary Sinus Floor Elevation – Lateral Window Technique . . . . . . . . . . . . . . . . . . . . . . . . . . . Transalveolar Sinus Floor Elevation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bone Splitting/Expansion and Immediate Implant Placement and Split-Ridge Techniques with Interpositional Bone Grafts and Delayed Implant Placement . . . . . . . . . . . Distraction Osteogenesis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Bone retention, bone augmentation, and bone regeneration are central topics in oral surgery, implantology, and periodontology. Bones and teeth are the only structure within the body where calcium and phosphate participate as functional pillars. Despite their mineral nature, both organs are vital and dynamic. The major sequel from human tooth loss is the loss of alveolar bone. After tooth extraction, the residual alveolar ridge generally provides limited bone volume because of ongoing, progresive bone resorption. The process of healing on bone defect in the region of alveolar ridge passes throught several stages from the coagulum formation to the mature lameral bone. The healing process within postextraction sites reduces the dimension of the socket over time. Bone grafts and bone graft substitutes support regeneration in bone defects and can be used for bone augmentation. Bone graft substitutes are clasiffied by their origin as autogenous bone grafts, bone graft substitutes (allogenic from human origin and xenogenic materials from animal origin), and synthetic (alloplastic) bone graft substitues, manufactured from mineral raw materials, whose composition is precisely defined and whose availability is is unlimited. Alveolar ridge augmentations are classified according to their morphology and severity. Bone augmentation techniques can be used for the application of socket defect grafting, horizontal ridge augmentation, vertical ridge augmentation, and sinus augmentation. Ridge augmentation methods are therefore very important developments and have so far been promising especially in view of the fact that life is incresingly prolonged especially in economically well-developed countries and the incidence of the disease is expected to further increase in the future.

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Keywords

Tooth extraction • Socket • Augmentation • Alveolar ridge • Bone regeneration • Grafting • Bone grafting techniques • Bone graft substutes • Ridge preservation • Autogenous • Alloplastic • Horizontal ridge augumentation • Vertical ridge augumentation • Guided bone regeneration • Guided tissue regeneration

Introduction Immediately after tooth extraction the bony walls of the alveolus present significant resorption, the central part of the socket is partly filled up with woven bone, and the extraction site becomes markedly reduced in size. Edentulous site diminishes in all dimensions, i.e., buccal-lingual, buccopalatal, and apical-coronal. At the same time, the soft tissues in the extraction site undergo adaptive changes that clinically may appear as deformations of the jaw. In health, the different structures of the alveolar process, the cortical and cancellous bone, are constantly undergoing remodelation in response to functional forces acting on the teeth. Once when teeth are lost, the attachment apparatus is destroyed, and the alveolar process, mainly the alveolar ridge, undergoes significant structural changes; these are referred to as disuse atrophy [1]. Alveolar ridge atrophy after loss of teeth occurs secondary to advancing age, to deterioration of general health, to systemic or metabolic diseases, and due to occlusion defects or to denture pressure. The condition causes serious problems for both the dentists and the patients. The toothless mandibular resorption or the high muscular attachments caused by senile atrophy produce unsuitable conditions for total denture. Resorption of the edentulous or partially edentulous alveolar ridge or bone loss due to periodontitis or trauma frequently compromises dental implant placement in a prosthetically ideal position. Therefore, augmentation of an insufficient bone volume is often indicated prior to or in conjunction with implant placement to attain predictable long-term functioning and an aesthetic treatment outcome [2]. The basic knowledge of alveolar augmentation included the use of bone grafting material with respect to the principles of biological mechanism. To optimize therapeutic approaches to bone augmentation and regeneration principles of osteogenesis, osteoconduction, osteoinduction, osteointegration, osteopercepcion, and osteopromotion can be used [3]. Osteogenesis – this term means that primitive, undifferentiated, and pluripotent cells are somehow stimulated to develop into the bone-forming cell lineage. One proposed definition is the process by which osteogenesis is induced. Osteogenesis has been described as the direct transfer of vital cells to the area that will regenerate new bone. Osteoconduction – includes the principle of providing the space and a substratum for the cellular and biochemical events progressing to bone formation. The term means that bone grows on a surface. An osteoconductive surface is one that permits bone growth on its surface or down into pores, channels, or pipes. Osteoconduction

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is a process by which bone is directed to conform to a material’s surface. The space maintenance requirement for many of the intraoral bone augmentation procedures allows the correct cells to populate the regenerated zone. Osteoinduction – embodies the principle of converting primitive, undifferentiated, and pluripotential mesenchymal-derived cells along an osteoblast pathway with the subsequent formation of bone. This term means that pluripotent cells are somehow stimulated to develop into the bone-forming cell lineage. This concept was established in 1965, with heterotopic ossicle formation induced by the glycoprotein family of morphogenesis known as the bone morphogenetic proteins (BMPs). A bone graft material that is osteoconductive and osteoinductive will not only serve as a scaffold for currently existing osteoblasts but also can trigger the formation of new osteoblasts, theoretically promoting faster integration of the graft. The most widely studied type of osteoinductive cell mediators are bone morphogenetic proteins (BMPs) [4]. Osteoperception – is the term used to describe the ability by patients with osseointegrated fixtures to identify tactile thresholds transmitted through their prostheses. It is a phenomenon of importance in both dental and orthopedic applications of osseointegration. The identification of osteoperception as a phenomenon of osseointegration was the result of work carried out in the dental sciences by Torgny Haraldson. Osteopromotion – involves the enhancement of osteoinduction without the possession of osteoinductive properties. For example, enamel matrix derivative has been shown to enhance the osteoinductive effect of demineralized freeze dried bone allograft (DFDBA) but will not stimulate from the new bone growth alone [5]. Osseointegration – Brånemark [6] introduced the term “osseo integration” to describe this modality for stable fixation of titanium to bone tissue. Osseointegration was originally defined as a direct structural and functional connection between ordered living bone and the surface of a load-carrying implant. It is now said that an implant is regarded as osseointegrated when there is no progressive relative movement between the implant and the bone with which it has direct contact. In practice, this means that in osseointegration there is an anchorage mechanism where no vital components can be reliably and predictably incorporated into living bone and that this anchorage can persist under all normal conditions of loading. Osseointegration provides an attachment mechanism for incorporation into living bone of nonvital components made of titanium. As a biological phenomenon it has been amply demonstrated and clinically tested, and it is now widely accepted. The present range of clinical applications is – In the field of oral surgery worldwide, more than 800,000 patients have been treated since 1965 until now with osseointegration dental reconstructions, according to Brånemark. The results indicate superiority over conventional prosthodontics, with respect to long-term success rates. – Facial prosthesis (extraoral applications of osseointegration include anchorage for craniofacial prostheses including ear, eye, and nose) finger prosthesis, etc.

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Aim of Alveolar Augmentation Atrophy of the alveolar ridge may cause aesthetic and surgical problems in prosthetic dentistry, especially when implant treatment is planned. After the loss of teeth atrophy of the alveolar processes occurs in a vertical as well as a horizontal plane. The term atrophy is defined in the dictionary as "a wasting away; a diminution in the size of a cell, tissue, organ, or part" [7]. This process is starting and continuous throughout life because of the lack of stimuli seen on alveolar process of the jaws. The importance of teeth for jaw bone health is extensively exploited in the contemporary scientific literature. When one or more teeth are missing, it can lead to jawbone loss at the site of the gap. This loss of jawbone can develop into additional problems, both with the patient’s appearance and overall health. Natural teeth are embedded in the jawbone and stimulate the jawbone through activities such as chewing and biting. Immediate alveolar ridge prophylaxis after tooth extraction includes – Preservation of the alveolar process by retention of endodontically treated roots (physiologically most accepted) – Immediate implant placement – Guided bone regeneration – The use of root analogues [8] Ethiopathogenesis of atrophy of alveolar ridge is teeth loss followed by loss of supporting alveolar bone which leads to alveolar defects, usually starts along the labial surface of the alveolar crest which results with loss of alveolar width, 40–60 % of bone loss within first 36 months with decrease to 0.25–0.5 % annually after 3 years, pneumatization of sinus, as well as periodontal disease associated with increased postextraction bone loss. When teeth are missing, the alveolar bone, or the portion of the jawbone that anchors the teeth in the mouth, no longer receives the necessary stimulation and begins to break down, or resorb. The body no longer uses or “needs” the jawbone, so it deteriorates. Without intervention (natural healing), the results of all nine studies showed a significant loss of ridge width (2.6 to 4.6 mm), and the results of five studies showed a statistically significant loss of bony ridge height (0.55 to 3.3 mm). The aims of alveolar ridge augmentation are to 1. Restore the function of the jaw in anterior, posterior, vertical, and lateral dimensions 2. Increase the bone tissue in cases where the lower jaw has been atrophied 3. Create an optimal support for dentures and better distribution of the jaw's functional forces 4. Fulfill the biomechanical requirement of the prosthesis 5. Restore the intermaxillary ridge relationship 6. Rehabilitate the dentures for efficient functioning and to produce better facial aesthetics

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Ensure good aesthetic results Re-establish adequate amount of bone volume for implant placement Provide biological acceptance of implants or transplants Obtain healthy bone to ensure osseointegration and survival of the implant

There seems to be no uniformity of opinion about which of the available methods provide the best anatomical and functional results. Among the produces proposed to restore the alveolar ridge, bone grafts were the first to be popularized. Kruger who favored this method recommended iliac grafts. Although costal grafts can perhaps better be adjusted to the mandibular arch, there can occur 50 % loss due to contraction. These results are akin to those of Steinhouser and Obwegeser who concluded that significant amount of atrophy and defects are observed of the mandibular or on the maxilla after bone grafting. Other studies have reported satisfactory results in general for treatment of atrophic ridge using hydroxyapatite with lesser percentage of neural injuries. Postoperative ridge resorption is observed only in 4–10 % of cases, a figure which compares favorably with other procedures aiming to correct alveolar ridge atrophy.

Bone Quality and Quantity Dental implants have become the most popular and reliable treatment option for restoring missing teeth. Since early times of implantation era preoperative studies include incision of gingiva in order to get a view of the bone surface. Preoperative studies are required because a jawbone must offer proper quality and adequate quantity of the bone. The overall dental implant success rate is considered to be influenced by both the volume (quantity) and density (quality) of available bone for implant placement (a great extent on the volume and quality of the surrounding bone). Therefore, it is important to measure the alveolar process precisely so that the proper system may be chosen. There are number of classifications suggested for assessment of the degree of atrophy of partially or fully edentulous jaws. Lekholm and Zarb [9] classify quality of residual alveolar bones into four types: – – – –

Type 1 Type 2 Type 3 Type 4

= = = =

large homogenous cortical bone thick cortical layer surrounding a dense medullar bone thin cortical layer surrounding a dense medullar bone thin cortical layer surrounding a sparse medullar bone

Classification of quality of residual alveolar bones indicates a good correlation with bone mineral content. The system for bone quality assessment with three classes, dense, normal, and soft bone, is proposed. Bone quality is classified into four groups according to the proportion and structure of compact and trabecular bone tissue: groups 1–4 or type I to IV (Fig. 1 Bone Quality Index-BQI) [10]. Also, it is important to know the bone quantity and quality of the jaws when planning implant treatment. Bone quantity of jawbone is classified into five groups

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Fig. 1 Bone quality index. 1- Type I: homogeneous cortical bone; 2. Type II: thick cortical bone with marrow cavity; 3. Type III: thin cortical bone with dense trabecular bone of good strength; 4. Type IV: very thin cortical bone with low-density trabecular bone of poor strength

(from minimal to severe, A–E), based on residual jaw shape and different rates of bone resorption following tooth extraction. During all stages of atrophy of the alveolar ridge, characteristic shapes result from the resorptive process. It is difficult to obtain implant anchorage in bone that is not very dense. Sufficient bone density and volume are therefore crucial factors for ensuring implant success [9, 10]. However, this classification, like many others, described changes only of jaw shapes in general and failed to indicate precise measurements. Alveolar ridge deformities are classified according to their morphology and severity. This classification has been done to standardize communication among clinicians in the selection of reconstructive procedures designed to eliminate these defects. • A class I defect has buccal-lingual loss of tissue with normal ridge height in an apicocoronal direction. • A class II defect has apical-coronal loss of tissue with normal ridge width in a buccal-lingual direction. • A class III defect has a combination of buccal-lingual and apical-coronal loss of tissue resulting in loss of height and width. Critical-sized alveolar ridge defects in the horizontal and vertical dimensions may occur following tooth loss, fractures, or different pathological processes. Such defects may compromise the ideal implant placement as prescribed prosthetically with an unfavorable outcome.

Bone Quality After Tooth Extraction Since the beginning of the twentieth century, the concept of osteoconduction in bony changes in the oral cavity showed a wide range of biomaterials and their osteoinductive potential that emerged gradually and has to a large extent improved

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the quality of the bone prior to the placement of an implant. Alveolar bone loss is a major concern after tooth extraction in patients, and therefore atraumatic extraction procedures should be followed to avoid further bone loss. To overcome the alveolar bone loss and to augment support for placing dental implants, many bone regenerative substitutes are available such as allografts, autografts, xenografts, synthetic biomaterials, and osteoactive agents [11]. Tooth extraction is one of the most widely performed procedures in dentistry today, and it has been historically well documented that this procedure may induce significant dimensional changes of the alveolar ridge. The dilemma that clinicians face is how to manage tooth extractions to provide for the future placement of a dental implant or to maximize ridge dimensions for the fabrication of a fixed or removable prosthesis. If performed inadequately, the resulting deformity can be a considerable obstacle to the aesthetic, phonetic, and functional results that both our patients and we clinicians expect at this current time [12]. Severe resorption of the maxillary alveolar crest presents a more demanding situation for the restorative team. Thus, it would be valuable to assess outcomes for this immediate loading treatment protocol in subjects with marked maxillary alveolar crest atrophy. Immediate loading of implants in the edentulous maxilla has previously been successfully performed and reported [13]. While the ability of various grafting materials to preserve extraction socket morphology has been adequately reviewed, the quality of the grafted bone in the socket is not as well understood. Based on a limited number of prospective comparative studies, the use of grafting materials for socket augmentation might change the proportion of vital bone in comparison to sockets allowed to heal without grafting. Whether these changes in bone quality will influence implant success and periimplant tissue stability remains unknown. In light of the steady progress in bone grafting techniques and graft materials, it has become possible to improve the volume, width, and height of bone in deficient areas of the oral cavity. These advances in regenerative dentistry thus facilitate an easy and convenient placement of an implant in an ideal position and angulations resulting in superior aesthetics and function. Bone grafting materials and their substitutes are the alternative filler materials, which facilitate to reduce additional surgical procedures, risks, or chances of cross-infection involved in placing autografts and allografts into the bony structures. This review literature highlights various biomaterials that are helpful in bone healing and thus creates an anatomically favorable base for ideal implant placement [11].

Techniques to Preserve the Bone After Tooth Extraction Several techniques can be used to preserve the bone and minimize bone loss after the tooth extraction. Immediate alveolar ridge prophylaxis after tooth extraction includes preservation of the alveolar process by – Retention of endodontically treated roots (physiologically most accepted) – Guided bone regeneration

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– Immediate implant placement – Use of root analogues In one common method, the tooth is removed and the socket filled with bone or bone substitute. It is then covered with gum, artificial membrane, or tissue stimulating proteins to encourage the body’s natural ability to repair the socket. With this method, the socket heals eliminating shrinkage and collapse of surrounding gum and facial tissues. The newly formed bone in the socket also provides a foundation for an implant to replace the tooth using autografts, allografts, xenografts, guided bone regeneration (GBR), and growth factors which are used with varying degrees of success in an effort to maintain the anatomical dimensions of the alveolus before implantation. There is suggestion for strategy of immediate implant placement into an extraction socket and simultaneous GBR. Guided bone regeneration techniques and the use of bone replacement materials have been shown to enhance socket healing and potentially modify the resorption process. Also, there are various recommendations regarding timing of implant placement after tooth extraction. The implant can be placed immediately following the extraction during the same surgical procedure (immediate implant placement), following a delay of 2–6 weeks (late implant placement), or following a delay of 3–6 months (delayed implant placement). Today, the combination of anatomically oriented implant designs, new biomaterials such as zirconia ceramics, and surface technologies has resulted in dental implants that are specially designed to replace each individual tooth. The use of root analogues as preimplant therapy can provide adequate quantity of bone and soft tissue for implant placement. Many authors showed that different bone substitute materials had been used as root analogues, some of them being dense hydroxyapatite, polyglycolic acid, polylactic acid, bioabsorbable polylacticpolyglycolic acid (PLGA), deproteinized bovine bone mineral integrated in a 10 % collagen matrix, β-tricalcium phosphate (β-TCP) combined with type I collagen and β- TCP/PLGA [14], and β-TCP coated with PLGA root analogue [8]. The development of new medical technologies enables use of achievements in material science, biochemistry, molecular biology, and genetic engineering while creating new combined synthetic materials for bone grafting (osteoplasty). Modification of their bulk structure, which brings their structure closer to natural bone tissue, including cytokine growth factors and morphogens into their composition, enables to provide synthetic materials with not only osteoconductive but also osteoinductive properties. This also enables control of the speed of biodegradation, bringing it closer to the kinetics of osteogenesis.

Bone Grafts and Donor Location Grafting refers to a surgical procedure when transplanting a tissue from one site to another on the body, or from another person, without bringing its own vascularization with it. Instead, a new blood supply grows in after it is placed. A graft also can

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be an artificially manufactured substance that has similar chemical composition with the tissue that replaces it. The main indications for bone grafting are – – – – – –

Socket preservation Ridge augmentation Defects following cyst removal/apicoectomies and periodontal defects Sinus lifts Distraction osteogenesis Implants placing

Bone grafting can be used for a localized or more generalized alveolar ridge augmentation like correction of intermaxillary relationships. This procedure is done by adding the patient’s own bone from a secondary location, or utilizing an organic or inorganic material like a bone substituent from another patient, species or artificially made. According to their origin there are four different types of bone grafts: – – – –

Autografts Allografts Xenografts Alloplasts – synthetic bone substitutes

Bone grafts in general can be made like bone blocks or particulates, in order to be able to adapt it better to a defect, and can be applied with inlay technique. Ideal bone graft material should have all properties like osseointegration, osteoconduction, osteoinduction, and osteogenesis. Osseointegration is the ability to chemically bond to the bone surface without scar; Osteoconduction means which provides matrix for bone growth, and support growth of bone over its surface. Osteoinduction is the characteristic when the graft can induce differentiations of pluripotential stem cells from surrounding tissue to an osteoblastic phenotype. Osteogenesis means directly producing new bone by osteoblastic cells present within the graft material. The ideal characteristics of a bone graft material have been described by Hammerle 1999 [15] as follows: 1. 2. 3. 4. 5. 6. 7.

Sterile Nontoxic Nonimmunogenic Osteoconductive or osteoinductive Favorable clinical handling Resorption and replacement by host bone Synthetic

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8. Available in sufficient quantities 9. Low in cost.

Autogenous Bone Graft (Autografts) Many clinicians as a first choice for alveolar augmentation use autogenous bone graft. It is considered as a “gold standard” by which all other graft materials are judged because of its osteogenic, osteoconductive, and osteoinductive properties [15]. According to the potency of their best features various types of autografts exist like cancellous, vascularized cortical, free vascular transfers (nonvascularized cortical), cortical-cancellous bone grafts, and bone marrow aspirates. The advantages that make this graft great are avoided rejection reactions and the possibility of antigenicity. But there are also some disadvantages like low patient acceptance of autogenous bone harvesting given the potential morbidity associated with such techniques, chronic postoperative pain, hypersensitivity, infection, etc. There are also other limitations as to how much bone tissue can be harvested, and harvesting requires an additional surgery at the donor site. For this kind of grafts bone development is occurring in two phases. The first one lasts almost 4 weeks. During this phase the living cells from the graft contribute to bone formation, and during the second, host cells are taking the main role. The endosteal lining cells and marrow stromal produce more than half of the new bone, whereas osteocytes make a small (10 %) contribution. Free hematopoietic cells of the marrow make a minimal contribution in this process.

Types of Autogenous Bone Graft Cancellous bone is the most commonly used source of autogenous bone graft. Due to the porous trabeculae lined with functional osteoblasts, good space filler can be easily revascularized; in other words it is a very effective osteogenic and osteoconductive material. A small quantity of growth factors in this bone graft can result in high osteoinductive power. Even though the initial support is not so strong, after some time the graft material is rapidly incorporating and ultimately achieves strength equivalent to that of a cortical graft after 6–12 months. When the implantation is finished there are still life donor osteocytes, and combined with graft porosity and local cytokines they promote angiogenesis and host mesenchymal stem cell recruitment. These recruited mesenchymal stem cells have the potential to differentiate into osteoblasts; thus, the graft may be fully vascularized within 2 days. New bone formation is observed within a few weeks and typically is remodeled by 8 weeks, with complete graft turnover by 1 year. This turnover occurs by the process of creeping substitution, defined as concomitant osteoblast deposition of new osteoid and osteoclast resorption of necrotic donor trabeculae.

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Cortical bone grafts are less biologically active compared to cancellous bone, less porous, with less cellular matrix, prolonged time to revascularization, but are more compact and can be used to provide better initial structural support when needed. Cortical grafts serve as highly osteoconductive material with minimal osteoinductive and osteogenic properties. The dense cortical matrix results with relatively slow revascularization and incorporation, as resorption must occur before deposition of new bone, and limited perfusion and donor osteocytes make this option poorly osteogenic. Although no vascularized cortical grafts provide immediate structural support, they become weaker than vascularized cortical grafts during the initial 6 weeks after transplantation, as a result of resorption and revascularization.

Vascularized Cortical Bone Grafts In order to improve graft incorporation and healing, cortical and corticocancellous grafts can be harvested with a vascular pedicle Corticocancellous Bone Graft This type pf bone grafts offers the advantages of both cortical and cancellous bone: an osteoconductive medium and immediate structural stability from cortical bone and the osteoinductive and osteogenic capabilities of cancellous bone. Bone Marrow Bone marrow grafts are highly osteoblastic materials due to stem cells found in bone marrow. Injections of autologous bone marrow provide a graft that is osteogenic and potentially osteoinductive through cytokines and growth factors secreted by the transplanted cells.

Allogenic Bone Grafts (Allografts) Allogenic bone is a nonvital osseous tissue taken from one individual and is transferred to another individual of the same species. Bone derived from cadavers has been widely used in implant dentistry as well as periodontology. These types of grafts are harvested from donors with good documented medical history, tested for safety from all common infectious diseases. Usually we use three forms of allogenic bone: fresh frozen, freeze dried bone allograft (FDBA), and demineralized freeze dried bone allograft (DFDBA) [11]. The type of autografts can be used as demineralized bone matrix, morselized and cancellous chips, corticocancellous and cortical grafts, and osteochondral and whole-bone segments. Demineralized bone matrix acts as an osteoconductive material providing a matrix for new bone formation and growth, and should be resorbed as part of the normal turnover of bone, but some particles appear to remain intact for some time after the graft has been placed and possibly as an osteoinductive material. The grafts are produced as particulates with a reasonably uniform grain size or as sheets and large blocks. Demineralized bone matrix revascularizes quickly. It also is a

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suitable carrier for autologous bone marrow. Demineralized bone matrix is prepared by a standardized process in which allogeneic bone is crushed or pulverized to a consistent particle size (74–420 m) followed by demineralization in 0.5 N HCL mEq/g for 3 h. Current methods of processing demineralized bone matrix follow the same basic steps, but refinements of the technique, many of which have been patented, have been developed by several companies and tissue banks. The biological activity of demineralized bone matrix is presumably attributable to proteins and various growth factors present in the extracellular matrix and made available to the host environment by the demineralization process. Fresh bone allografts are highly antigenic and have limited time to test for immunogenicity or diseases.

Xenogenic Bone Grafts (Xenografts) These grafts are harvested from animals, usually cows. That is why this is processed to make it sterile and totally biocompatible. Advantages of xenografts: – Only one procedure is needed as the bone is not being harvested from the patient. – Natural bone growth is encouraged. Disadvantage of xenografts is the minimal risk of bovine sponge form encephalopathy due to the fact that all organic components of the bone are extracted. The most widely used xenograft bone is deproteinized bovine bone mineral like Bio-Oss (Geistlich Pharma, Wolhusen, Switzerland). Bio-Oss has similar properties to human cancellous bone, in its macrostructure and its crystalline content; it also has similar physical properties to the human bone. This is a purely mineral graft and is osteoconductive but also can go through some resorption, so its use also has limitation. When used in a particulate form it is mixed with the patient’s blood and packed into the defect. Some authors [11, 15] have described improved results when combining this with a membrane to protect the blood clot. Bio-Oss Collagen is the latest version of this material combining bone mineral and collagen to produce a block of material that can be carved to shape or used as a particulate graft.

Donor Location (Site) Autogenous bone grafts have been used for many years for ridge augmentation and are still considered the gold standard for jaw reconstruction. The use of autogenous bone grafts with osseointegrated implants originally was discussed by Brånemark and colleagues, who often used the iliac crest as the donor site. Other external donor sites include calvarium, rib, and tibia. For repair of most localized alveolar defects, however, block bone grafts from the symphysis and ramus buccal shelf offer advantages over iliac crest grafts, including close proximity of donor and recipient sites, convenient surgical access, decreased donor site morbidity, and decreased cost.

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There are several donor sites that are available for autogenous bone harvesting. Grafts can be taken from intraoral and extraoral sites, depending of the indications and the volume of defect for augmenting, as well as advantages and disadvantages of the grafts. Therefore the selection of this location requires careful preoperative planning to obtain adequate block size. Extraoral harvesting requires hospitalization of the donor and general anesthesia, two surgery sites, and more cost. On the contrary intraoral harvesting can be performed with local anesthesia, intravenous sedation, or premedication. The most valuable advantages of intraoral harvesting are that local donor sites have convenient access and the bone is very short time ischemic. Also there is no morbidity due to second operation field since the same person is donor and recipient. The most utilized location for extraoral harvesting is the ilium, ribs, calvarium, tibia, and sometimes distal part of radius. Extraoral harvesting of the iliac crest is most often used for major jaw reconstruction. Both marrow and cortical iliac bone autografts are known as the most sure for bone growth of all. Bone can be taken from anterior and posterior iliac crest. Considerable large amounts of bone supply can be provided especially from the posterior iliac crest. Complications associated with the use of fresh iliac bone and marrow included root resorption and ankylosis, in regard to bone grafting around teeth. Later, these complications were minimized by either freezing the bone graft in a storage medium or adding autologous intraoral bone to the harvested iliac crest bone graft mixture. Intraoral autogenous bone grafts have been harvested from various intraoral sites including mandibular symphysis, mandibular ramus and retromolar area, maxillary tuberosity, coronoid process, mandibular and palatinal tori, and zygomatic bone. Augmentation of local alveolar defect when implanting can be done using bone harvested from maxillary tuberosity. Main sources of cortical bone suitable for onlay grafting are the mandibular and retromolar ramus. Mandibular symphyseal bone can be used for secondary alveolar cleft bone grafting maxillary sinus and rafting alveolar defects before implants placement.

Autogenus Bone Graft Harvested Form: Mandibular Symphysis From all intraoral donor sites mandibular symphysis produces biggest amount of graft bone. Maximum average bone block size is approximately 21  10  7 mm or 4.7 ml bone. But in clinical study conducted by Misch the approximate volume for the symphysis graft was 1.74 ml. This donor location is indicated when the surgeon needs cortical bone, but if he needs more cancellous bone an alternative donor site is preferred. In some cases bicortical bone blocks can be harvested, but there are some complications afterward. Usually monocortical block bone is harvested and cancellous bone taken with use of curettes for scooping [16]. Access to the mandibular symphysis area can be achieved by one of three different incision designs: (1) sulcular, (2) attached gingiva, or (3) vestibular [17]. Block bone grafts harvested from the symphysis can be used for predictable bone augmentation up to 6 mm in horizontal and vertical dimensions. The range of this cortical cancellous graft thickness is 3–11 mm, with most sites providing 5–8 mm. The density of the grafts is D-1 or D-2, and up to a three-tooth edentulous

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site can be grafted. In contrast, the ramus buccal shelf provides only cortical bone with a range of 2–4.5 mm (with most sites providing 3–4 mm). This site is used for horizontal or vertical augmentation of 3–4 mm. Symphysis block graft – indications: • Horizontal augmentation 4–7 mm (up to three-tooth defect) • Vertical augmentation 4–6 mm (up to three-tooth defect)

Autogenic Bone Graft Harvested Form: Mandibular Retromolar Area and Ramus This donor location for autogenic bone graft is relatively easy for producing a cortical bone harvest. The usually used area is the buccal site at the second and third molar area distal to the molar. The average amount of bone that can be taken from site is in volume 0.9 ml and in dimensions 35  10  4. Harvesting of bone from this area requires knowledge of the mandibular canal anatomy to prevent nerve injury. There is less postoperative complication compared to symphysis site, and depending on the amount of harvested bone these complications can become more severe. The deficiency of the block taken from the mandibular ramus harvest site will regrow and may be reharvested at a later date. Autogenic Bone Graft Harvested Form: Maxillary Tuberosity This donor site is usually used for producing cancellous bone grafts, and the volume of the graft is very limited. This procedure of harvesting can be done when additional bone is needed for extending the bone volume with other intraoral harvest. Other Intraoral Bone Donor Sites Other locations which can be used for bone harvesting are mandibular and palatal tori, zygomatic bone, and bone from coronoid proccesus. If there is a palatal or mandibular torus alveolar augmentation on edentulous ridges can be mainly done with cortical bone harvested with burs or suction traps. Zygomatic bone as donor site is a technique with very mild complications and can be done with local anesthesia. It can be used as donor site for alveolar bone reconstructions, and the amount harvested is 0.5–1.5 ml in volume. A new method for the reconstruction of small anterior maxillary alveolar bone defects was described using donor bone from the zygomatic buttress region. One ramus buccal shelf can provide adequate bone volume for up to a three- and even four-tooth segment. Bone density is D-1 with minimal, if any, marrow available. Some sites require extensive bone graft volume, which necessitates simultaneous bilateral ramus buccal shelf and symphysis graft harvest. For graft volume of more than 6–7 mm thickness, a secondary block graft can be used after appropriate healing of the initial graft. Ramus buccal shelf block graft – indications: • Horizontal augmentation 3–4 mm (up to four-tooth defect) • Vertical augmentation 3–4 mm (up to four-tooth defect)

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Despite the many advantages block grafts offer for alveolar ridge augmentation, complications can occur when mandibular block autografts are used for horizontal and vertical augmentation. Morbidity with this grafting protocol is associated with donor and recipient sites. Symphysis donor site morbidity includes intraoperative complications, such as bleeding; mental nerve injury; soft tissue injury of cheeks, lips, and tongue; block graft fracture; infection; and potential bicortical harvest. Pain, swelling, and bruising occur as normal postoperative sequel and are not excessive in nature. Complications associated with the recipient site include trismus, bleeding, pain, swelling, infection, neurosensory deficits, bone resorption, dehiscence, and graft failure. Trismus is expected if the recipient site is the posterior mandible, which affects the muscles of mastication.

Bone Augmentation Techniques and Material (Horizontal, Vertical) Dentistry has entered an era in which patients no longer need to accept an edentulous or partially edentulous condition or any other situation when their chance for dental implants can be dismissed because of insufficient alveolar bone volume, height, or width. A significant dimensional change in the alveolar bone appears as a result of tooth extraction, fractures, or pathological processes. These events eventually conduct to reduction in ridge height and width, with notable changes in both bone crests, whereupon buccal crest resorbs more quickly than the lingual one. This loss of alveolar bone volume has negative effect of the final result to restore the lost dentition. The consequences of lost bone height and ridge width can make it difficult to get an ideal placement of an implant and can result in compromised aesthetics of the prosthetic restoration. Therefore, augmentation of an insufficient bone volume is often indicated prior to or in combination with implant placement to achieve predictable long-term functioning and an aesthetic treatment outcome [2].

Clinical and Radiological Assessments Jaw dental segment (JDS) is defined as a vertically cut jaw segment with tooth, alveolar bone, and all or part of the basal bone. The location of bone suitable for implantation is identical with the former location of a tooth in the jaw. The number of the JDS describing the position of a planned implant in the jaw can be shown. If the JDS is edentulous, the term edentulous jaw dental segment (eJDS) is used (Fig. 2). To obtain clinical and radiological assessments of the eJDS, evaluation was begun at the widest point of each segment. Because the crest of the alveolar process was often thin, it was necessary to save it and thus produce a plane surface for the planned implant installation. In such cases, the heights of eJDS would have been

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Fig. 2 Edentulous jaw segments: (a) upper jaw, (b) lower jaw (A1, A2, A3, A4: width; B: height; C: length)

Fig. 3 Alveolar ridge vertical position (RVP). The vertical component of the alveolar bone defect is measured from the lowest point of the defect to an imaginary tangent running through the necks of the adjacent teeth (A – in aesthetic zone; B – in no aesthetic zone). A distance of 3 mm constitutes a significant cosmetic defect

shortened by 1–3 mm; this change had to be considered when performing dental segment height evaluation. The height of the alveolar process (H) is the distance between the crest of the alveolar process and the important vital structures of the jaws (nasal sinus floor, mental foramen, anterior loop of mental nerve). Alveolar ridge vertical position (RVP) is the distance from the lowest points of alveolar ridge crest to the cervicoenamel line of the adjacent teeth. This parameter is important for achieving of implant-supported restoration length equability to contralateral tooth (Fig. 3).

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The main goals of radiological jawbone examination are to determine the quantity, quality, and angulations of bone; select of the potential implant sites; and verify absence of pathology. Clinicians should choose proper radiographic method which provides sufficient diagnostic information with the least possible radiation dose.

Bone Augmentation Procedures Bone augmentation procedures can be done some time before implant placement (two-stage procedure) or at the same time as implant placement (one-stage procedure), using various materials and techniques [18]. A two-stage augmentation procedure is usually indicated to place an implant with the desired dimensions in an alveolar ridge with insufficient height or width. The first outcome parameter of interest in a two-stage bone augmentation procedure is the possibility of implant placement in an ideal position for the later prosthetic restoration. The long-term goal, for both one-stage and two-stage augmentation procedures, is the stability of the augmented bone volume, allowing unhindered masticatory function and optimal aesthetics, as expressed by implant survival, bone stability, and soft tissue stability [2]. There are different indications, numerous alternative techniques, and various biologically active agents and biomaterials currently used to augment bone. Therefore, which bone augmentation technique will be used to reconstruct different ridge defects depends on the horizontal and vertical extent of the defect. The predictability of the corrective reconstructive procedures is influenced by the span of the edentulous ridge and the amount of attachment on the neighboring teeth; typically, reconstructive procedures are less favorable in defects that exhibit horizontal and vertical components [3]. Thus, the choice of augmentation technique will depend on the size of defect, horizontal or vertical, anatomical structures, and size of area to be augmented. Some surgical techniques used to augment bone volume include 1. 2. 3. 4. 5. 6.

Onlay bone grafting Interpositional bone graft Ridge split technique Guided bone regeneration Distraction osteogenesis Sinus augmentation

Horizontal bone augmentation procedures: any technique aimed at making the recipient bone wider or thicker in order to receive dental implants of adequate diameter (usually of a 3.5 mm diameter or wider). For horizontal defects mostly used is onlay bone graft, ridge split, and guided bone regeneration. Vertical bone augmentation procedures: any technique aimed at making the recipient bone higher in a vertical dimension in order to receive dental implants of

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adequate length (usually 9 mm or longer). In many instances a combination of horizontal and vertical bone augmentation is needed, and these procedures were included in the vertical augmentation group. Techniques used for vertical defects are onlay bone graft, distraction osteogenesis, interpositional bone grafting, sinus lift, and guided bone regeneration.

Onlay Bone Grafting The graft material is laid over the defective area to increase insufficient alveolar bone volume, height or width, or both of them. The host bed is usually perforated with a small bur to encourage the formation of a blood clot between the graft and recipient bed. The graft is immobilized with screws and plates or with dental implants. Onlay bone grafting can be made by block onlay grafts or particulate bone grafts. Block Grafting Technique When using autogenous block graft approaches for bone augmentation, a considerable amount of horizontal augmentation can be added predictably to the defect area [4]. Intramembranous bones (calvarias, mandible) have reduced rates of resorption compared to endochondral bone (iliac crest) due to a more dense cortical structure and architecture, have more structural integrity than particulate bone, and show less absorption. The graft should be 3–4 mm larger than recipient site to allow contouring, adaptation, and resorption of graft. The thickness of the bone graft should be slightly larger than the planned width or thickness. The stabilization and intimate contact of these block grafts to the recipient bed has been considered crucial to a successful outcome [19]. This can be achieved with the use of bone fixation screws [4] or the simultaneous placement of dental implants. Aggressive recipient bed preparation with decortication, intramarrow penetration, and inlay shaping also has been supported, because of increases in the rate of revascularization, the availability of osteoprogenitor cells, and the increased rate of remodeling [4, 19]. Vertical onlay graft has two major concerns: • Increased risk of graft exposure due to overexpansion of the soft tissue. Covering of membrane with PRP gel reduces chance of graft exposure. • Adequate adaptation of the bone graft. In cases where cortical bone grafts are too thin, a stack technique can be used. All spaces should be filled with particulate bone and covered with a membrane. All grafts must be rigidly fixated with a screw to prevent movement of the graft. Movement will disrupt blood clot adhesion and will compromise vascularity. Particulate Bone Grafting Technique This augmentation technique has advantages of rapid vascularization, higher rate of resorption, and must be covered with membrane. Today allogenic and xenogenic particulate grafts are growing in popularity. Particulate onlay grafts must be used with a membrane. When a rigid membrane is used, the choice of material is less

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important as the membrane will maintain space and protect the graft. The particulate graft must have the strength to resist deformation when a nonrigid membrane is used. As any other technique, onlay bone grafting has advantages and disadvantages. This technique is less invasive compared to the other augmentation techniques and can be used for both horizontal and vertical augmentation. There is no injury to inferior alveolar nerve. A disadvantage is their high incidence of bone resorption (41.5 % during first 6 months (Cordaro et al.); 17.4 % during first 6 months (Proussaefs et al.); 20 % during first 5 months (Pikos)). Bone resorption is higher in vertical grafts. It can be reduced with use of membranes or metallic mesh and must be kept in position for entire healing period. Block graft is a better option although both can be used.

Interpositional Bone Graft This is an augmentation technique where a section of jawbone is surgically separated and graft material sandwiched between two sections. It is based on theory that bone placed between two pieces of pellicled bone with internal cancellous bone will undergo rapid and complete healing and graft incorporation. Grafts heal with rapid vascularization and bone remodeling in the bone gap. There is minimal bone resorption, and these grafts are almost indistinguishable from the surrounding tissue at 12 weeks post op. Interpositional bone grafting technique is useful in reconstruction of atrophic alveolar ridges. Apart from the ridge split technique, interpositional bone grafts are mainly used for vertical defects. It can however also be used in conjunction with horizontal onlay grafts. Its use is however restricted and abandoned due to the risk of nerve damage and lack of bone retention after grafting. Ridge Split Technique Ridge splitting is an alternative to the various techniques described for horizontal ridge augmentation where the alveolar ridge is split longitudinally and parted to widen it and allow placement of an implant or graft material or both in the void. The longitudinal split can be limited by placing transverse cuts in the bone. This technique is used for horizontal defects to widen the alveolar ridge. There must be at least 2 mm of crestal bone width and minimal reflection of surrounding soft tissue in order to maintain blood supply. Bone is forced buccal with interpositional bone grafts to keep segments separated. It is important to keep ridge relationship between mandible and maxilla in mind. The segments must be stabile to prevent disruption of vascularization of new bone. Guided Bone Regeneration (Bone Augmentation with Barrier Membrane Technique) The concept of GBR was described first in 1959 when cell-occlusive membranes were employed for spinal fusions. The terms “guided bone regeneration” and “guided tissue regeneration” (GTR) often are used synonymously and rather inappropriately. GTR deals with the regeneration of the supporting periodontal apparatus, including cementum, periodontal ligament, and alveolar bone, whereas GBR

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refers to the promotion of bone formation alone. GBR and GTR are based on the same principles that use barrier membranes for space maintenance over a defect, promoting the ingrowth of osteogenic cells and preventing migration of undesired cells from the overlying soft tissues into the wound. Protection of a blood clot in the defect and exclusion of gingival connective tissue and provision of a secluded space into which osteogenic cell from the bone can migrate are essential for a successful outcome. The sequence of bone healing is not only affected by invasion of nonosteogenic tissue but more so by the defect size and morphology [4]. Grafting materials can be categorized in one of the following groups [2]: • • • • • •

No graft (coagulum) Autograft block (extraoral or intraoral donor site) Autograft particulate Autograft from bone trap Membrane alone (nonresorbable or resorbable) Allograft (freeze-dried bone allograft [FDBA] or demineralized freeze-dried bone allograft [DFDBA]) • Xenograft (demineralized bovine bone mineral [DBBM], algae-derived, or coralderived) • Alloplast (hydroxyapatite [HA],  tricalciumphosphate [TCP], bioglass, or calcium sulfate or allograft + alloplast).

Autogenous Bone Grafts Autogenous bone corresponds to bone graft obtained from the same individual and used to build up the deficient area. It is considered to be the material of choice (Palmer 2000), i.e., it is the gold standard [18]. Autogenous grafts are biologically compatible as they are from the same patient and provide a scaffold into which new bone may grow. They are not immunogenic and contain osteoblasts and osteoprogenitor stem cells, which are capable of proliferating. These grafts, therefore, are osteoinductive. Autogenous grafts are presented as blocks or particles and can be used isolated or associated with allogenic or alloplastic grafts. Autogenous grafting may include cortical, cancellous, or cortical-cancellous bone. The donator area can be from intraoral sites (mentonian region, retromolar area, maxillary tuberosity) and extraoral sites (iliac crest, rib, cranium, tibia, and fibula) [20]. Because the survival of osteocytes depends on the presence of a vascular supply within a distance of 0.1 mm, cortical bone grafts lacking vascular and cellular pools on endosteal and periosteal surfaces may not be able to sustain cellular viability. Cancellous bone grafts may have a greater likelihood of supporting cell survival because of the possibility of diffusion of nutrients and revascularization from the recipient bed [21]. Nevertheless, this type of graft may cause morbidity in the donator area, hematoma, edema, infection, chronic pain, and vascular and nerve lesions. In addition, this technique spends more time for surgical procedure and is limited for large

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reconstructions. So, biomaterials have been suggested as an alternative to solve those limitations and reduce the gap between bone and implant [20, 22]. Allografts Allografts are bone grafts that are harvested from cadavers. Allograft material is freeze dried, demineralized, irradiated, and treated with ethylene oxide to prevent the transmission of disease. They contain no viable cells. Allografts are resorbable and supplied by specially licensed tissue banks in several convenient ways such as bone particles or large blocks [9]. Examples of allografts are fresh-frozen bone, freezedried bone, and demineralized freeze-dried bone. Freeze-dried bone and demineralized freeze-dried bone allografts are reported to be less immunogenic than fresh-frozen bone allografts. These materials are frequently used in mixtures with osteoinductive autogenous bone or bone substitutes [21]. The advantages of using allografts are that the material is available in large quantities and there is no donor site within the patient. The disadvantages are that the process for preparing the graft (i.e., freeze-drying and irradiating) decreases the material’s integrity and osteogenic potential, and the immunological response to it may diminish its incorporation into the recipient bone. Xenografts Xenografts are made of naturally derived deproteinized cancellous bone from animals such as cow or bone-like minerals (calcium carbonate) derived from corals or algae [18]. Deproteinized bovine bone is the most researched grafting material and is widely used in dentistry because of its similarity to human bone. Proteins in deproteinized bovine bone have been extracted to avoid immunological rejection after implantation; however, as the deproteinizing procedure eliminates the osteoinductive capacity, deproteinized bovine bones act solely as an osteoconductive scaffold [21]. The risk of transmission of diseases such as bovine spongiform encephalopathy is negligible because the bone’s organic component is extracted. The advantages and disadvantages of using xenografts are similar to those of using allografts. Alloplasts Alloplasts are synthetic bone substitutes. They are made of biocompatible, inorganic materials including synthetic hydroxyapatite, tricalcium phosphate, calcium carbonate, and bioactive glass. Whether synthetic hydroxyapatite is resorbable or nonresorbable depends on the temperature at which it is prepared. High-temperature preparation of hydroxyapatite results in a nonresorbable, nonporous, dense material, which is used as a filler. Tricalcium phosphate acts as a filler and is partially resorbable. Calcium carbonate, which is derived from coral, is biocompatible and resorbable so that it acts as a filler, which eventually may be replaced by new bone. Bioactive glass is a silicone-based, osteoconductive material that bonds to bone through the formation of carbonated hydroxyapatite.

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Some surgeons use these materials in combination with autogenous bone grafts or allografts [18]. The advantage of alloplasts is that they have no potential for disease transmission. Barrier Membranes for Guided Bone Regeneration (GBR) This technique uses special barrier membranes to protect defects from the ingrowth of soft tissue cells so that bone progenitor cells may develop bone uninhibited. Ingrowth of soft tissue may disturb or totally prevent osteogenesis in a defect or wound. Examples of membranes are expanded polytetrafluoroethylene, porcine collagen, and polyglactin. Membranes can be resorbable or nonresorbable [18]. Bone Morphogenetic Proteins (BMPs) and Platelet-Rich Plasma (PRP) BMPs are a family of proteins naturally present in bone and responsible for activation of bone development (Valentin-Opran 2002). BMPs may encourage bone formation. They may be incorporated into any of the above graft types. Growth factors and PRP are used to promote bone formation [18]. Many techniques exist for effective bone augmentation. The approach largely is dependent on the extent of the defect and specific procedures to be performed for the implant reconstruction. It is most appropriate to use an evidence-based approach when a treatment plan is being developed for bone augmentation cases [4]. Furthermore, every type of augmentation material can be used combined with a variety of different surgical techniques. There is possibility of many treatment combinations, and the situation is rather complicated. In addition, new techniques and “active agents” are continuously introduced in clinical practice [18]. Future bone augmentation approaches likely will use molecular, cellular, and genetic tissue engineering technologies.

Bone Substitutes The need for bone substitution materials has been increased due to tooth loss, trauma, and tumor and bone reconstructive surgery. A variety of grafting materials are used for bone augmentation in modern dentistry. All osteoplastic materials can be divided into four groups by origin: autogenic (the donor is the patient), allogenic (the donor is another person), xenogenic (the donor is an animal), and synthetic (on the basis of calcium salts) [14]. Although the autogenous bone is still considered as first option for bone reconstruction in implantology some limitations, such as donor site morbidity availability and unpredictable graft resorption, have stimulated the search for suitable synthetic grafting material [23]. However, the perfect grafting material has yet to be identified. There are four characteristics that an ideal bone graft material should demonstrate which include osseointegration, osteoconduction, osteoinduction, and osteogenesis. Only autogenous bone graft satisfies all of these requirements. Synthetic bone grafts at most possess only two of these four characteristics (osteointegration, osteoconduction).

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The ideal bone graft substitute is biocompatible, bioresorbable, osteoconductive, osteoinductive, structurally similar to bone, has no risk to disease transmission, appropriate lifetime, easy manipulation, able to form a suitable shape easily, and accessible cost [20]. From a mechanical point of view synthetic bone graft substitutes should have a similar strength to that of the cortical/cancellous bone being replaced. This needs to be matched with a similar modulus of elasticity to that of bone in an attempt to prevent stress shielding as well as maintaining adequate toughness to prevent fatigue fracture under cyclic loading. Synthetic materials that demonstrate some of these properties are composed of calcium, silicon, or aluminum. Alloplastic bone substitutes represent a large group of chemically diverse synthetic calcium-based biomaterials. They vary in chemical composition, structure, and mechanical and biological properties. Some of them are nonresorbable, while others are chemically resorbable with a concomitant release of bioactive ions. Their pore size is a very important determinant of the ability to form bone. Alloplastic graft materials with pore sizes 300 μm show enhanced formation of new capillaries and bone. Most of the current commercial materials do not exhibit any pores [21]. Synthetic resorbable materials were intended as an inexpensive substitute for natural bone. Synthetic graft materials include various types of ceramics: – Tricalcium phosphate (CP) – Bioglass – Hydroxyapatite (HAP) and its compositions with collagen, sulfated glycosaminoglycans such as keratin and chrondroitin sulfate, as well as sulfate and – Calcium phosphate. Now, many various forms of porous nanostructured calcium phosphate ceramics, bone cements, biohybrids, and biocomposite compounds have been created [24, 25].

Calcium Phosphates The calcium phosphate family of alloplastic graft materials has both osteoconductive and osseointegrative characteristics. Osseointegration results from the formation of a layer of hydroxyapatite (HA) shortly after implantation. The Ca2+ and PO4 2 ions required to establish this layer are derived from the implant and surrounding bone. The pathways of both Ca2+ and PO4 2 ions have been traced in serum and urine without any significant elevation in serum levels from which it can be concluded they are handled as part of the normal body ion pool. They have an excellent record of biocompatibility with no reports of systemic toxicity or foreign body reactions. Calcium phosphate materials such as tricalcium phosphate (TCP) and hydroxyapatite (HA) for the first time were established for clinical use in the 1980s. They can be found in a variety of forms (e.g., pastes, putties, solid matrices, granules). Based

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upon their chemical composition, calcium phosphates can be separated into hydroxyapatite (HA), tricalcium phosphate (TCP), and composite grafts [26].

Beta Tricalcium Phosphate (bTCP) Beta tricalcium phosphate (β-TCP) was one of the earliest calcium phosphate materials used for a bone augmentation. In 1920 Albee and Morrison reported that the rate of bone union was increased when β-TCP was injected into the gap of a segmental bone defect. The β -tricalcium phosphate is one of the most frequently used alloplasts in implant dentistry and has been regarded as a material of choice because it is osteoconductive, absorbable, and nonosteoindictive [20]. β-TCP may be a suitable bone substitute that will biodegrade and be replaced by newly mineralizing bone tissue without fibrous tissue proliferation [23]. These biomaterials facilitate attaching, proliferation, migration, and phenotypic expression of the bone cells, which leads to apposition growth of the bone on the graft surface [14]. When particles of β-TCP are mixed with the blood clot and surrounded by the bony walls of the alveolar socket, osteogenic cells, including undifferentiated mesenchymal stem cells, start migrating from the existing bone surface between and over the surface of the particles, stimulated mostly by an adhesive glycoprotein, called fibronectin, a component of the forming blood clot [23]. Because of their capability of adsorbing proteins, function of osteoclasts and osteoblasts is stimulated, wherefore the function of competing cells is inhibited [23]. Despite the aforementioned positive biological properties, the drawback of most CP materials is poor mechanical durability and slow resorption in the body tissues, which limits its use in bone augmentation procedures performed for aesthetic purposes [21]. It has been found to be brittle and weak under tension and shear but resistant to compressive loads. Beta tricalcium phosphate undergoes reabsorption via dissolution and fragmentation over a 6–18- month period. Unfortunately the replacement of βTCP by bone does not occur in an equitable way. That is, there is always less bone volume produced than the volume of βTCP reabsorbed. Beta tricalcium phosphate is available in porous or solid form as either granules or blocks. Structurally porous βTCP has a compressive strength and tensile strength similar to cancellous bone.

Nanosized HAP HAP also could be found in the form of nanosized crystals. This material has two elements which are most important for the physiology of bone tissue: they are in a dynamic equilibrium with their biological environment in the remodeling cycle (resorption/mineralization) and manifest a high level of mechanical properties. Nanocrystalline HAP possesses an enhanced capability to adsorb proteins required

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for the vital activity of the cells, as well as discrimination regarding the function of the cells which form osseous and fibrous tissues. Earlier preclinical studies have shown that nanostructural HAP obtained at temperatures below 60  С possesses a significantly larger capability to stimulate reparative osteogenesis compared with its polycrystalline (high-temperature) analogue. Nanocrystals of biological HAP make the bone harder and stiffer, whereas collagen fibers ensure elasticity and high cracking resistance as well as an adequate resorption and bone regeneration rate [14, 27].

Synthetic Hydroxyapatite (HA) Hydroxyapatite C10 (PO4)6(OH)2 forms the principal mineral component of bone. Synthetic HA comes in ceramic or nonceramic form as porous or solid, blocks or granules. Ceramic refers to the fact that the HA crystals have been heated (sintered) at between 700  C and 1300  C to form a highly crystalline structure. Ceramic HA preparations are resistant to reabsorption in vivo, while nonceramic HA is more readily reabsorbed. Synthetic HA have good compressive strengths but are weak in tension and shear. Synthetic HA in solid block form is difficult to shape, does not permit fibro-osseous ingrowth, and has a much higher modulus of elasticity than bone. Synthetic HA has been successfully used to coat metal implants to enhance their osseointegration [28]. Synthetic HAP is used in the form of nonporous (nonresorbable) and porous (resorbable) ceramics. With nonporous ceramics no osteogenesis is taking place directly in the area occupied by the bone and porous HAP ceramics is an osteoconductor. One of the forms of porous ceramics used is its granulate form after implantation of high-temperature ceramic granulates into bone defects, with extension growth of the connective tissue and the osteogenic cells present within it. This served as the base for using this material as a surface coating for endoprostheses, osteosynthesis constructions, and dental implants. The process is most intensive primarily near the surface of HAP particle conglomerates close to the source of osteogenic cells (bone defect walls) [14]. Porous granular form has been used alone or with bone graft to fill voids.

Coralline Hydroxyapatite Another form of alloplast graft material developed in 1971 with a purpose to create a HA implant with a consistent pore size and improved interconnectivity is Coralline HA. Coralline HA utilizes the genetically determined highly regular and permeable structure of marine coral, which closely resembles that of cancellous bone. The replamineform process involves processing of the calcium carbonate coral to remove the bulk of its organic matter. It is then subject to both extreme pressure and heat in an aqueous phosphate solution. This converts the calcium carbonate coral skeleton

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entirely to calcium phosphate (HA) as well as sterilizing it at the same time. Mechanically coralline HA is only slightly greater in compressive strength than cancellous bone. Like the other HA preparations it is weak in tension, brittle, and difficult to shape. Its main advantage is that its interporous structure allows complete ingrowth of fibro-osseous tissue. Fifty to eighty per cent of the void is filled within 3 months. Coralline HA initially does not possess the strength of trabecular bone nor the plastic properties because it lacks a collagen matrix; but with completion of fibroosseous ingrowth the coralline HA becomes stronger but is less stiff than cancellous bone. Coralline HA has been successfully used in nonweightbearing applications such as maxillofacial, periodontal augmentation, and also for distal radial fractures, spinal fusions, and orbital restorations.

Calcium Sulfate Calcium sulfate (plaster of Paris) has been used in craniofacial surgery for more than 100 years, although its use is documented for fracture treatment by the Arabs even in the tenth century. In 1892 a German by the name of Dreesman successfully used plaster of Paris medicated with a 5 % phenol solution to fill skeletal defects, and De Leonardis and Pecora have used the material for sinus floor augmentation in implant dentistry [21]. Calcium sulfate is thought to act as an osteoconductive matrix for the ingrowth of blood vessels and associated fibrogenic and osteogenic cells. For this to occur it is critically important that the implanted calcium sulfate is adjacent to viable periosteum or endosteum. Over a period of 5–7 weeks the calcium sulfate is reabsorbed by a process of dissolution. Calcium sulfate resorbs quickly and is substituted by new bone. The rapid resorption rate can pose a potential problem because the volume of the graft may not be maintained for a sufficiently long period of time to yield reliable grafting results in the aesthetic zone [21]. Its compressive strength is greater than the cancellous bone and has slightly less tensile strength. However, calcium sulfate needs a dry environment to set, and if it is re-exposed to moisture it tends to become soft and fragment. Because of this, it has no reliable mechanical properties in vivo, and its application should be limited to a contained area.

Bioactive Glasses Two families of silicon-based compounds have the ability to bond directly to bone. These are the bioactive glasses and the glass ionomers. Bioactive glasses are hard, solid (nonporous), materials consisting of four components: sodium oxide, calcium oxide, silicon dioxide (silicate, which is the main component), and phosphorous pentoxide [25]. By varying the proportions of sodium oxide, calcium oxide, and silicon dioxide, forms can be produced that are soluble in vivo (solubility being proportional to the sodium oxide content) right through to those that are essentially nonresorbable.

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Bone substitute materials within the group of bioactive glass display osteoinductive and osteoconductive properties. They are bioactive, as they interact with the body. Bioactivity depends upon the SiO2 content; the bonding between bone and glass is best if the bioactive glass contains 45–52 % SiO2 [26]. A mechanically strong bond between bioactive glass and bone forms as a result of a silica-rich gel layer that forms on the surface of the bioactive glass when exposed to physiological aqueous solutions. Within this gel Ca2+ and PO4 2 ions combine to form crystals of hydroxyapatite (HA) similar to that of bone, hence a strong chemical bond. The extracellular proteins attract macrophages, mesenchymal stem cells, and osteoprogenitor cells. Subsequently, the osteoprogenitor cells proliferate into matrix-producing osteoblasts [26]. Bioglass is slowly resorbed, so it takes 12–16 months before the graft is replaced by newly formed bone, and that is a factor that should be considered when planning the graft healing time. The studies by Tadjoedin et al. and Turunen et al. suggest that bioglass can be used in a mixture with autogenous bone at the floor of the maxillary sinus, thus decreasing the amount of autogenous bone required [21].

Glass Ionomers Glass ionomer cements were first introduced in 1971 for dental use where cement was required to bind tooth enamel in a moist environment. Ionomeric cement consists of calcium/aluminum/ fluorosilicate glass powder (0.001–0.1 mm diameter) which is mixed with polycarboxylic acid. This results in an exothermic reaction (56  C) with CO2 evolution to produce a porous cement paste. The paste sets hard in approximately 5 min after which it is water insoluble. Prior to this it must be protected from wound fluids which will dissolve it. After 24 h it has a compressive strength (180–220 MPa) and modulus of elasticity comparable to cortical bone. It is biocompatible and osseointegrates in a manner similar to bioactive glasses. Its porous structure aids osteoconduction and subsequent bone ingrowth. It is non-reabsorbable and therefore is not replaced by bone. Except in dentistry, it is also used in nose, ear, and throat surgery and in maxillofacial reconstructive surgery, but its use in contact with neural tissue or cerebrospinal fluid (CSF) is contraindicated because the release of aluminum ions and polyacid in its unset form is neurotoxic.

Combined Synthetic Materials Combined synthetic materials for bone augmentation are a polymeric matrix (polylactide, polyglycolic acid, polyoxybutyrate, and their combinations) and nano-HAP as a filler. Development of composites made of synthetic HAP in different forms (powders, granules, gels) in combination with other preparations (protein collagen, alginate, polysaccharides chitosan, hyaluronic acid, peptides, embryonic stem cells) expanded the opportunities of bone reconstruction [14].

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Mechanism of Bone Regeneration After Ridge Augmentation Following tooth removal, the normal healing process takes place over approximately 40 days, starting with clot formation and culminating in a socket filled with bone covered by connective tissue and epithelium. Complete preservation and restoration of the original ridge volume after tissue remodeling would be ideal for future implant placement. Unfortunately, this is usually not the case. In fact, without further treatment, crestal bone resorption is common and unavoidable which can lead to significant ridge dimensional changes. These changes range from an average vertical bone loss of 1.5–2 mm and an average horizontal ridge width loss of 40–50 % over 6–12 months healing. Most of the dimensional changes occur during the first 3 months and can continue over time, with as much as an additional 11 % of volumetric bone loss during the following 5 years [28–31]. Tooth extraction resulted in approximately 40–60 % loss of bone height and width respectively within 2–3 years. More often, greater bone resorption occurs in the horizontal plane than in the vertical plane, leading to more severe loss of alveolar width. The presence of bone dehiscences or fenestrations during extraction may increase postextraction alveolar remodeling, leading to an even more severe buccal concavity after healing.

The Terms “‘Guided Bone Regeneration” (GBR) and “Guided Tissue Regeneration” (GTR) One technique of ridge augmentation is guided bone regeneration (GBR). GBR is a surgical procedure that uses barrier membranes with or without particulate bone grafts or/and bone substitutes. Guided bone regeneration (GBR) and guided tissue regeneration (GTR) are surgical techniques that utilize barriers made by different membranes to provide navigation of the growth of new bone or gingival tissue. These techniques were first described by Melcher (1976) and are used in everyday practice for more than three decades. These two terms, guided bone regeneration and guided tissue regeneration, often are used synonymously, but that is inadequate. GTR primarily is linked with regeneration of the supporting periodontal apparatus, including periodontal ligament, cementum, as well as alveolar bone. GBR is oriented only to the promotion of bone tissue formation. GBR and GTR are based on the same biological and surgical principles. In both barrier membranes are used to make space maintenance over a defect, promoting the pluripotential of osteogenic cells and preventing migration of undesired cells from the overlying soft tissues into the wound. So, guided bone regeneration is similar to guided tissue regeneration but is focused on development of bone tissues in addition to the soft tissues of the periodontal attachment. GTR and GBR are used at sites that are having insufficient quantity or quality of bone or gingiva. Nowadays, guided bone regeneration is predominantly applied in dentistry to get support for new bone tissue growth on an alveolar ridge to make

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adequate placement of dental implants. In fact guided bone regeneration originate from guided tissue regeneration, but these techniques are directed to regenerating tissue in osseous defects adjacent to natural teeth. For adequate and successful results from these procedures, protection of a blood clot in the wound and exclusion of gingival connective tissue is necessary. Also provision of a secluded space into which osteogenic cell from the bone can migrate is essential for adequate outcome. The sequence of bone healing is not only affected by invasion of nonosteogenic tissue but more by the size of the defect and its morphology. The patterns of bone regeneration are very complex and include various pathological and pathophysiological changes. This process primarily involves angiogenesis (of new blood weasels) and ingress of osteogenic cells from the defect periphery toward the center. In this way is created new well-vascularized granulation tissue in the surgical defect. This provides a scaffold for woven bone proliferation and bone apposition within the defect. A very important factor for GTR and GBR is the size of the defect. The dimensions of the defect have influence on the bone healing capacity. In circumstances where the wound is too large to generate a biomechanically stable central scaffold, bone formation is limited to the marginal stable zone with a central zone of disorganized loose connective tissue. Thus, combined use of bone grafts or bone replacement substitutes with barrier membranes are the best for bone regeneration of larger defects. Biologically, no matter if it is GBR or GTR, these three factors have the main role: 1. Progenitor cells that can produce appropriate tissue and supporting structures also 2. Growth factors for coordinating these activities 3. Scaffolds to provide space and a structure framework for deposition of extracellular matrix Except the differences between GTR and GBR, these two different surgical techniques share some common prognostic factors. For adequate GTR or GBR there is a need of stabile blood clot, to achieved wound healing per primam intentionem, isolation of the primary defect from the gingival tissue, and space provision. Also, GBR and GTR today have benefits from contemporary advances in different fields of regenerative medicine including gene therapy, cell therapy, or usage of growth factors.

Guided Tissue Regeneration Guided tissue regeneration consists of positioning barriers to cover the bone and periodontal ligament, thus separating them from the gingival epithelium temporarily. With this surgical technique the gingival epithelium and connective tissue are excluded from the bone surface in the postsurgical phase. In this way is made

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prevention of epithelial migration into the surgical wound and favors repopulation of different mesenchymal cells from the periodontal ligament and alveolar bone [32]. Guided tissue regeneration is primary periodontal occupation and is theoretically based on the potential that only periodontal cells possessed for regeneration of the attachment apparatus of the tooth. This procedure prevents epithelial migration along the dental cement of the pocket with help of biomembrane. The final result from this surgical procedure should be reconstruction of the attachment and coverage of the cementum. GTR is induced when there is loss of periodontal tissue and infrabony defects. The aim of this technique is recontouring of the periodontal ligament which have adequate organized and oriented collagen fibers in newly formed cementum and alveolar bone. Today, contemporary GTR surgical techniques have minimally invasive approach with minimal patient morbidity and reduced need for membranes. In periodontology are used particulated grafts from autogenous, xenogenic, allogenic, and rarely synthetic origin. In the GTR different types if membranes have high importance. Applying membranes between the gingival epithelium and connective tissue and the periodontal ligament and alveolar bone prevents migration of the gingival epithelium and gingival connective tissue cells into the defect along the root surface. Progenitorial cells form the periodontal ligament and from the alveolar bone colonize the blood clot, and periodontal regeneration is induced [33]. Three different outcomes can be noticed after the GTR. If mesenchymal cells from the periodontal ligament or from the perivascular region proliferate into the defects, regeneration occurs. If bone cells migrate and make adherent on the root surface resorption or ankyloses occurs. If gingival epithelial cells proliferate along the root surface, the result will be long functional epithelial attachment. But, if gingival connective tissue migrates along the root surface, a connective tissue attachment can be the result. Also, this can result with root resorption.

Guided Bone Regeneration Guided bone regeneration is defined as a surgical procedure in which barrier membranes with or without particulate bone grafts or/and bone substitutes are used. Biologically, as is noticed, this type of regeneration by GBR depends on the migration of pluripotential and osteogenic cells (osteoblasts derived from the periosteum and/or adjacent bone and/or bone marrow) to the bone defect site and exclusion of cells impeding bone formation (epithelial cells and fibroblasts). This is the biggest difference from the guided tissue regeneration. Guided bone regeneration is indicated for 1. Filling bone defects after tooth extraction 2. Preparation of the bone for successful implantation 3. Building up of new bone tissue around dental implants in immediate implantation

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4. Filling bone defects after cystectomy 5. Filling bone defects after impacted tooth replacement The main goals of GBR are preservation of the postextraction alveolar ridge quantity and quality (normally it is spontaneously reduced), reconstruction of the alveolar bone after extraction for realizing implantation procedure, correction of peri-implant dehiscence and fenestrations, and reconstruction of the lost bone after peri-implantitis [34]. For GBR particulate or block grafts from autogenous, xenogenic, allogenic, and rarely synthetic origin can be used. These two scaffolds can be used to fill the defect and provide support and space for regeneration. Membranes used in GBR have a specific purpose – to prevent gingival epithelial cells or fibroblasts access into the wound, which results with fibrous connective tissue forming. The membrane must be carefully positioned, so a space must be created beneath the membrane, but it must isolate the defect from the overlying oral soft tissue. This space primarily is filled with fibrin clot, which is scaffold for settlement of progenitor cells. This cell originates from adjacent bone or bone marrow. Today, development of new improved osteoplastic bone grafting materials and growth and differentiation factors and tissue engineering made this surgical method very simple for realization. GBR today becomes a predictable surgical method to enhance new bone formation in peri-implant bone deficiencies and alveolar ridge augmentation, requiring excellent surgical skills [33]. After the surgical procedure is done, series of sequences in the wound take place. In the first 24 h the space which was made is filling with blood clot. It releases different substantives such as growth factors (mainly platelet-derived growth factor) and cytokines (IL-8). They attract two types of cells: neutrophils and macrophages. The blood clot is absorbed and replaced with young granulation tissue rich with new blood vessels that result from an enormous angiogenesis in this tissue. Through these vessels except nutrition, stem cells capable for osteogenesis can be transported. They are responsible for osteoid formation. This osteoid is a template for apposition of mineral and forming a lamellar bone. All of this at the end results with newly formed bone with compact and reticular bone tissue and mature bone marrow. This occurs 3–4 months after the surgery [35].

Membrane Membranes are used in periodontal and oral surgery to disable migration of cells, primarily epithelial, into the defects. The main role of the membranes is to prohibit epithelial cells, which have bigger regeneration capacity, into areas with tissue that has slower regeneration capacity, like bone tissue. These membranes are dominantly used for guided tissue regeneration (GTR) and guided bone regeneration (GBR) [36]. Membranes used in oral and periodontal surgery should fulfill some basic requirements as

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1. Biocompatibility (the interaction between membranes and host tissue should not induce negative effect on these tissues) 2. Space-making (the membrane should have ability to maintain a space for cells from surrounding bone tissue to migrate for stable time duration) 3. Cell occlusiveness (the membrane should prevent fibrous tissue that delays bone formation to invade into the defect site) 4. Mechanical strength (the membrane should have proper physical properties to allow and protect the healing process, including protection of the underlying blood clot) 5. Degradability (this characteristic is typical for resorptible membranes and means adequate degradation time matching the regeneration rate of bone tissue to avoid a secondary surgical procedure to remove the membrane), nonimmunogenicity (membranes should not provoke immune response of the host), nontoxicity (membranes should not act toxic on the organism) [33, 36] In history, the first membranes developed and fabricated were nonresorptible. The biggest disadvantage of these types of membranes is necessarily of second surgical intervention, after the healing process is finished. This disadvantage of the nonresorptible membranes is exceeded with development of the resorptive membranes. Thus the need of second surgical intervention is surpassed. Membranes used in dentistry can have natural or synthetic origin. Based on their capability of resorption, membranes are subdivided in three big groups: (a) Resorptive (b) Synthetic nonresorptive (c) Natural biodegradable Using of nonresorptible membranes has one more disadvantage – bigger risk of loss of some of the regenerated bone further to flap reflection. That’s why these types of membranes in contemporary dentistry are replaced with bioresorptive membranes.

Nonresorptible Membranes As it is mentioned nonresorptible membranes require second surgical intervention to be displaced, because they retain their primary form and structure in the tissue. In this way additional trauma and resorption of the alveolar bone may occur; additional patient discomfort also can be caused, and it has social meaning; this intervention costs more.

Polytetrafluoroethylene Nonresorbable membranes made from polytetrafluoroethylene (PTFE) can be divided into expanded polytetrafluoroethylene (e-PTFE) and high-density polytetrafluoroethylene (d-PTFE). This classification is made according to the structure.

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PTFE has unique chemical structure, it is a simple polymer because it is built up from two chemical elements: carbon and fluorine. This chemical structure has a long, straight carbon backbone to which the fluorine atoms are connected. Both the C–C and C–F connections are extremely strong. In addition, the electron cloud of the fluorine atoms forms a uniform helical sheath that protects the carbon backbone. The even distribution of fluorine atoms makes it nonpolar and nonreactive. The combination of strong bonds, a protective sheath, and no polarity makes PTFE extremely inert as well as thermally stable. Biologically PTFE is chemically identical, causes minimal inflammatory reaction in different tissues, and allows tissue ingrowth [36]. The e-PTFE has micropores, which allows nutrients to be supplied through multiple pores although it can allow the invasion of bacteria when the membrane is exposed to the oral cavity. The d-PTFE completely blocks the penetration of food and bacteria, and thus even if it is exposed to the oral cavity, the d-PTFE membranes exert good guided tissue regeneration (GTR) effects [37]. In contrast, other polymeric membrane materials have some or all of the fluorine atoms replaced with hydrogen or other elements. This results in weaker bonds and a more polar, reactive molecule. The substitution also increases the surface free energy. Therefore, these polymers are less hydrophobic, less thermally stable, and more reactive than PTFE.

Titanium Mesh Titanium mesh is another nonresorptible membrane which can be used as an option in GBR. Titanium mesh has mechanical strength and rigidity, so it can be used in large bone or tissue defects. Titanium mesh has low density and elasticity and rigidity corresponding with low weight. It also has high resistance of corrosion, although this is a metal. As metal titanium is highly reactive, but soon after its exposure is covered by layer of titanium oxide it becomes uncreative and corrosion stabile. Titanium mesh is good alternative for e-PTFE, because it makes good bone graft stabilization, provides extensive space for GBR and GBT, and prevents contour collapse. Its stability prevents graft displacement and its plasticity enabling adequate adaptation of any type of defect. Its elasticity prevents mucosal compression. Its smooth surface makes smaller chance for bacterial colonization. But titanium mesh has its own disadvantages as irritation of the mucosal flap and its sharp edges may cause cutting of the surrounding tissue [38]. The most important characteristic of this mesh is presence of macroporosity, which maintained good blood supply and diffusion of extracellular matrix through the membrane. These make difficulties in removing the membrane in the second surgical intervention and also with it excellent pathway for microorganisms can be done. That is the reason for increased presence of infections after using of titanium mesh.

Resorptible Membranes Resorptible membranes are widely used in oral and periodontal surgery. They are subdivided in two groups: polymers of lactic or glycolic acid. These types of

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membranes have a lot of advantages in comparison with the nonresorptible membranes like (a) There is no need of secondary surgical intervention for removal of the membrane. (b) Eliminate the need of follow-ups; reducing the possibility of secondary infection during the secondary intervention. (c) Patients’ discomfort is minimized or totally reduced. (d) Eliminate potential surgical complications and in this way. (e) The cost is reduced [39]. As indicated by their name they disintegrate in the tissue. Disintegration of the resorptible membranes starts soon after placement in the tissue. A specific characteristic of these membranes is different speed of disintegration. Persistence of these types of membrane is between 4 weeks and several months. Resorptible membranes release free acid that can cause inflammation and compromise the wound healing and regeneration outcome.

Collagen Membranes Collagen membranes are used for GTR, GBR, and root coverage. A variety of collagen materials are available including soluble collagen, collagen fibers, sponges, membranes, and bone implants allowing diverse usage of this material, for example, collagen fibers and sponges for hemostatic, collagen membranes for wound covering or implantation, and injections of soluble collagen in plastic surgery. They come from collagen type I and II from pigs or cows. With their biocompatibility and capacity of prompting wound healing they are widely used in oral and periodontal surgery, as well as in dental implantology, bone defect, and ridge augmentation nowadays [40]. Collagen membranes are resorptible membranes, and duration of their disintegration is between 4 and 40 weeks. They have a lot of advantages because they are used in dentistry to inhibit epithelial migration and promote attachment of new connective tissue. Also collagen membranes prevent blood clot by platelet aggregation in early stage of clot formation and wound healing [41]. The use of autogenous periosteum taken from the hard palate as a membrane is described in the literature. In this way there is a possibility of this periosteum to be used for GTR and GBR [42, 43].

Sinus Floor Augmentation, Bone Splitting, Distraction Osteogenesis Soft and hard tissue defects result from a variety of causes, such as infection, trauma, and tooth loss. These create an anatomically less favorable foundation for ideal implant placement [44].

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Survey of various literatures using Internet sources, manual searches, and common textbooks on dental implants shows that a thorough knowledge of conventional augmentation procedures such as bone augmentation techniques, guided bone regeneration, alveolar distraction, maxillary sinus elevation techniques with or without grafting, and contemporary techniques of implant placement provides effective long-term solutions in the management of the atrophic maxilla [45]. Autogenous bone grafting is the best and the gold-standard technique for bone augmentation procedures prior to implant placement. If the amount of available intraoral donor bone is insufficient, it is necessary to harvest bone graft from extraoral sites, such as calvarias. Although this technique is well established, only a few case reports show the histological analysis of the grafted bone at the moment of implant placement [46].

Maxillary Sinus Floor Elevation – Lateral Window Technique It is a surgical technique using a window into the lateral wall of the maxillary sinus to gain access to the maxillary sinus membrane. Following mobilization and elevation of the sinus membrane, bone augmentation materials (i.e., autografts, allografts, alloplasts, xenografts, or combination mixtures) are used to elevate the sinus floor and allow the placement of dental implants. If the original bone height permits sufficient primary implant stability, then a simultaneous procedure can be used. Sinus floor elevation is commonly used in cases where alveolar bone resorption has led to insufficient bone height for the placement of dental implants. Lateral wall sinus elevation is carried out when the bone is severely deficient. Although this procedure has a high rate of success, it may present surgical problems. A description of the anatomy of the maxillary sinus and lateral wall augmentation techniques leads to a discussion of the various challenges and complications that may arise and their management [47].

Transalveolar Sinus Floor Elevation A transalveolar approach for sinus floor elevation with subsequent placement of dental implants was first suggested by Tatum in 1986. In 1994, Summers described a different transalveolar approach using a set of tapered osteotomies with increasing diameters. The transalveolar approach of sinus floor elevation, also referred to as “osteotomy sinus floor elevation,” the “Summers technique,” or the “Crestal approach,” may be considered as being more conservative and less invasive than the conventional lateral window approach. The surgical technique uses a transalveolar approach to elevate the sinus floor by using osteotomy instruments. Anatomic aspects, such as an oblique sinus floor or insufficient bone height, can limit the use of this delicate surgical technique in daily practice. This is reflected by the fact that more than 9 out of 10 patients who experienced the surgical procedure are willing to undergo it again.

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The main indication for transalveolar sinus floor elevation is reduced residual bone height, which does not allow standard implant placement. Contraindications for transalveolar sinus floor elevation may be intraoral, local, or medical. The surgical approach utilized over the last two decades is the technique described by Summers, with or without minor modifications. The surgical care after implant placement using the osteotomy technique is similar to the surgical care after standard implant placement. The patients are usually advised to take antibiotic prophylaxis and to utilize antiseptic rinses. The main complications reported after performing a transalveolar sinus floor elevation were perforation of the Schneiderian membrane in 3.8 % of patients and postoperative infections in 0.8 % of patients [48]. The scientific literature demonstrated that maxillary sinus grafting is a reliable surgical technique which permits implants to be placed in the atrophic posterior maxilla with an excellent long-term prognosis. Similar results have been obtained with different grafting materials, such as autogenous bone, allografts, xenografts, alloplastic materials, and mixtures of these materials. Survival rates of implants placed in grafted sinuses are consistent with those of implants placed in no grafted edentulous maxillae.

Bone Splitting/Expansion and Immediate Implant Placement and Split-Ridge Techniques with Interpositional Bone Grafts and Delayed Implant Placement Bone splitting/expansion seems to be a reliable and relatively noninvasive technique to correct narrow edentulous ridges. Survival and success rates of implants placed in the expanded ridges are consistent with those of implants placed in native, nonreconstructed bone. The gap created by sagittal osteotomy/expansion undergoes spontaneous ossification, following a mechanism similar to that occurring in fractures. New bone formation permits a consolidation between the oral and buccal bone plates of the alveolus, and implants placed in expanded ridges seem to withstand the biomechanical demands of loading. However, some considerations have to be made. Bone splitting/expansion can be applied only when the buccal and palatal/lingual plates are separated by spongy bone. Therefore, the indications are more limited as compared to onlay grafts and GBR, which can be also applied in cases presenting with severe horizontal atrophy. Another limitation is represented by unfavorable inclination of implants placed in expanded areas. This procedure may lead to excessive buccal inclination of implants, which may create problems from a functional and aesthetic viewpoint. In the case of unfavorable bone angularity, GBR or bone grafting techniques seem to represent more adequate surgical procedures [48].

Distraction Osteogenesis Despite the limited number of patients and implants placed in the retrieved articles, the following conclusions can be drawn:

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• Distraction osteogenesis provides an opportunity to obtain a natural formation of bone between the distracted segment and basal bone in a relatively short time span, thus avoiding the necessity of autogenous bone harvesting. This leads to a reduction of morbidity and a shortening of operating times. Soft tissues can follow the elongation of the underlying bone (neohistogenesis), and there is a lower risk of infection of the surgical site (0 % in this case series). Both limited and extended (fully edentulous patients) defects can be treated. Some disadvantages of vertical distraction osteogenesis must be emphasized: • Frequent lingual/palatal inclination of the distracted segment has been reported by some authors, with an incidence varying from 13 % to 35.4 %, probably due to local muscle pull, inappropriate device positioning, and/or poor device trajectory. To solve this complication, different solutions have been suggested, including the use of fixed or removable prosthodontic and orthodontic devices to guide the distracted segment to its proper final position. Ideally, a multidirectional alveolar distraction device would allow the vector to be modified and guided in several planes of space [49].

Summary Recently there has been significant improvement in understanding bone repair. This has notable implications for the future management of bone loss. Currently autogenous and allograft bone are the main sources for bone grafting procedures. Concerns related to the use of both autograft and allograft has led to the search for alternatives. The synthetic bone graft substitutes as yet offer only a part solution to the treatment of localized bone loss. They have some of the desired mechanical qualities of bone as well as osseointegrative/conductive properties but are largely reliant on viable periosteum/ bone for their success. Ideally a synthetic bone graft substitute should mimic the native bone in both mechanical and osteogenic properties. The identification of osteoinductive proteins and other factors involved in the promotion of osteoblastic proliferation, differentiation, and function has enhanced the potential for manipulating local repair in a beneficial manner. The restoration of an adequate blood supply and the ability to maintain stability and controlled loading during the repair process are also important in achieving a satisfactory outcome. In order to obtain optimum results it is likely that manipulation of this process will require a combination of strategies depending on the clinical situation. A large but heterogeneous body of literature was available regarding augmentation of localized bone defects in the alveolar ridges after including all levels of clinical evidence except expert opinions. The major development in aesthetic dentistry, and more so the introduction of implant dentistry, led to significant developments aimed to regenerate or restore bony defects and bone loss in the edentulous ridge. Most clinical efforts in the developments in bone augmentation

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procedures are related to either simplifying clinical handling or influencing of biological processes. Many techniques exist for effective bone augmentation. The approach largely is dependent on the extent of the defect and specific procedures to be performed for the implant reconstruction. It is most appropriate to use an evidence-based approach when a treatment plan is being developed for bone augmentation cases.

References 1. Tal H, Artzi Z, Kolerman R, Beitlitum I, Goshen G (2012) Augmentation and preservation of the alveolar process and alveolar ridge of bone, bone regeneration. In: Prof. Haim Tal (ed) ISBN: 978-953-51- 0487–2. InTech, Available from: http://www.intechopen.com/books/ bone-regeneration/augmentation-and preservation-of-the-alveolar-process-and-alveolar-ridgeof-bone 2. Jensen SS, Terheyden H (2009) Bone augmentation procedures in localized defects in the alveolar ridge: clinical results with different bone grafts and bone-substitute materials. Int J Oral Maxillofac Implants 24(Suppl):218–236 3. Dimova C, Papakoca K, Papakoca V (2014) Alveolar bone augmentation. Key Eng Mater 614: 89–94. # Trans Tech Publications, Switzerland. doi:10.4028/www.scientific.net/KEM.614.89 4. McAllister BS, Kamran H (2007) AAP-commissioned review bone augmentation techniques. J Periodontal 78(3):337–396 5. Nguyen Ngoc Hung (2012) Basic knowledge of bone grafting, bone grafting. In: Dr Alessandro Zorzi (ed) ISBN: 978-953-51-0324-0. InTech, Available from: http://www.intechopen.com/ books/bone-grafting/basicknowledge-of-bone-grafting 6. Brånemark R, Brånemark P-I, Rydevik B, Myers RR (2001) Osseointegration in skeletal reconstruction and rehabilitation. J Rehabil Res Dev 38(2):175–181 7. The Academy of Prosthodontics. The glossary of prosthodontic terms (2005) J Prosthet Dent 94 (1): 10–92 8. Koković V, Lj T (2011) Preimplantation filling of tooth socket with β-tricalcium phosphate/ polylactic polyglycolic acid (β-TCP/PLGA) root analogue: clinical and histological analysis in a patient. Vojnosanit Pregl 68(4):366–371 9. Lekholm U, Zarb GA (1985) Patient selection and preparation. In: Branemark PI, Zarb GA, Albrektsson T (eds) Tissue integrated prostheses: osseointegration in clinical dentistry. Quintessence Publishing, Chicago, pp 199–209 10. Gulsahi Ayse (2011) Bone quality assessment for dental implants, implant dentistry – the most promising discipline of dentistry. In: Ilser Turkyilmaz (ed) pp 437–452. ISBN 978-953-307481-8. InTech, Available from: http://www.intechopen.com/books/implant-dentistry-the-mostpromising-discipline-of-dentistry/bone-qualityassessment-for-dental-impl 11. Nazirkar G, Singh S, Dole V, Nikam A (2014) Effortless effort in bone regeneration: a review. J Int Oral Health 6(3):120–124 12. Horowitz R, Holtzclaw D, Rosen PS (2012) A review on alveolar ridge preservation following tooth extraction. J Evid Based Dent Pract 12(Suppl 3):149–160. doi:10.1016/S1532-3382(12) 70029-5 13. Thor A, Ekstrand K, Baer RA, Toljanic JA (2014) Three-year follow-up of immediately loaded implants in the edentulous atrophic maxilla: a study in patients with poor bone quantity and quality. Int J Oral Maxillofac Implants 29(3):642–649. doi:10.11607/jomi.3163 14. Ivanov SY, Mukhametshin RF, Muraev AA, Solodkaya DV (2013) Synthetic materials used for the substitution of bone defects. Critical review. Ann Oral Maxillofac Surg 1(1):4 15. Palmer RM, Howe LC, Paul J (2012) Implants in clinical dentistry, 2nd edn. Informa Healthcare, 37–41 Mortimer Street, London 16. Brener D (2006) The mandibular ramus donor site. Aust Dent J 51(2):187–190

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17. Toscano N, Shumaker N, Holtzclaw D (2010) The art of block grafting. A review of the surgical protocol for reconstruction of alveolar ridge deficiency. J Implant Adv Clin Dent 2(2):45–66 18. Esposito M, Grusovin MG, Felice P, Karatzopoulos G, Worthington HV, Coulthard P (2009) Interventions for replacing missing teeth: horizontal and vertical bone augmentation techniques for dental implant treatment. Cochrane Database Syst Rev Issue 4. Art. No.: CD003607. doi:10.1002/14651858.CD003607.pub4 19. De Carvalho PS, Vasconcellos LW, Pi J (2000) Influence of bed preparation on the incorporation of autogenous bone grafts: a study in dogs. Int J Oral Maxillofac Implants 15:565–570 20. Santos PL, Gulinelli JL, Telles CS, Betoni WJ, Okamoto R, Buchignani VC, Queiroz TP (2013) Review article bone substitutes for peri-implant defects of postextraction implants. Int J Biomater Article ID 307136, 7 pages 21. Hallman M, Thor A (2000) Bone substitutes and growth factors as an alternative ⁄ complement to autogenous bone for grafting in implant dentistry. Periodontology 47(2008):172–192 22. Trejo PM, Weltman R, Caffesse R (2000) Treatment of intraosseous defects with bioabsorbable barriers alone or in combination with decalcified freeze-dried bone allograft: a randomized clinical trial. J Periodontol 71(12):1852–1861 23. Brkovic B, Prasad HS, Konandreas G, Radulovic M, Antunovic D, Sándor GKB, Rohrer MD (2008) Simple preservation of a maxillary extraction socket using beta-tricalcium phosphate with type I collagen: preliminary clinical and histomorphometric observations. J Can Dent Assoc 74(6):523–528 24. Dimova Cena (2014) Socket preservation procedure after tooth extraction. Key Eng Mater 587: 325–330. # (2014) Trans Tech Publications, Switzerland. doi:10.4028/www.scientific.net/ KEM.587.325 25. Byrne G (2011) Socket preservation of implant sites. A critical summary of Ten Heggeler JMAG, lot DE, Van der Weijden GA., Effect of socket preservation therapies following tooth extraction in non-molar regions in humans: a systematic review (published online ahead of print Nov. 22, 2010). Clin Oral Implants Res 22(8):779–788 26. Van der Stok J, Van Lieshout EMM, El-Massoudi Y, Van Kralingen GH, Patka P (2011) Bone substitutes in the Netherlands – a systematic literature review. Acta Biomater 7:739–750 27. Dorozhkin SV (2009) Nanodimensional and nanocrystalline apatites and other calcium rthophosphates in biomedical engineering, biology and medicine. Materials 2 (4):1975–2045 28. Strnad Z, Strnad J, Povysil C, Urban K (2000) Effect of plasma sprayed hyroxyapatite coating on the osteoconductivity of commercially pure titanium implants. Int J Oral Maxillofac Implants 15:483–490 29. Pikos MA (2005) Mandibular block autografts for alveolar ridge augmentation. Atlas Oral Maxillofac Surg Clin N Am 13:91–107 30. Gellrich NC, Held U, Schoen R, Pailing T, Schramm A, Bormann KH (2007) Alveolar zygomatic buttress: a new donor site for limited preimplant augmentation procedures. J Oral Maxillofac Surg 65(2):275–280 31. Liu J, Kerns David G (2014) Mechanisms of guided bone regeneration: a review. Open Dent J 8:56–65 32. Giardino R, Aldini NN, Fini M, Giavaresi G, Torricelli P (2002) Guided tissue regeneration in dentistry. J Trauma 52(5):933–937 33. Bottino MC, Thomas V, Schmidt G, Vohra YK, Chu TM, Kowolik MJ, Janowski GM (2012) Recent advances in the development of GTR/GBR membranes for periodontal regeneration–a materials perspective. Dent Mater 28(7):703–721 34. Nyman S (1991) Bone regeneration using the principle of guided tissue regeneration. J Clin Periodontol 18:494–498 35. Farzad M, Mohammadi M (2012) Guided bone regeneration: a literature review. J Oral Health Oral Epidemiol 1(1):3–18 36. Zhang Y, Zhang X, Shi B, Miron RJ (2013) Membranes for guided tissue and bone regeneration. Ann Oral Maxillofac Surg 1(1):10

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37. Sculean A, Nikolidakis D, Schwarz F (2008) Regeneration of periodontal tissues: combinations of barrier membranes and grafting materials – biological foundation and preclinical evidence: a systematic review. J Clin Periodontol 35(Suppl 8):106–116 38. Rakhmatia YD, Ayukawa Y, Furuhashi A, Koyano K (2013) Current barrier membranes: titanium mesh and other membranes for guided bone regeneration in dental applications. J Prosthodont Res 57(1):3–14 39. Duskova M, Leamerova E, Sosna B, Gojis O (2006) Guided tissue regeneration, barrier membranes and reconstruction of the cleft maxillary alveolus. J Craniofac Surg 17 (6):1153–1160 40. Bunyaratavej P, Wang HL (2001) Collagen membranes: a review. J Periodontol 72(2):215–229 41. Patino MG, Neiders ME, Andreana S, Noble B, Cohen RE (2002) Collagen as an implantable material in medicine and dentistry. J Oral Implantol 28(5):220–225 42. McCabe J, Yan Z, Al Naimi O, Mahmoud G, Rolland S (2011) Smart materials in dentistry. Aust Dent J 56:3–10. doi:10.1111/j.1834-7819.2010.01291 43. Dori F, Huszar T, Nikolidakis D, Arweiler NB, Gera I, Sculean A (2007) Effect of platelet-rich plasma on the healing of intra-bony defects treated with a natural bone mineral and a collagen membrane. J Clin Periodontol 34(3):254–261 44. McAllister BS, Haghighat K (2007) Bone augmentation techniques. J Periodontol. doi:10.1902/ jop.2007.060048, 377 45. Ali SA, Karthigeyan S, Deivanai M, Kumar A (2014) Implant rehabilitation for atrophic maxilla: a review. J Indian Prosthodont Soc 14(3):196–207. doi:10.1007/s13191-014-0360-4, Epub 2014 Apr 22 46. Bastos AS, Spin-Neto R, Conte-Neto N, Galina K, Boeck-Neto RJ, Marcantonio C, Marcantonio E, Marcantonio E Jr (2014) Calvarial autogenous bone graft for maxillary ridge and sinus reconstruction for rehabilitation with dental implants. J Oral Implantol 40 (4):469–478. doi:10.1563/AAID-JOI-D-11-00090 47. Caudry S, Landzberg M (2013) Lateral window sinus elevation technique: managing challenges and complications. J Can Dent Assoc 79:d101 48. Mazor Z, Peleg M, Garg AK, Chaushu G (2000) The use of hydroxyapatite bone cement for sinus floor augmentation with simultaneous implant placement in the atrophic maxilla. A report of 10 cases. J Periodontol 71:1187–1194 49. Chiapasco M, Casentini P, Zaniboni M (2009) Bone augmentation procedures in implant dentistry. Int J Oral Maxillofac Implants 24(Suppl):237–259

CAD-CAM Processing for All Ceramic Dental Restorations

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Alexandru Eugen Petre

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Dental Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Overview of All-Ceramic Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Success of Prosthetic Restorations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Dental CAD/CAM Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Characteristics of All-Ceramic CAD-CAM Restorations . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

High-strength ceramic materials such as polycrystalline alumina and zirconiabased ceramics can be used for the fabrication of various implant- or teethsupported prostheses. Zirconia based ceramics can be successfully used for multiple unit – long span bridges or allow for the thickness reduction of the restorations in non-critical bio-mechanical situations. These materials have been introduced in dental practice, along with new technologies like the computerassisted design/computer-assisted manufacturing (CAD/CAM) of dental restorations. Keywords

Dental ceramics • High-strength ceramics • Computer-assisted design/computerassisted manufacturing (CAD/CAM) of dental restorations

A.E. Petre (*) Department of Prosthodontics, Discipline of Fixed Prosthodontics and Dental Occlusion, University of Medicine and Pharmacy “Carol Davila”, Bucharest, Romania e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_49

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Introduction Fixed and removable dental prostheses can be produced using a plethora of dental materials and laboratory technologies. Classically, the longevity of dental prostheses has been accomplished mainly by the use of metallic alloys and precision casting technologies. However, due to the past decades’ public demand for aesthetics and improved biocompatibility, an active search for metal-free prostheses has been pursued. Traditionally, the best aesthetic results for dental prostheses are obtained with glassceramics like the feldspathic porcelain, but their use for metal-free restorations is limited by their limited strength especially for support (premolar and molar) area crowns and fixed partial dentures (dental bridges). In this process many high-strength ceramic materials such as polycrystalline alumina and zirconia-based ceramics can be traced back in the late 1960s. Consequently, several types of zirconia were used for the fabrication of dental implants and various restorations and parts for implant- or teethsupported prostheses. Zirconia based ceramics can be successfully used for multiple unit – long span bridges or allow for the thickness reduction of the restorations in non-critical bio-mechanical situations. These materials have been introduced in dental practice, along with new technologies like the computer-assisted design/computerassisted manufacturing (CAD/CAM) of dental restorations. In this chapter we present (1) dental ceramics, (2) dental CAD/CAM systems, (3) clinical characteristics of all-ceramic CAD-CAM restorations, and (4) further developments for all-ceramic restorations and manufacturing technologies.

Dental Ceramics Ceramic materials are used in dentistry for over 200 years. On May 11, 1791, following the ideas of the pharmacist Alexis Duch^ateau, Nicolas Dubois de Chémant (1753–1824), a French dentist, patents in England the “teeth of mineral paste.” In 1889 Charles H. Land patented the all-porcelain jacket crown. Since then, mechanical strength, aesthetics, and stability were proven as the main advantages of dental ceramics, but the fragility caused by their brittle nature made these materials rarely used without a metal substructure – first developed by Abraham Weinstein in the late 1950s. Metal-ceramic restorations are until today the first choice for fixed dental prostheses as they blend in a convenient and durable way – marginal adaptation, mechanical strength, biocompatibility, and aesthetics [11, 15].

Overview of All-Ceramic Materials In 1965, MacLean and Hughes developed metal-free crowns from an alumina reinforced porcelain core with a flexural strength of 120–150 MPa [16], and after that several classes of high-strength ceramics have been fabricated. An overview of these materials is presented in Table 1 as opposed to the ISO specifications for ceramics used for fixed prosthetic restorations (Table 2).

Commercial products

Vitadur ® (Vita), Vitablocs ® Mark II (Vita), Optec ® HSP (Jeneric/Pentron), Mirage™ (Chameleon Dental Products)

Material Feldspathic porcelain

Reinforced feldspathic porcelain (aluminous, leucite, fibers) Alumina, leucite, fibers

Crystalline phase Amorphous glass

Powder slurries for the layering and firing technique hot press

Processing method Powder slurries for the layering and firing technique

Table 1 Types of dental ceramics vs. dental enamel and dentine – modified after [22] Physical properties: σ, flexural strength; KIc, fracture toughness; H, hardness; E, Young’s modulus; CTE, coefficient of thermal expansion σ, 60–70 MPa; KIc, 0.92–1.26 MPa.m½; E, 70 GPa; H, 6 GPa; CTE, varying depending on application σ, 120–150 MPa; KIc, 1.5 MPa.m½; CTE, 16.0–17.3  106 K1 (multi-indication alloys); CTE, 13.8– 15.2  106 K1 (high gold content, reduced precious metal content, palladium-based and non-precious metal alloys); CTE, 16.0– 17.5  106 K1 (goldpalladium-silver alloys); CTE, 7.2–7.9  106 K1 (alumina, spinell, and zirconia glass-infiltrated ceramics); CTE, 9.0–10.5  106 K1 (zirconia)

CAD-CAM Processing for All Ceramic Dental Restorations (continued)

Resin-bonded laminate veneers and crowns, veneering of multiunit bridges

Clinical indication Resin-bonded laminate veneers and veneering of metallic and ceramic cores

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IPS Empress ®, IPS Empress ® CAD (Ivoclar), Optimum Pressable Ceramic OPC ® (Jeneric/ Pentron), Finesse ® (Dentsply), Authentic ® (Jensen Dental), PM™ (Vita) IPS Empress II ®, now IPS e.max ® (Ivoclar)

Leucite glassceramic

Lithium disilicate glass-ceramic

Commercial products Dicor ® (Corning Inc.) (Dentsply), Cera Pearl ® (Kyocera Corp)

Material Glass-ceramics (fluormica tetrasilicic mica, apatite glassceramic)

Table 1 (continued)

Hot press CAD/CAM

Hot press CAD/CAM

Li2Si2O5

Processing method Castable CAD/CAM

KAlSi2O6 (tetragonal phase)

Crystalline phase K2Mg5Si8 O20F4

σ, 360/400 MPa; KIc, 2.25/ 2.75 MPa. m½; H, 5.8 GPa (Vickers); E, 95 GPa; CTE, 10.2  106 K1 (100–400  C). CTE, 10.5  106 K1 (100–500  C)

Physical properties: σ, flexural strength; KIc, fracture toughness; H, hardness; E, Young’s modulus; CTE, coefficient of thermal expansion σ, 150 MPa, KIc, 1.4–1.5 MPa.m½; H, 362 MPa (Knoop); E, 70.3 GPa; CTE, 7.2  106 K1; S, 127/147 MPa; KIc, 1.4–1.5 MPa.m½; H, 3.3–3.5 GPa (Knoop); E, 68 GPa; CTE, 6.4  106 K1 σ, 160 MPa; KIc, 1.3 MPa. m½; H, 6.2 GPa (Vickers); E, 65 GPa; CTE, 16.6  106 K1 (100–400  C); CTE, 17.5  106 K1 (100–500  C)

Resin-bonded laminate veneers, inlays and onlays, crowns, bridges in the anterior region up to premolars

Resin-bonded laminate veneers, inlays, onlays, and crowns

Clinical indication Resin-bonded laminate veneers, anterior crowns, posterior inlays

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Al2O3 (hexagonal phase) ZrO2 (yttriastabilized tetragonal phase) Al2O3 (hexagonal phase)

Vita In-Ceram ® Alumina

Vita In-Ceram ® Zirconia

Procera ® (Nobel Biocare)

Glass-infiltrated alumina

Glass-infiltrated zirconia

Pure alumina (corundum)

ZrO2 (yttriastabilized tetragonal phase) 90 % Ca5(PO4)3OH (hydroxyapatite)

MgAl2O4

Vita In-Ceram ® Spinell

Glass-infiltrated spinell

Yttria-stabilized tetragonal zirconia polycrystals (3Y-TZP) Enamel

Ca5(PO4)3F

IPS e.max ® Ceram (Ivoclar)

Fluorapatite glassceramic

CAD/CAM

CAD/CAM

Slip casting CAD/CAM

Slip casting CAD/CAM

Slip casting CAD/CAM

Powder slurries for the layering and firing technique

σ, 500–700 MPa; KIc, 4.5 MPa.m½; E, 270–380 GPa; H, 12 GPa; CTE, 7.106 K1 σ, 900–1400 MPa; KIc, 6–10 MPa.m½; E, 205–210 GPa; H, 13.9 GPa; CTE, 10.5  106 K1 σ, 261–288 MPa (10 MPa if not supported by dentin); KIc, 0.6–1.5 MPa. m½; E, 70–100 GPa; H, 3–5 GPa

σ, 400 MPa; KIc, 2.7 MPa. m½; E, 185 GPa; CTE, 7.7  106 K σ, 500 MPa; KIc, 3.9 MPa. m½; E, 280 GPa; CTE, 7.4  106 K1 σ, 600 MPa; KIc, 4.4 MPa. m½; E, 258 GPa; CTE, 7.8  106 K1

Slightly higher strength and fracture toughness of the pressable vs. CAD/CAM ceramic σ, 90 MPa; H, 5.4 GPa (Vickers); CTE, 9.5  106 K1 (100–400  C)

CAD-CAM Processing for All Ceramic Dental Restorations (continued)

Laminate cores, anterior and posterior crowns and bridges, abutments and implant bridges

Laminate cores, crowns, 4-unit bridges

Anterior and posterior crowns and bridges

Laminate cores, crowns, abutments, anterior bridges

Resin-bonded laminate veneers, veneering of lithium disilicate glassceramic and zirconia frameworks Inlays and anterior crowns

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Material Dentin

Table 1 (continued)

Commercial products

Crystalline phase 70 % Ca5(PO4)3OH (hydroxyapatite) Processing method

Physical properties: σ, flexural strength; KIc, fracture toughness; H, hardness; E, Young’s modulus; CTE, coefficient of thermal expansion σ, 232–305 MPa; KIc, 3.1 MPa.m½; E, 15–30 GPa; H, 0.6 GPa

Clinical indication

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Table 2 Classification of ceramics for fixed prostheses by intended clinical use (ISO 6872:2008)

Class 1

2

3 4

5 6

Recommended clinical indications (a) Aesthetic ceramic for coverage of a metal or a ceramic substructure (b) Aesthetic ceramic: single-unit anterior prostheses, veneers, inlays, or onlays (a) Aesthetic ceramic: adhesively cemented, single-unit, anterior or posterior prostheses (b) Adhesively cemented, substructure ceramic for single-unit anterior or posterior prostheses Aesthetic ceramic: nonadhesively cemented, single-unit, anterior or posterior prostheses (a) Substructure ceramic for nonadhesively cemented, single-unit, anterior or posterior prostheses (b) Substructure ceramic for three-unit prostheses not involving molar restoration Substructure ceramic for three-unit prostheses involving molar restoration Substructure ceramic for prostheses involving four or more units

Mechanical and chemical properties Chemical solubility Flexural strength maximum, mg minimum (mean), cm2 MPa 50 100

100 100

100 2000

300

100

300

2000

500

2000

800

100

Clinical Success of Prosthetic Restorations It has been proposed a quantitative rating scale for which may be used for evaluating clinical success for new types of restorations [2]: 1. Superior performance: Survival of all FDPs (100 %) for at least 5 years and a success rate of 95–100 % 2. Excellent performance: Survival of 95–100 % of all FDPs for at least 5 years and a success rate of 90–95 % 3. Good performance: Survival of 90–95 % of restorations for at least 5 years and a success rate of 90–95 % 4. Poor performance: Survival of less than 90 % of restorations or a success rate of less than 90 %

Survival of Metal-Ceramic Restorations There are several reports about survival rates of metal-ceramic restorations: at 5-year 95.6 % for metal-ceramic single crowns [20] 94.4 % for metal-ceramic fixed partial dentures (FPD) [21]; 10-year survival rates of 92 % and 74 % 15-year survival for FPD [24] 18-year survival rates of 75 % for crowns on vital teeth and 79 % on nonvital teeth were found in a retrospective evaluation [8] or 78 % for FPDs [18].

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Survival of All-Ceramic Restorations For anterior teeth, it is accepted that clinicians may select from any all-ceramic system for any restoration type. For example, the survival of lithium disilicate restorations shows a cumulative survival rate over a 10-year period of 96.7 % [19]. For molar restorations there are few all-ceramic systems that provide predictable long-term success. Clinical complications of all-ceramic multiunit fixed dental prostheses are frequent, even for those with enlarged connectors [14]. The mediumterm survival of lithium disilicate restorations for fixed partial dentures is only 70.9 % with the lateral areas being the most affected by failure [19]. Kaplan-Meier survival rates were reported for 3-unit fixed partial dentures for end points ranging from 1 to 10 years, with a mean end point of approximately 5.6 years [14]. Hydrothermal aging of zirconia can cause significant transformation from tetragonal to monoclinic crystal structure, which results in a decrease in mechanical properties [10].

Dental CAD/CAM Systems In 1972, François Duret filed in France a patent application and in 1973 completed a doctoral thesis on the theme “le rayon laser en dentistrie et prothèse.” Duret invented optical impression in dental prosthetics and thereby created the possibility of computer aided designing and fabrication of dental prostheses [9]. Werner Mörmann (Switzerland) and Marco Brandestini (Italy) developed the first commercial system – the CEREC, acronym for Chairside Economical Restoration of Esthetic Ceramics [17]. The system allows dentists to chairside fabricate single ceramic restorations in one appointment, and it has been upgraded several times in the past decades. In the 1980s Diane Rekow developed a CAD-CAM system in the United States. In 1983 Matts Andersson from Sweden developed the Procera method of manufacturing high-precision dental crowns which was acquired in 1989 by Nobelpharma, later Nobel Biocare (http://corporate. nobelbiocare.com/en/our-company/history-and-innovations/ at 13.11.2014). Since the early 1980s, the CAD-CAM systems continued to develop, but in the last years, the growth was exponential (Table 3). Table 3 PubMed search result in number of publications in the past two decades with CAD-CAM keyword and DENTISTRY as major topic

Year 2012–2013 2010–2011 2008–2009 2006–2007 2004–2005 2002–2003 2000–2001 1998–1999 1996–1997 1994–1995

No. of references 467 375 307 211 161 140 109 85 47 48

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Currently, commercial CAD-CAM systems allow for the fabrication of practically all types of fixed restorations: anatomical copings and bridge frameworks, full anatomical crown and bridges, inlays, onlays, inlay bridges, veneers, multilayer crowns and bridges [7], wax-up, digital temporaries, post and cores, telescopes, customized implant abutments, bars and bridges, full dentures, removable partial dentures, custom trays and models for classic technology, occlusal splints, orthodontic analysis and appliances, and surgical guides for dental implants (http://www.3shapedental.com/restoration/dental-lab/digitallab/next-step-for-cadcam-dentistry/ at 13.11.2014). Comparing virtual models can also compare and measure teeth movements during different dental treatments [4]. The main advantages of the CAD-CAM technology are the speed and the predictable quality of the final restorations at increasingly competitive costs due to the time-saving procedures and increased commercial offer [25]. There are three ways in which CAD-CAM technology can be used: ALL LAB 1. Classical impression (analogue, with impression materials) 2. Model fabrication with removable dies and articulation 3. Model digitization and virtual model creation In this variant, clinical work is completely separate from CAD-CAM technology, used exclusively in the dental lab. The virtual model and the resulting dental prosthesis are likely to include all the errors from the impression and modelpouring steps. On the other hand, the existence of a physical model makes possible the try-in and potential adjustments as well as subsequent phases such as the traditional aesthetic ceramic veneering of metal or zirconia frameworks. Due to the versatility in accommodating both fabrication methods, the all-lab CAD-CAM technology is for the moment the most commonly used. IMPRESSION SCANNING 1. Classical impression 2. Digitization of the impression and transformation into a virtual model without pouring a model Impression scanning can be done either in the dental office or in the laboratory and has the advantage that it avoids casting and preparing the model and therefore the possible errors of this phase. Impression scanners can be dedicated devices for in-office or in-lab use or classic laboratory scanners with an impression scanning add-in. Any impression scanner can digitize also models, but the dedicated lab scanners with the impression scanning option are usually more expensive than their model-only counterparts. Scanning mainly negative reliefs is sometimes difficult and often requires shortening impression borders or the use of antireflective powder before digitization, which might pose problems with custom or universal metal trays with high borders that cannot be shortened without risking deformation of the impression.

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Pouring models in the scanned impression is of limited utility for the subsequent lab phases. The prostheses fabricated after digitized impressions present an uncertain fit on these models mainly because of the soft tissue areas. Whenever the CAD-CAM process does not start from a poured model but there is the need for further steps like the aesthetical veneering, there are two options: – Producing a physical model from the virtual one through a model designing software. – Doing an intraoral try-in and making a position impression of the framework; further stages are then continued in the traditional dental lab. Dedicated impression scanners could be an economic solution for digitalization in prosthetic dentistry as they allow for an in-office creation of the virtual model and a faster clinical feedback regarding the quality of the preparations. Besides, there are cases in which intraoral scanning is not feasible because of large edentulous spans, highly mobile mucosa or other mobile anatomical structures in the close vicinity of the preparations, or deep and narrow interdental spaces. As a consequence of the need, in some cases, for traditional impression, it is likely that further developments of intraoral scanners could include the option for impression scanning. INTRAORAL SCANNING 1. Intraoral scanners enable the work in an entirely virtual environment which brings numerous advantages: – Impression accuracy and fit of prosthetic structures produced in this technology are considered within clinically acceptable limits as the flaws in the classical impressions are notoriously frequent [5]. In one study, the mean marginal (internal) discrepancies for restorations produced after several intraoral scanners were significant but clinically acceptable: iTero 90 (92) μm, TRIOS 128 (106) μm, CEREC AC with Bluecam 146 (84) μm, and Lava COS 109 (93) μm [23]. – The control of the preparations on the scanner desktop, at an important magnification, enables the clinician to validate the preparations in real time (Fig. 1). – Corrected preparations or inaccurate reproduced areas from the digital model can be selectively removed and re-added to the digital impression, which makes complete retaking of the impression never necessary, as in the classical counterparts (Fig. 2). – In the cases in which one takes the conformative approach of the occlusion, intraoral scanning allows immediate validation of the reference mandibulomaxillary position – by real-time observation of the coincidence of intraoral occlusal contacts with the virtual models (Fig. 3). – The incorporation in the virtual model of the pre-preparation geometry is very simple; it can also be scanned and used in the design process of the provisional restorations or mock-up’s geometry (Figs. 3 and 4). – Transport and storage do not change the virtual model as in the case of classical impressions.

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Fig. 1 Intraoral direct scanning of a pre-preparation situation for the prosthetic restoration of teeth 22 23 24

Fig. 2 Intraoral direct scanning of the preparations for the prosthetic restoration of teeth 22 23 24

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Fig. 3 Scanning can be visualised either in true color display for the correct tissue identification or color evaluation or monochromatic for a better evaluation af the scanning process

Fig. 4 On the digital models one can make several measurements and sectional views for better clinical evaluation

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– The oldest commercial CAD-CAM system includes a milling machine that provides immediate, ceramic or composite blocks of single prosthetic elements and other milling machines small enough for an in-office use which have been developed very recently and are now commercially available (https://www. amanngirrbach.com/products/milling-cam/ceramill-mikro/ at 13.11.2014). – Some commercial intraoral scanners are able to make color measuring (http:// www.3shapedental.com/news/using-shade-measurement/ at 13.11.2014). – The online or direct connection communication with the dental lab is largely facilitated by the built-in modules of the scanner and design software. Disadvantages of CAD-CAM technology: – The high initial investment and periodic license/support /upgrade fees. – The need for some computer literacy and specific training for scanning, designing, and machining applications even though the software is often quite intuitive. – Compatibility and interconnection problems when using a mixed environment, e.g., different types of intraoral scanner, design software, and cam software and hardware. – The need of producing high-precision physical models from the digital data by printing or milling which add to the restoration cost; alternatively the model can be produced in a traditional manner, after a position impression of the milled framework. – Few available high aesthetic solutions for multiple restorations. – Lack of a true virtual articulator: the current systems require mounting the working models in a real articulator and then transferring the parameters in the virtual environment which excludes impression and/or intraoral scanning from their use without producing a physical model; there are several reports of experimental methods of transferring the location of the maxillary dental arch from the patient directly to a virtual articulator (virtual facebow) [26–28].

Clinical Characteristics of All-Ceramic CAD-CAM Restorations It has been demonstrated that milling titanium and zirconia frameworks produces some peri-implant strain. There are reports in which the zirconia frameworks produced significantly less strain than titanium. Combining the qualitative and quantitative information indicates that the implants were under vertical displacement rather than horizontal. The vertical fit was similar for zirconia (3.7 μm) and titanium (3.6 μm) frameworks; however, the zirconia frameworks exhibited a significantly finer passive fit (5.5 μm) than titanium frameworks (13.6 μm) [1]. Dental zirconia specimens sintered at 1450  C for 1 h combined good mechanical properties with the best resistance to low-temperature degradation independent of the commercial brand. Larger temperatures and times were detrimental for the structures [12].

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Table 4 Surface average roughness values after different zirconia surface treatments – modified after [3] Surface treatment Airborne-particle abrasion 125 μm Al2O3 particles applied for 10 s at 60–100 psi 9.5 % hydrofluoric acid etching for 90 s Selective infiltration etching (SIE) Experimental etching (10 min, 30 min, 60 min) methanol (800 ml), 37 % HCl (200 ml), and ferric chloride (2 g) Non-treated

Surface average roughness (S.D.) nm 7.11 (1.1) 5.23 (0.9) 26.02 (8.8) 54.22 (29.0), 81.79 (8.0), 103.02 (31.2) 6.94 (1.3)

A marginal preparation angle smaller than 60 may increase the risk of cervical chipping when using some milling systems [29]. The surface of zirconia ceramic is damaged during grinding which may affect the mechanical properties of the material. The micromotor produced a significantly higher temperature (127  C) than a high-speed handpiece (63  C) [13]. Ceramic surface polishing can be obtained by airborne-particle abrasion, polishing kits with or without polishing paste, or autoglaze at 621  C for 3 increase of 83  C/min up to 918  C for 30 s. For all ceramic types, the smoothest surfaces are obtained after autoglazing [6]. Among different cement types proposed for cementing zirconia restorations, those containing 10-MDP-based resin luting agents seem to have the best adhesive properties, but the bonds may be more effective and durable if associated with micromechanical retentions. Several surface treatments have been investigated (Table 4).

References 1. Abduo J, Lyons K, Waddell N, Bennani V, Swain M (2012) A comparison of fit of CNC-milled titanium and zirconia frameworks to implants. Clin Implant Dent Relat Res 14(Suppl 1): e20–e29 2. Anusavice KJ (2012) Standardizing failure, success, and survival decisions in clinical studies of ceramic and metal–ceramic fixed dental prostheses. Dent Mater 28(1):102–111 3. Casucci A, Osorio E, Osorio R, Monticelli F, Toledano M, Mazzitelli C, Ferrari M (2009) Influence of different surface treatments on surface zirconia frameworks. J Dent 37 (11):891–897 4. Chen H, Lowe AA, de Almeida FR, Wong M, Fleetham JA, Wang B (2008) Three-dimensional computer-assisted study model analysis of long-term oral-appliance wear. Part 1: Methodology. Am J Orthod Dentofacial Orthop 134(3):393–407 5. Christensen GJ (2005) The state of fixed prosthodontic impressions: room for improvement. J Am Dent Assoc 136(3):343–346 6. Coşkun Akar G, Pekkan G, Çal E, Eskitaşçıoğlu G, Özcan M (2014) Effects of surface-finishing protocols on the roughness, color change, and translucency of different ceramic systems. J Prosthet Dent 112(2):314–321 7. Davidowitz G, Kotick PG (2011) The use of CAD/CAM in dentistry. Dent Clin N Am 55 (3):559–570

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8. De Backer H, Van Maele G, Decock V, Van den Berghe L (2007) Long-term survival of complete crowns, fixed dental prostheses, and cantilever fixed dental prostheses with posts and cores on root canal-treated teeth. Int J Prosthodont 20(3):229–234 9. Duret F, Preston JD (1991) CAD/CAM imaging in dentistry. Curr Opin Dent 1(2):150–154 10. Flinn BD, deGroot DA, Mancl LA, Raigrodski AJ (2012) Accelerated aging characteristics of three yttria-stabilized tetragonal zirconia polycrystalline dental materials. J Prosthet Dent 108 (4):223–230 11. Goodacre CJ, Bernal G, Rungcharassaeng K, Kan JYK (2003) Clinical complications in fixed prosthodontics. J Prosthet Dent 90(1):31–41 12. Inokoshi M, Zhang F, De Munck J, Minakuchi S, Naert I, Vleugels J, Van Meerbeek B, Vanmeensel K (2014) Influence of sintering conditions on low-temperature degradation of dental zirconia. Dent Mater 30(6):669–678 13. İşeri U, Özkurt Z, Yalnız A, Kazazoğlu E (2012) Comparison of different grinding procedures on the flexural strength of zirconia. J Prosthet Dent 107(5):309–315 14. Land MF, Hopp CD (2010) Survival rates of all-ceramic systems differ by clinical indication and fabrication method. J Evid Based Dent Pract 10(1):37–38 15. Libby G, Arcuri MR, LaVelle WE, Hebl L (1997) Longevity of fixed partial dentures. J Prosthet Dent 78(2):127–131 16. McLean JW, Hughes TH (1965) The reinforcement of dental porcelain with ceramic oxides. Br Dent J 119(6):251–267 17. Mormann WH (2006) The evolution of the CEREC system. J Am Dent Assoc 137 (Suppl):7s–13s 18. Napankangas R, Raustia A (2011) An 18-year retrospective analysis of treatment outcomes with metal-ceramic fixed partial dentures. Int J Prosthodont 24(4):314–319 19. Pieger S, Salman A, Bidra AS (2014) Clinical outcomes of lithium disilicate single crowns and partial fixed dental prostheses: A systematic review. J Prosthet Dent 112(1):22–30 20. Pjetursson BE, Br€agger U, Lang NP, Zwahlen M (2007) Comparison of survival and complication rates of tooth-supported fixed dental prostheses (FDPs) and implant-supported FDPs and single crowns (SCs). Clin Oral Implants Res 18:97–113 21. Sailer I, Pjetursson BE, Zwahlen M, H€ammerle CHF (2007) A systematic review of the survival and complication rates of all-ceramic and metal–ceramic reconstructions after an observation period of at least 3 years. Part II: fixed dental prostheses. Clin Oral Implants Res 18:86–96 22. Saint-Jean SJ (2014) Chapter 12 – Dental glasses and glass-ceramics. In: Shen JZ, Kosmač T (eds) Advanced ceramics for dentistry. Butterworth-Heinemann, Oxford, pp 255–277 23. Schaefer O, Decker M, Wittstock F, Kuepper H, Guentsch A (2014) Impact of digital impression techniques on the adaption of ceramic partial crowns in vitro. J Dent 42(6):677–683 24. Scurria MS, Bader JD, Shugars DA (1998) Meta-analysis of fixed partial denture survival: Prostheses and abutments. J Prosthet Dent 79(4):459–464 25. Service, U. D. E. C. (2011). Synopsis of CAD/CAM systems 26. Solaberrieta E, Mínguez R, Barrenetxea L, Etxaniz O (2013) Direct transfer of the position of digitized casts to a virtual articulator. J Prosthet Dent 109(6):411–414 27. Solaberrieta E, Minguez R, Barrenetxea L, Sierra E, Etxaniz O (2013) Novel methodology to transfer digitized casts onto a virtual dental articulator. CIRP J Manuf Sci Technol 6(2):149–155 28. Solaberrieta E, Otegi JR, Minguez R, Etxaniz O (2014) Improved digital transfer of the maxillary cast to a virtual articulator. J Prosthet Dent 112(4):921–924 29. Giannetopoulos S, van Noort R, Tsitrou E (2010) Evaluation of the marginal integrity of ceramic copings with different marginal angles using two different CAD/CAM systems. J Dent 38(12):980–986

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Florin Miculescu, Lucian Toma Ciocan, Marian Miculescu, Andrei Berbecaru, Josep Oliva, and Raluca Monica Comăneanu

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Overview of Dental Prostheses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . All-Ceramic Dental Prostheses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Metal-Ceramic Dental Prostheses . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Dental Prostheses Failure Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Causes of Failure for Ceramic Dental Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Causes of Failure for Metal-Ceramic Dental Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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To accurately assess the real causes of failure of dental prostheses, the assemblies, the materials, and their surfaces should be analyzed to highlight the exact morphological and compositional aspects, which is also the aim of this chapter.

F. Miculescu (*) • M. Miculescu • A. Berbecaru Faculty of Materials Science and Engineering, University Politehnica of Bucharest, Bucharest, Romania e-mail: [email protected]; [email protected]; [email protected] L.T. Ciocan Dental Medicine Faculty, “Carol Davila” University of Medicine and Pharmacy from Bucharest, Bucharest, Romania e-mail: [email protected] J. Oliva Clinica Oliva Dental, Barcelona, Spain e-mail: [email protected] R.M. Comăneanu Faculty of Dental Medicine, Titu Maiorescu University, Bucharest, Romania e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_56

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The experimental demonstration of the main causes and types of failure of dental prostheses consists of analysis of removed dental implants, fixed, single, and multitooth prostheses. These were collected from dental laboratories after their failure. Within the presented analyses prospects related to the cross-section sample microstructure or prosthetic surface (where the failure appeared due to surface defects) were targeted. Restorative dental materials include representatives of the main classes of metallic materials, polymers, ceramics, and composites. Most of the restorations are described by their physical, chemical, and mechanical parameters resulted from laboratory tests. Improvements of these characteristics may seem attractive for laboratory studies, but the real test of the materials’ performance is done in the mouth cavity environment. Although, at this point, the dental materials became of high performance, many types of dental prostheses fail fast or after a certain period of time, smaller than the estimated one. Keywords

Failure analysis • Dental prostheses • Dental ceramics • All-ceramic dental prostheses • Metal-ceramic • SEM • Surface analysis • Morphology • Interface • Fracture analysis

Introduction During the last century, dentistry has significantly transformed, becoming a very complex subject area. With this transformation, materials obtained a crucial role in every stage of dental treatment. Dental materials are in a continuous development as the technology progresses. Also, an increasingly important requirement of the patients is the natural appearance of the teeth, which leads to a higher focus on the aesthetic component [1–4]. Developments in the field of materials science and engineering have dramatically changed the way we look on the human anatomy replacing components, and thus, the dental restoration materials represent the foundation on the tooth structure replacement. The form and function of the dental prostheses contribute enormously to the quality of life. Proper functioning of the mouth elements – teeth and soft tissue – is essential for speech, chewing, swallowing, and breathing. Dental restorative materials allow the reconstruction of hard dental tissues. Due to their success in long-term use, patients often expect dental restoration to have more quality than the natural teeth. The use of material science in the dental field is unique because the mouth has a high complexity. One can find bacteria presence, high stresses, variable pH, and a warm, fluid environment. The oral cavity is considered to be as one of the harshest environments for a material in the body [2, 4]. Restorative dental materials include representatives of the main classes of metallic materials, polymers, ceramics, and composites. Most of the restorations are

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described by their physical, chemical, and mechanical parameters resulted from laboratory tests. Improvements of these characteristics may seem attractive for laboratory studies, but the real test of the materials’ performance is done in the mouth cavity environment. Although, at this point, the dental materials became of high performance, many types of dental prostheses fail fast or after a certain period of time, smaller than the estimated one. Therefore, to accurately assess the real causes of failure of dental prostheses, the assemblies, the materials, and their surfaces should be analyzed to highlight the exact morphological and compositional aspects, which is also the aim of this chapter [5–8]. In this respect, within this chapter the experimental demonstration of the main causes and types of failure of dental prostheses consists of analysis of removed dental implants, fixed, single, and multitooth prostheses. These were collected from dental laboratories after their failure. Within the presented analyses prospects related to the cross-section sample microstructure or prosthetic surface (where the failure appeared due to surface defects) were targeted. Prosthesis surface characterization was performed without the special preparation of the samples. For the cross-section analysis, the samples were kept cool (the temperature during the polymerization and preparation of the samples did not exceed 45  C) in a two-component polymeric resin, then were sectioned to reveal the sagittal and transverse planes, grounded and polished until ogling surfaces were obtained. Sample preparation was made using a Buehler metallographic sample preparation system. Electron microscopic analyses of the prosthesis were performed with scanning electron microscope Philips XL 30 ESEM TMP, equipped with an EDS EDAX Saphire spectrometer, at University Politehnica of Bucharest, Department of Metallic Materials Science and Physical Metallurgy. The chemical compositions were analyzed on each component and each material constituent component.

Overview of Dental Prostheses Dental prostheses are natural bodies made of special materials in order to restore morphofunctional dent alveolar tissues. Generally, dentistry involves the use of two such categories: dental prostheses, which replace the missing tissue morphology and restore the functions of the affected maxillary and also the use of prosthetic device, to prevent, correct, or maintain certain ratios of the dentodental, dental-alveolar, or interarch [4]. Depending on the nature of fixity or mobility to the remaining teeth in the mouth, the dental prostheses can be fixed (aggregated to dental tissues intimately and cemented for long periods of time), movable (partially edentulous prosthetic field maintained by anchoring or use of special sliding), and mobile (total edentulous prosthetic field maintained by the phenomenon of suction, adhesion, muscle tonicity, anatomical retention) [3, 4] (Fig. 1). The fixed prostheses can be microprostheses (single-tooth implants with small dimensions – Jacket Crown) or bridges (two or more teeth) and can be metallic (high

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Fig. 1 The main types of dental prostheses: (a) fixed; (b) movable; (c) mobile

noble alloys, noble alloys and common – steel, bronze), nonmetallic, acrylic resins, and dental ceramics or metal-acrylic metal-ceramic and composites [4]. Depending on production technology, the small prostheses can be molded, stamped, embossed (stamped), and soldered (crown from two pieces), polymerized, obtained by synthesis or obtained by combined techniques, cast synthesized, cast polymerized, or cast photopolymerized. Regardless of the type of the prosthetic restoration, all sides must be polished without micropores [6, 9]. In this chapter we consider some of the most common ceramic-ceramic and metal-ceramic technologybased dental prosthesis failure possibilities.

All-Ceramic Dental Prostheses In the vast field of dentistry, ceramic materials have become among the most used materials. Ceramics are one of the oldest known materials and are represented by compounds of one or more metals with a nonmetallic element (typically oxygen). They are made of chemical and biochemical substances that are stable, resistant, hard, brittle, inert, and that do not conduct heat and electricity [9–11]. Dental practice has proved that dental ceramics are dental restorative materials that can realistically duplicate hard dental structures. Although composite resins have a similar aesthetic potential, the major difference is that dental ceramics are way

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more durable, wear resistant, and virtually indestructible in the oral environment. They are immune to the oral fluids and absolutely biocompatible. Due to their huge potential, there are many ways of research and development in dental ceramics. Ceramic applications include inlays and onlays, dental veneers, natural teeth, ceramic crowns, fixed partial dentures in the short- and long-term aesthetic component to metal crowns and bridges (metal-ceramic), artificial teeth (dentures – total or partial), columns, ceramic cores, and ceramic orthodontic appliances [1, 12]. Getting an aesthetic and durable material that can accurately reproduce missing teeth or tooth structure has always been a priority. Prior to using porcelain, dental crowns were made entirely of gold and other alloys. As the aesthetic appreciation grew, so has the use of colored resin tooth shades as a layer on the metal surface. Porcelain crowns were introduced in dentistry in the early 1900s, but back then they had many shortcomings: difficulty of manufacture, not being set well, and an easi tendency to fracture (half moon type). In the early 1960s, McLean developed a ceramic that could be deposited on metal. This led to the possibility of obtaining metal-ceramic prosthesis, which is now a significant proportion of the dental restorations [12, 13]. However, research on all-ceramic crowns continued. Although metal-ceramic restorations were a big success, they were not the final solution. The metal component, unlike the natural teeth, prevented light from reflecting and passing through, and in certain conditions, these crowns appear dense, dark, and opaque. The ideal aesthetic in this perspective would have involved the opportunity to reflect accurately the color of the dentin. Also, restoration edges appear dark, even when hidden under the gums, as they develop a bluish tint [4]. The first major breakthrough in full ceramic restoration was done in 1965, when McLean and Hughes proposed an alumina-reinforced core material, which increased the strength of porcelain. Even so, resistance was not strong enough for later use, and there is also the problem of marginal adaptation. The 1990s were the years in which all-ceramic crowns and fixed partial dentures have made major improvements. Restoration resistance increased alongside the rise of technology and with new types of porcelain. The new generation of ceramics included pourable glasses, uncompressed cores, cores that are modeled by injection, infiltrated alumina cores with high-strength glasses, CAD-CAM ceramics, etc. [13–15]. Modern all-ceramic dental restoration has partially solved one of the biggest problems of the first ceramics: low mechanical strength. However, there is still place for improvement as ceramic systems are complicated; they involve expensive manufacturing processes and equipment and dentist and dental technician extensive experience [6]. For the presentation and for fully understanding the types and causes of failure in prosthetic restorations with ceramics, we will present briefly the types of materials and possibilities of obtaining dental prostheses. Based on the sintering temperature, dental ceramics can be fused at high temperature (above 1300  C), fused at medium temperatures (1100–1300  C), fused at low temperature (850–1100  C), or fused at ultralow temperature (below 850  C). From the perspective of the manufacturing process, ceramic materials can be condensed, glass infiltrated, hot pressed, pourable, mechanically processed, or various combinations of the above processes [4, 12, 15, 16].

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Ceramic materials with possible use in dental prosthetic technologies are feldspatic porcelain, leucite-reinforced glasses, glasses based on tetrasilicon fluoride, ceramics based on lithium disilicate, alumina-reinforced ceramics, zirconia-reinforced ceramics, and ceramics reinforced with spinel. The structure of these can be vitreous, crystalline, or vitrocrystalline. The biggest disadvantage of the porcelain is that it is brittle. This liability drastically limits its use. For fixing this problem there have been developed numerous systems that prevent the formation and propagation of the cracks in the inner layer of the porcelain restoration. A possible approach involves the use of a pure alumina core on which the porcelain crown can be built. Alumina is a really hard material, opaque, which is less susceptible to cracks than porcelain. Another possible approach is the use of alumina inserts. These are small plates of alumina that are usually placed in the back of the crown to not affect the aesthetics [4]. The alumina powder may be added to the porcelain composition to achieve a significant increase in resistance. These improvements are obtained not only as a result of the superior mechanical properties of alumina but also because of the good compatibility between alumina and porcelain. The two materials have very close values of the coefficients of thermal expansion and elastic modulus. Therefore the interface region between the porcelain and the alumina is substantially free of tension and does not encourage propagation of cracks around the particles of alumina [11, 17]. Porcelain that contains alumina is called aluminous porcelain, and the designated alumina content is typically 40 %. Although aluminous porcelain has clear advantages in terms of mechanical properties, it is opaque; therefore it can only be used to achieve the inner dental restoration. It is an accepted fact, since the internal area is the one in which cracks appear, so that has to be reinforced [17–19]. Using additions of alumina for porcelain reinforcement was taken a step further by introducing sintered alumina core. For such a system, the first stage for the restoration manufacturing involves the formation of duplicated plaster dies from a special plaster. A strip of alumina is then created from aluminum powder and water on the matrix. The moist from the strip is then absorbed by the plaster, leaving behind a layer of alumina powder with the ideal thickness of 0.5 mm. This is later sintered using frittation at 1120  C for 2 h. Sintering makes the matrix's material to compress, making it easy to remove the sintered alumina core. The external surface of the core is then coated with a suspension of glass powder; after that another frittation once at 1100  C is done in order to liquefy the glass that will flow and fill the spaces between sintered alumina particles. In this case the type of glass named lanthanum aluminosilicate is used. Lanthanum reduces viscosity and assists infiltration. It also increases the refractive index of the glass and improves the translucency of the ceramic. The excess glass is then removed by sablation and fritted at 960  C to ensure a good infiltration of the glass into the alumina [16]. A more advanced development of this method is making a sintered alumina core that contains a significant amount of zirconium oxide to gain better mechanical strength and bending resistance of approximately 800 MPa. A limitation of this method is given by the quite high opaqueness of the resulted core that can be hard to

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Fig. 2 Lithium disilicate all-ceramic crown with no defects

hide with glass infiltrations, thus limiting the aesthetic qualities of the final restoration [20, 21] (Fig. 2). The ones obtained by injection and by compression represent another class of ceramic materials. These materials have been introduced in the early 1980s for manufacturing full-ceramic anterior and posterior crowns. The first commercial method that was based on this principle involved the production of a core for crown by injection, eliminating the need to use a platinum foil and improving the marginal adaptation of the crown. Noncompression properties are obtained by incorporating a substantial amount of magnesium oxide in the ceramic material. During fritation, it reacts with the alumina and forms a mixed metallic oxide named spinel. Spinel is less dense than the original mixture, and its making leads to an expansion that compensates compression during fritation. The latest pressing approach is based on disilicate lithium ceramics. The resulted piece can reach values of flexural strength similar to ceramics with sintered alumina core (300–400 MPa). This type of material is suitable for the manufacturing of three-unit bridges (fixed partial dentures) used to replace anterior or premolar teeth [4, 11]. Currently, the methods that are part of the polycrystalline ceramics and the poured glass materials category are only used for the manufacturing of unitary crowns and are substituted for other mentioned methods. The melted ceramic is centrifugally poured in the matrix at approximately 1350  C. The result is a transparent glass crown that is afterward thermally treated in an oven at 1075  C for 10 h. This thermal treatment induces a partial crystallization, creating crystals that contain K, Mg, Si oxides and significant quantities of fluorine, which have dual effect; they slowly reduce translucency and increase strength [7]. Color matching is achieved by applying to the surface a series of porcelains with different shades and refrittation. The development of new materials in which the shade can be incorporated in the crown will undermine this method. Glazes used at the surface for establishing the color can cause a really realistic effect. Unfortunately, if any adjustment is needed for the shape of the crown, then these are removed and the crown needs to be reglazed. The pouring technique gives the ability to create

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all-ceramic crowns that can be accurately adjusted and that can be strong enough to be used in the posterior region without the use of a metal substructure. Another material that is introduced in dentistry is based on yttrium – tetragonal zirconia polycrystals (Y-TZP). Yttrium oxide is mixed with zirconia oxide to create a multiphase material known by the name of partially stabilized zirconia. This form of restoration can be processed using pouring methods or can be molded from monolithic blocks of partially of fully sintered material. For these kinds of Y-TZP ceramics there have been reported really high resistance values, both at flexural (900–1200 MPa) and yield point. The transformations from the crystalline lattice resulted in the volume growth that leads to the appearance of compression tensions around the beginning of the crack points. This minimizes the possibility of crack propagation [12]. One of the most modern technologies involves the use of ceramic restorations with CAD-CAM. The short name CAD-CAM comes from Computer-Aided-Design – Computer-Aided Manufacturing. This is a powerful approach that provides patients with sustainable restoration in natural colors. The method involves the recording of an optical impression, and afterward a restoration can be made with the help of a computer. The image details will be used to build the restoration with the help of a high-precision drill that will cut a block of ceramics under computer control. This technique is quite flexible from the point of view of the restorations that can be made. Restorations with three or four complex surfaces and ceramic facets can be achieved. The optical impression is recorded with a miniature video camera that scans the tooth for approximately 10 s. The surface of the tooth needs to be clean, dry, and covered with a reflecting powder in order to maximize the quality of the image. This layer of powder needs to be flat and as thin as possible. It can capture also an optical image of the antagonist teeth in order to determine the normal pattern of teeth, thus achieving a proper restoration surface anatomy [4, 7, 9].

Metal-Ceramic Dental Prostheses The mixed metal-ceramic crown is a modern form of treatment in the field of prosthetic coronary restoration and reduced edentation solving from their indicated perspective as dental bridge aggregation elements. Metal-ceramic dental prostheses are a part of the mixed dentures category, made out of a metal structure intimal contacted to the dental blunt and an aesthetic component. The aesthetic component partly or fully covers the metal structure, and in a mixed restoration it can be represented by thermal polymerizable acrylic resins, preparing ceramic masses and composites. Metal-ceramic restorations have been developed in 1965. The molded metal core has significantly increased the strength of porcelain restorations and, shortly after, it became the most used method of ceramic restoration. A study in the 1990s showed that approximately 90 % of dental restorations are metal-ceramic based [4]. Metal-ceramic restoration methods can be classified into poured metal-ceramic restorations (feldspathic porcelain cast and noble alloys, master alloys and titanium

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castings and feldspathic porcelain cast porcelain fusion and ultralow temperature) and forged metal-ceramic restorations (gold sheet covering, platinum sheet covering). The description of a metal-ceramic restoration can be realized from more points of view: from the aesthetic point of view, from the masticator point of view, and from the biological point of view. The aesthetic point of view is ideal because of the chromatic stability, the shades of color, translucency, and permanent gloss. These features are determined by the presence of inorganic dyes (mineral oxides), by the inertia of the compositional elements, and by the compact and waterproof structure of the ceramic mass. The aesthetic aspect is durable in time, being superior as against the aesthetic aspect of the mixed metal-acrylic crowns that modifies their color shades. Because of the existence of the metallic component on which the ceramic mass is sintered and because of the physicochemical phenomenon of oxidation that causes the bonding between the two materials, it is not possible to obtain the same optical effect as the Jacket porcelain crown or of natural teeth. For the correction of the aesthetic aspect, ceramic masses are being created and applied on the surface of the metallic cap instead of the primer. This has the advantage that it can be used way more easily than the primer. They are burnt two times in a row to fully opacify the metal component. These substances are deposited on a 0.2 mm width layer [4, 7] (Fig. 3). From the masticatory point of view and from the functional occlusion stability point of view, the palatine faces for the frontal group and the occlusal faces for the

Fig. 3 Porcelain fused to metal prostheses (a) CAD-CAM design, (b) metal framework, (c) PFM restorations “in situ”)

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side group that are made out of ceramic mass or from a metallic alloy with morphofunctional individualized pattern that can efficiently participate in the food crushing process and maintain stable dentodental contacts. The ceramic mass, because of the glazed layer, is more resistant to wear than the ones without the layer, but the resistance at compression is the same. From the clinical point of view, to obtain multiple simultaneous and stable contacts, the required grinding can’t be made after the cementing. Areas with no glazing, after a period of use, favor the appearance of wear facets participating in this way to the maxillary selfbalancing [21]. From the biological point of view, it is not possible for a statement to be made to exhibit preference for this type of crown instead of the metal-acrylic one, because for both types of crowns, it is necessary to reduce the vestibular side to create the necessary space for the metallic and nonmetallic (aesthetic) component. Metallic component wall thickness is 0.3–0.5 mm. The thickness of the ceramic mass burned in the vestibular face is 1 mm. At the edges of the incision and occlusal sides, the layer thickness is larger than 1.5–2 mm. This space is obtained by grinding the natural tooth crown tissues. The ceramic masses’ biological tolerability is far superior to the acrylic resins. Ceramics are well supported by the tissue, being inert [22]. Inflammatory reactions of the marginal periodontal that can be observed after the cementation of the metal-ceramic restoration are determined by the existence of a cervical margin of an oversized mixed crown. The ceramic mass layer and the cap wall overpass the dental abutment threshold, compressing the marginal periodontium including interdental papilla. The edge oversizing has two causes: tooth abutment preparation with an insufficiently sized threshold or the desire to obtain a particular aesthetic by applying a thick ceramic mass. Indications and contraindications of the mixed metal-ceramic crown are required to be discussed in terms of what characteristics define the qualities, adding the comparative analysis with two other microprostheses that the specialists possess in the treatment options and may prefer in a given clinical situation, represented by mixed metal-acrylic crown and replacement crown. Metal-ceramic crown indications are ideal aggregation element in terms of physiognomy, biology, masticatory, and functional occlusion of dental bridges in all types of edentulism; restores coronary morphology of anterior and lateral teeth showing the shape and volume changes due to the lack of substance decay products, dysplasia, fractures, and abrasions; restores the color appearance of devital teeth; morphologically restoring a tooth group to provide functional occlusion conditions [7]. The mixed metal-ceramic crown contraindications are represented by the natural tooth crown morphology. All low-volume crowns particularly determined by the cervical incisal or occlusal very low size (primary or secondary) cannot get a mixed crown cover as the side surfaces are not effective to achieve the microprostheses retention on the dental abutment. Casted metal-ceramic prostheses are very popular. Due to the metal frame being really resistant, it makes possible the completion of some long-term fixed dental restorations. They can also be used in difficult cases where the all-ceramic restorations can't be used due to the existence of really large tensions. Feldspathic porcelain

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is used for fixation on metal. The ceramic composition used for metal-ceramic restorations is different than the composition of the porcelain used for full-ceramic restorations, being richer in potassium oxide and sodium oxide. High content of alkaline was necessary to increase the coefficient of thermal expansion. Unfortunately, this increased the tendency of the ceramic material to devitrify and gave it a cloudy aspect [4]. The ceramic masses made for the use of being burnt on metallic alloys have in their composition 10 % potassium oxides, 15 % aluminum oxides, and 55 % silicon oxides, to which sodium and calcium oxides are added. For some special ceramic materials, titanium and cesium are added. The main physicochemical characteristics of ceramic masses burned on metal alloys are summarized below. Usually, these materials are impenetrable to the oral environment, property conferred by the inorganic compact structure. Technologically speaking, it is possible to obtain the required colors after sintering, due to the selection of the material from the vials with the help of the color key. These ceramic masses have good plasticity, and so after the preparation of the paste, a mass will result that can be put on and molded into those surfaces. Due to evaporation of the slurry preparation liquid during the firing they have a low contraction coefficient. The fact that it does not change its volume, which would favor the production of cracks, is really important, and so it has thermal stability at temperature variations and are unfavorable in the firing cycle. The links between the metallic material and the ceramic material can be the chemical bond type at the porcelain-metal interface or the mechanical adhesion of porcelain to metal type. The chemical bonding is considered to be the main linking mechanism. An adherent oxide layer is essential for a good bonding. In the case of base alloys, the chromium oxide is responsible for the bond, but in the case of noble alloys this role is played by the tin alloy, indium oxide, and possibly iridium oxide. Inappropriate oxide formation and oxide excess may lead to poor bond formation and result in the delamination of the porcelain layer. The main advantages are better resistance to tearing due to the metallic reinforcement and better marginal fixing because of the metal frame. Significant disadvantages involve lower aesthetics in comparison to total-ceramic restorations, because the metal layer and the opaque reduce the general translucency of the tooth, and also the fact that this metallic frame is sometimes visible at the edge of the tooth, resulting in dark edges. Regardless of the type of prosthesis, ceramic-ceramic or metal-ceramic, it needs to be mounted inside oral environment [22–26]. The final cementation is done for longer periods of time expressed in more or less years. The final cementation is realized if a cement that was manufactured and marketed for this purpose is used.

Dental Prostheses Failure Analysis Dental restoration materials are subjected to a hostile environment, in which the PH, the saliva quantities, and the mechanical charge are fluctuating constantly and often quickly. These challenges have resulted in substantial research and development in

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Fig. 4 Deposits of biological material on the surface of a dental prosthesis-emergence of degradation phenomena

order to provide the dentists with functional products. These things are possible due to the application of some fundamental concepts of material science. Understanding the polymers, ceramics, and metal properties is crucial in order to select and achieve the dental restoration plan. No material property can define its quality by itself. To describe the quality of the material, usually more properties taken from standardized laboratory and clinical tests need to be appreciated. Standardization of laboratory tests is essential to quality control and to allow the comparison of results between different researchers [27–29]. While it is important to know the comparative values of restorative material properties, it is also essential to analyze the hard and soft tissues that support the prosthetic restoration (Fig. 4). Many dental restorations fail clinically due to fracture or deformation. This is an issue of material property. Also, some restorations flawlessly realized become useless after the failure of the supportive tissue. This is an interface or substrate failure. Therefore, when designing a restoration and when interpreting laboratory tests, it is important to note that the success of dental restorations depends not only on the physical and mechanical properties of the used materials but also on the biophysical and physiological features of the support tissues. According to a study on the causes of failure of dental prosthetics, 30 % are due to oral diseases and 70 % for mechanical reasons [23]. And so, the failure of dental materials can be the result of different factors such as biological (saliva, plaque deposition), mechanical (wear, fractures), technological (voids, inclusions, incorrect overlaps), or the patient (occlusive tension size, pH value in the mouth). Also, in the mouth, restorative materials are exposed to changes in the chemical, thermal, and mechanical nature. These may cause deformation of the material. The performance of the material in the oral environment depends on concepts such as elastic, plastic, and viscoelastic deformations and mechanical quantities such as force, stress, strength, hardness, and toughness [23, 30] (Fig. 5). Occlusal forces that are formed between the teeth of adults are higher in the posterior region and descend from molars to incisors. The maximum occlusive forces range from 200 to 3500 N. Forces on the first and second molars range from 400 to 800 N, and the average load on the bicuspids, cuspids, and incisors is

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Fig. 5 Peeled tartar deposition on the surface of fully ceramic dentures

Fig. 6 Metal-ceramic and all-ceramic prosthesis degraded by abrasion in the mouth

300, 200, and 150 N. In the case of dentures, occlusal forces are generally smaller than the natural dentition forces. For patients with movable partial dentures generate occlusive forces in the range of 65–235 N. For patients with complete artificial teeth, the molars and bicuspids medium force is about 100 N and for incisors 40 N. Variations in the age and sex of population contributes to the large variation in forces. In general, the occlusive force produced by women is 90 N less than that those produced by men. The shape and the define degree of facial muscles are also factors that determine the magnitude of the occlusive force [7, 31]. The maximum occlusal force and the response of surrounding tissue change with anatomical location, age, with the occlusal scheme and placement of dental implants. When designing a restoration and when the materials are selected, it is important to take into account the location, the opposing teeth, and the patient's ability to generate force. These factors can often be estimated by the success or failure of other restorations in the mouth. A material or a model of restoration may be suitable for occlusive forces of the anterior segment but not strong enough for posterior segment [4, 7] (Fig. 6). Tensions are crucial. When a constraint force is acting on a body, this resists to the initial force. This internal reaction is equal in size and opposite in direction with the

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Fig. 7 The failure of prostheses due to ceramic layer mechanical exfoliation (left) and a fullceramic restoration cracking following tensions and forces applied (right)

applied force and it's called stress, denoted by σ. As the tension in a structure varies proportionally with the force and inversely proportionally with the area, that area on which force is applied is very important to consider. This is even more important for dental restorations, where the areas over which large forces are acting are often very small [4, 32]. For example, at the corner of a tooth, the contact surface may have a section area of only 0.15–0.015 cm2 (Fig. 7). The stress is always normalized on an area of 1 m2, but dental restoration in the form of a small occlusive cavity may have a surface area of up to 4 mm2, if the side of the restoration is 2 mm. If an occlusive force of over 400 N is concentrated in this area, the developed stress will be around 100 MPa. So, tensions equivalent to several hundreds of MPa are often encountered in dental restorations. Therefore, when an occlusion is balanced, multiple simultaneous occlusive connections are required. Distribution of occlusal forces on larger surfaces reduces local occlusive stress. Very rarely do forces and stresses happen to be isolated on a single axis. The forces applied individually may be defined as axial, of shear, of torsion, and of bending [4, 33, 34]. Each type of tension is capable of producing a deformation corresponding to that body. Deformation from a stretching tension is an elongation of the material, while the one from the compressive tensions is a compression of the material. The deformation, ε, is described as being the change in length (ΔL = LL0) on the original length (L0) of the body that is subjected to a mechanical load. Deformation is often reported as a percentage. This will be different depending on the type of material and the applied force magnitude. Deformation is an important factor in the analysis of dental restorative materials, such as orthodontic wires or implant screws, which deforms extensively before failure. The wires can be bent and adjusted without fracture [7]. Determining the yield limit of the dental material is important for several reasons. Any dental restoration that is permanently deformed due to masticatory forces may be considered to a certain extent a functional failure. For example, a fixed partial

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denture, which is permanently deformed by excessive occlusive forces, will present altered occlusal contacts. Restoration is permanently deformed because a tension equal to or greater than the yield limit was generated. It is important to note that a distorted restoration is possibly subjected to higher stresses than the original because of the occlusion, which was previously spread over a larger number of contacts, it is now carried out on a much smaller number of contacts. Under these conditions, fracture does not occur if the material is capable of plastic deformation. And so the permanent deformation of the restoration resulted, which is an example of destructive deformation [11]. Permanent deformation of dental materials and the application of tensions higher than the elasticity limit are desired when seeking the shaping of an orthodontic arch wire or when you adjust the bracket of a removable partial denture. In these examples, stresses need to exceed the flow limit in order to bend permanently the element. Elastic deformation occurs when the wire or fastener attaches and detaches from the cervical area of the tooth. The retention is achieved by elastic deformation on a smaller scale. The tension at which a brittle material is fractured is called ultimate strength. This is very important for ceramic materials. It should be noted that the material does not necessarily break at the point in which the largest stress occurs. After a maximum stretching force is applied to the ductile materials, the sample begins to elongate excessively, which leads to the bottleneck phenomenon, or the considerable decrease in the sectional area. However, this does not occur with brittle materials. The tension calculated with the force and cross-sectional area may decrease prior to the final fracture. In specific cases of many dental alloys and ceramics subjected to stresses, ultimate strength and braking strength values are similar [35]. The elasticity of a material is described by the term tensile modulus, or Young's modulus. Other materials such as elastomers and other polymers have low elastic module values, while different metals and ceramics have higher values. The modulus of elasticity is the stiffness of a material in the elastic interval. It can be determined from a stress–strain curve with the equation: modulus of elasticity (E) = Stress (σ)/ strain (ε). For ceramics, this curve has a specific form. Strength is the maximum stress a material can withstand before its failure. Depending on the applied tension, the resistance can be compressive or tensile. Basic strength of an alloy used in dentistry specifies both maximum supported load and minimum sectional area when developing a restoration. An alloy that was loaded up to the breaking limit will remain permanently deformed, so a restoration receiving that amount of tension during operation becomes unnecessary. An error range needs to be incorporated into restoration achievement plan and the material selection to ensure that the breaking limit is not reached in the normal operation of the device. Yield point is often of greater importance than the breaking point for the selection of the material as it is an approximation of the moment in which the material will begin to deform permanently. Most dental materials have those values clearly determined and recorded in the product data [4, 36]. The resilience of a material is the resistance to breaking and indicates the energy required to cause a fracture. The area under the elastic and plastic portions of the

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curve of stress/strain is the toughness of the material. In the case of ceramics for dental use, the resilience may be superior to that of dentin but usually is less than or close to the enamel’s value. Tenacity is the energy necessary to strain the material to the point where it fractures. Brittle materials tend to have low toughness because small plastic deformations occur before failure, so the area near the elastic and plastic zones of the curve is not that much different from the area just beside the elastic zone. From this perspective, dental ceramics have low values of toughness [4]. Recently, the concept of fracture mechanism has been applied to a number of problems in dental materials. The fracture mechanism characterizes the behavior of materials with cracks and other defects and can usually be identified by means of microscopy (stereo or scanning electron). Defects and cracks may occur naturally in a material after it has been in use for a while. Any defect weakens the material, and as a result, fracture occurs at lower stress than the flow limit. Catastrophic fractures often occur in the case of brittle materials, which don’t have the ability to be plastically deformed and to redistribute tensions. The field of breaking mechanisms analyzes the behavior of materials during this type of failures. For brittle materials such as glass, the absence of local plastic deformation is associated with fractures, while for a ductile material, a plastic deformation such as the ability to bend happens with no fracture [37–39]. The ability of a material to be plastically deformed without breaking is called fracture toughness. Generally, when a defect is larger, the stress required to cause fracture is smaller. This thing occurs because tensions, which are normally supported by the material, now concentrate on the starting point of the defect. Fracture toughness was measured for various important restorative materials such as acrylic ones for the denture base, composite, corrective ceramic braces, cements, alloys, natural enamel, and dentin. Typically, the binder addition in polymers can significantly increase the fracture toughness. Strengthening mechanisms are believed to be binder-matrix interactions, but this is still not clear. Similarly, the addition of up to 50 % of the weight of zirconia in ceramics increases fracture toughness [40].

Causes of Failure for Ceramic Dental Materials Alumina (aluminum oxide) is the only solid oxide of aluminum (Al2O3). Alumina was firstly used in the 1970s, but clinical applications from this period showed a fracture rate of 13 %. Failure in this first generation of ceramic is due to the fact that they could not be processed to a high final density. A second improved generation presented a higher density and smaller grains. Fracture rates associated with the second generation decreased with 5 %. Today, a third generation of ceramic components is available, characterized by high purity, maximum density, and fine microstructure [15]. However, despite the best properties and the potential possibility of use as a structural material it has been greatly limited by typical fracture strength of ceramic material. The cracks rapidly propagate in the ceramic, and so they unexpectedly fail during use, and in many cases catastrophically, even when the impact force is below

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the resistance of the ceramic material. The resistance to crack propagation is the ability to activate a mechanism of increasing resistance, such as the deflection [16, 17]. A current study indicates that certain ceramic crowns have a good wear resistance, similar to metal-ceramic crowns. The role of ceramic surface treatments that may be responsible for changing the wear rate is still determined by clinical trials. Tooth structure wear is a natural process that cannot be avoided and occurs when two teeth or a tooth and a dental restoration are in contact and they slide. This natural process can be accelerated by the introduction of a restoration whose wear properties differ from those of the tooth structure of which it is in contact. It can be shown that enamel can be loaded to high levels of wear when in contact with a ceramic material. So far, restoration materials with a wear behavior similar to natural enamel is aimed, because excessive wear can lead to serious clinical problems such as loss of vertical dimension of occlusion, reduced masticatory function associated with temporomandibular joint remodeling, dentin hypersensitivity, etc. [41]. In the mouth there are many factors that contribute to the wear of enamel and dentine, such as the nature of occlusal contacts with opposing teeth (attrition), chewing, tooth brushing, dust breathing (abrasion), acid attack due to consumption of certain fruits or drinks, inhaling industrial acid, or gastric regurgitation when suffering of bulimia and anorexia (corrosion). To observe and quantify the wear it is necessary to know the teeth wear mechanism and how it can be measured and evaluated, both clinically and in the laboratory. The attrition, abrasion, and corrosion terms are often used to identify the various mechanisms that lead to wear or dental restoration failure. Tooth-to-tooth contact causes a wear type called attrition, this taking place without the presence of food or other foreign substances during swallowing and clenching; it is typically characterized by the sides of tooth on an opposite tooth. It gets worse during bruxism [7]. Abrasion is the wear caused by friction between a tooth and an exogenous element. Masticatory abrasion occurs usually when friction is given by the food presence, while abrasion is the result of harmful habits such as nails, pencils, and other hard object biting, opening hair clips with the teeth, etc. Occupational abrasion can occur when repeating a harmful gesture is related to the individual restoration (such as musicians). Ceramic has a high hardness, which can be defined as nonbiological. Even normal cleaning process may cause abrasion of the used tooth restoration material, over time. In developed countries, the main factor that leads to abrasion is the toothpaste that affects to a greater extent the enamel dentine. Brushing without toothpaste has absolutely no effect on the enamel and negligible clinical effects on dentin [42]. The degradation is the loss of material from the tooth surface caused by chemical dissolution. Depending on the acid source that produces dissolution, there are two types of degradation: intrinsic and extrinsic. Tooth degradation (corrosion) caused by intrinsic sources such as bulimia and gastroesophageal reflux gives a translucent and thin enamel surface. In addition, consumption of food and drinks with a pH value of less than 5.5 can cause degradation and demineralization of teeth (Fig. 8).

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Fig. 8 Ceramic prostheses degradation by abrasion phenomena (left-macroscopic analysis and right-detailed images of the same area)

Weakened enamel that is exposed to saliva for an appropriate period can regain minerals, thus increasing its mechanical strength. On the other hand, it was proven that fluoridated toothpaste has a protective effect on enamel corrosion progress. In vivo and in vitro studies indicate that the wear mechanisms rarely act alone, but instead they interact with each other so that teeth wear is the result of three processes: abrasion, corrosion, and attrition [23]. Gaseous inclusion appearance is one of the most important issues that may arise during the obtaining of ceramic bodies. All the ceramic masses that are burned at atmospheric pressure have interwoven structure, with more or less air microcavities. In dental ceramic masses burned in vacuum, most of the air is absorbed from the substance particles, allowing the melted feldspar to flow in the gaps. Gas inclusions in the structure of dental ceramic have negative influence to their physical properties. The bigger the number and the volume of gas inclusions the greatly diminished transparency (Fig. 9). The pores in the ceramic structure create unfavorable conditions for polishing the Jacket crown’s surfaces. Open pores are retentive microspaces that cause debris deposition and promote discoloration. The worst is that gas inclusions decrease the resistance of the walls making it possible to fracture more easily [6]. The main causes of gas inclusions are described below. Powder dental ceramic masses are prepared by mixing it in distilled water or alcohol. The amount of liquid is minimized by vibration and drying (buffering), and after the evaporation of the

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Fig. 9 All-ceramic restoration surface degradation through chemical dissolution phenomena

liquid (drying and firing) gaps remain. They are more and bigger when not built up conscientiously (vibration and buffering). Also, dental-ceramic masses have nonplastic material composition (feldspar and quartz). Organic matter is added after drying in order not to crumble (dextrin, sugar, starch). These ingredients burn without leaving a trace, though creating blanks after burning. An additional reason is due to the fact that during the preparation of the paste and deposition of the ceramic mass, there is the possibility of air incorporation, which remains between particles of different substances. At the same time, the molten silicates incorporate gases at high temperatures, which are not released when decreasing temperature [43]. Another possibility of the pore occurrence is due to the fact that during agglutination, small mineral powder gaps occur at the particles’ boundary. They form a small part of the total pore volume. It is known that in the ceramic mass containing particles of a single size case, keeping as closely as possible is the most effective means of condensation, the space between them representing 45 % of the total volume. If the particles present two dimensions, the space may be reduced to a minimum of 25 %; when three or more dimensions are used then it can be reduced to 22 %. Different powder particle size of ceramic masses favors the approach between them, creating, after burning, more compact and resistant bodies (Figs. 10 and 11). By using a high-vacuum oven, it is possible obtain the removal of gas from the unburned structure. The burning conducted exclusively in vacuum creates a surface with crater-like depressions. The bursting of bubbles close to the surface forms them. It has been proven that only by combining vacuum combustion and atmospheric pressure, a smooth surface can be obtained. *The vacuum is maintained until the end point of the agglutination is done; after that it stops in order to obtain a smooth surface [44, 45]. Although dental restorations must meet strict aesthetic criteria, the most important aspect remains the functional need. Recent research and development of new technologies have led to the introduction of new ceramic materials for dental prosthetics, with its aim of obtaining an optimal combination between strength and aesthetics. Numerous clinical studies have shown that the rate of fracture of

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Fig. 10 Pores – gaseous inclusions in the ceramic layers of coverage media for zirconia-ceramic prostheses

Fig. 11 Microstructural inhomogeneity in all the layers of all-ceramic prostheses

porcelain faces mounted on zirconium core varies between 6 % and 15 % over a period of 3–5 years, while it is only 4 % for conventional metal-ceramic restorations. The types of fractures that can occur with these types of dental restorations are • Cohesive: the fracture is located in the internal structure of the porcelain coating (cutting). • Adhesive: porcelain – zirconium interface. • Complete: complete rupture of the crown.

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As a result of laboratory research, it can be said that porcelain restoration on zirconia core has a higher proportion of cohesive fracture (72 %), as compared with the metal-ceramic ones (8 %); However, metal-ceramic restorations show a much higher rate of adhesive breakage (92 %), as compared with porcelain-zirconium (25 %). Both crowns are susceptible to deformation of the outer layer in the occlusal zone, with fracture projecting from the central point of the force application area to the periphery, where the porcelain delamination occurs. To date, few studies are made on the resistance of dental restorations with zirconia core covered with porcelain in comparison with the abundance of information on metal-ceramic crowns. The zirconia-based dental crown fractures are found most frequently within the porcelain layer, which are more than in the case of metal-ceramics, the top product failure being the chipping. Until this moment, there have not been identified any chemical bonds between zirconium and the porcelain layer. The two substances are fixed by mechanical interposition and by the forming of compressive stress resulting from thermal contraction during cooling after sintering. Therefore, we can say that the most common fractures occur in the layer of porcelain and at the zirconia-porcelain interface [46].

Causes of Failure for Metal-Ceramic Dental Materials Factors that may lead to the failure of dental metal-ceramic prosthetic restorations include technical factors, factors owed to the dentist, material properties, direction, size and frequency of the applied force, environmental factors, etc. Because ceramic materials are a very important component of metal-ceramic prostheses, some of the causes of failure of these prostheses are similar to those that can occur in all-ceramic prostheses. Currently, a significant part of the metallic materials used in dentistry is titanium (titanium alloys), Co-Cr alloys, and Ni-Cr alloys. Good physicalmechanical properties of these materials, relatively low Young’s modulus, fatigue strength, and good corrosion and biocompatibility are all very important. Choosing a metallic material type in metal-ceramic technology holds several technical, medical, and economical considerations [47] (Fig. 12). Regardless of the particular alloy and the technology and materials used, inherently, there may be different types of defects that lead to failure of dental prosthetic restoration [4, 11]. The main causes of defect occurrence in these cases are • • • • • • • •

The mass and alloys have not been compatible. Bonding or oxidation have been misused. Cape made with deformable walls. Stress insertion on the blunt. The cement with high viscosity, the insert was obtained under high pressure. Reduced size of the ceramic layer. Adhesion oxides were thick. Sintering was performed with incorrect technical regimes.

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Fig. 12 Identification of wear and delamination phenomena by abrasion of ceramic layer on a metal-ceramic restoration

• Restoration carried out in adverse conditions, many impurities in the working space [11]. Metal-ceramic prostheses are durable, but most often the failure is due to premature fracture of the ceramic coating. It has been proven that there is a direct relationship between the fracture of the porcelain and the durability of fixed partial dentures. Compared to uncoated ceramic prosthesis, the metal-ceramic show, after 10 years, a much higher risk of failure [6]. It is well known that the fracture resistance is severely reduced when the ceramic material is deposited on an oxidized metal substrate. At the same time, the risk of fracture of the ceramic component of the prosthesis increases with the increasing thickness of the oxide layer above a certain limit [11] (Fig. 13). A very important factor of the compatibility of metal-ceramic prosthesis components is the expansion coefficient of metal and ceramic material. The stress concentration at the metal-ceramic interface is due to the major difference between the expansion coefficients of the two components. To reduce the possibility of failure for this reason, the coefficient of thermal expansion of the ceramic material should be slightly smaller than that of the metal parts of the dental prosthesis [11]. It is also known that the mechanical strength of the prosthesis decreases with the increase of the ceramic layer thickness. The ceramic component at the interface with the metal substrate is usually tensioned, as the metal shrinks more than the ceramic. In this sense, to minimize the possibility of fracturing the ceramic layer

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Fig. 13 The appearance of cracks in the ceramic layer longitudinally in the opaque layer of metalceramic restorations

Fig. 14 Metal mass casting pore defect type

deposited on the metal, it is recommended to obtain a uniform thickness of the ceramic deposit (Fig. 14). The identified defects (pores and cracks) in the ceramic masses deposited on metal frames appear due to technical problems during the application of ceramic materials. Cracks may occur due to the incomplete densification that induces the appearance of angular residual pores. This type of failure may occur, however, when the metal substrate doesn’t have a modulus of elasticity big enough to withstand mechanical loads of the dentomasticatory device. An important aspect that should be followed during the obtaining of metal-ceramic restoration, in order to avoid cracks, is the heating and cooling rates. From this perspective, repeated thermal cycling and overheating the sintering furnace can induce the appearance of superficial or deep imperfections in the ceramic material [7, 11]. Poor clinical interventions are always followed by other problems in the lab. When probing, the axial and transversal inadequacy in the cervical zone may be noted . When one finds axial inadequacy, the short cape requires rebuilding after a new mask or on the same model if the defect appears on the model. Transversally inappropriate cape (large) is recovering after another masking and model manufacturing. The long caps are retouched step by step. At the same time, the

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Fig. 15 Pore type defect in the opaque layer and first ceramic layer of metal-ceramic restorations

occurrence of some casting defects represented by pores is possible. Defects in the form of pores determine the restoration of the cape. Isolated pluses that are well contoured are polished [6]. Additionally, the sizes and cross-sectional outlines of all components have important effects on the stability and strength of the entire metalceramic prosthesis structure [4]. The connector must be thick enough to support the occlusal loads, but gingival and occlusal embrasures must be obtained so as to be aesthetic (Fig. 15). Failure factors correlated to the dentist are different. Within them, the most important are related to the fact that metal anterioposterior substructures are curving when loaded complexly or excessively, which leads to the fracture of the ceramic coating of the prosthesis. Also, a major risk of failure presents the prosthetic replacement restorations of three posterior teeth. For the risk avoidance associated with uneven distribution of forces and possible mechanical failures in these cases it is recommended to use implant supported prosthetics or partially mobilized teeth [11]. One factor that can lead to failure of the prosthesis because of the dentist is the inappropriate preparation of the teeth (resulting interocclusal space insufficient for the metal structure and overlapping the ceramic deposits). Therefore, cracks may occur during the technological process or shortly after starting. That’s why the experience of the dentist is very important in order to obtain a correct design, this inherently leading to achieving a long-lasting prosthetic restoration (Figs. 16 and 17). Detailed analysis of denture fractures is rare in the literature and dental biomechanics. In most cases, the fracture surface of the prosthesis, which is essential for fracture analysis, is destroyed or severely damaged over the point where

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Fig. 16 Opaque layer peeling from the metal substrate and the appearance of pore type defects in the ceramic coating

Fig. 17 Fully fractured outer ceramic layer and opaque coverage; identification of pores in the ceramic layer

fractographic analysis can be performed [6]. In addition to the incomplete oxidation, the presence of fractures and cracks that cause separation is determined by the presence of impurities, by the fact that the ceramic slurry drying wasn’t made progressively, by the physical and chemical incompatibility between the alloy and the ceramic mass, by the metal’s unsuitable size, by the ceramic layer’s nonuniform size, by sudden changes in the thermal values, or by the forced insertion of the model abutment or dental abutment if it has flaws [46–49]. Since the 1950s it was proved that water could act chemically on the crack edges, reducing the strength of ceramics. This phenomenon is called crack propagation chemically assisted or static fatigue. Inside wet environments the metal-ceramic bond strength is reduced by about one-fourth, which leads to crack propagation along the microcracks and the final prosthetic restoration failure. In addition, low pH beverages induce cracks generating in the vitreous materials for dental restorations [11]. Detaching the ceramic from the metal parts after cementing is the most serious fault that is remedied only after removal of the coping, generally by sectioning. Intervention is difficult, accompanied by unfavorable repercussions for that dental cabinet. After removal shall also mark the prosthetic field to resume all technological stages [6] (Figs. 18 and 19).

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Fig. 18 Degradation of metal-ceramic restorations by partially peeling the layer ceramic

Fig. 19 Fracture of metal-ceramic restorations mounted on a Ti implant

Fig. 20 Abrasion on the surface of metal-ceramic restorations and identification of pores in the outer ceramic layer

Implant-supported prostheses are susceptible to failure compared to fixed natural teeth. In these cases, failure is generally due to fracture of ceramics (especially in patients with bruxism habits) [38]. The most likely explanation is that in these cases the natural tooth (along with periodontal) provides a sensitive detection of the occlusal loads, while implants are not receiving this mechanism and cannot absorb shocks. These effects can be limited by the use of nonrigid connectors for decreasing the forces applied to the superstructure [41] (Fig. 20).

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If the fracture surface shows ductile fracture and fatigue ribs, the main cause of fracture is most likely the presence of transgranular stress that determines corrosion and cracking. Corrosion can be the main factor that has caused the failure of the prosthesis, in which case one searches for products of corrosion or corrosion points. Also, the deposition of calcium and phosphorus on the surface of the prosthesis may be an integral part of the mechanism that leads to failure [11].

Conclusions The result of the studies and research carried out in this work confirms the theoretical data on metal-ceramic dental restorations and full-ceramic single and multitooth. Of all dental restorative materials analyzed in this study, the material showing the best features is the yttrium-doped zirconia ceramics type. In this case major defects that could lead to failure dentures were not found. However, a major disadvantage of this type of dental restoration is the high cost of both the material itself and the processing technology and labor in the dental cabinets. In the case of metal-ceramic prostheses, there are many factors that can lead to failure, especially because of the ceramic fracture. These factors may be associated to technician, dentist, environment, prosthesis design, or inherent metal or ceramic material structure and microstructure. Regardless the type of the used prosthesis, possible factors leading to the failure of the dental restoration prosthesis are abrasion, peeling, delamination, the presence of material defects that become sources of crack propagation, occlusive stress, etc. Common types of ceramic like porcelain and alumina ceramics have many internal defects such as pores, voids, and gas inclusions. Also their surface is susceptible to flaking, abrasion, and corrosion. Although all-ceramic restorations with superior mechanical properties are an ideal of the aesthetics, their strength is far below the metal-ceramic. Dental metal-ceramic prostheses are more accessible in terms of manufacturing technology and cost of materials, but their main disadvantage is aesthetically (often the metal component may change the prosthesis color or become visible in contact with the gum tissue). There is also the risk of adhesive cracking at metal-ceramic interface, the risk that does not appear in all-ceramic restorations. To eliminate the main causes of failure, highest strength characteristics of the materials must be exploited, and the weakest should be removed. Making the future of all-ceramic restorations with mechanical properties better than metalceramic ones is a priority in modern dentistry. Acknowledgments This work was supported by a grant of the Romanian National Authority for Scientific Research and Innovation, CNCS – UEFISCDI, project number PN-II-RU-TE-2014-40590.

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Part VI Clinical Performance in Bioresorbable and Load-Bearing Applications: Other Applications

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Bioceramics and Composites for Orbital Implants: Current Trends and Clinical Performance Francesco Baino

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Essential Medical Background . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Historical Overview: A Century of Evolution (From the Late Nineteenth Century to 1980s) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Currently Used Ceramic Porous Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . HA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Alumina . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Glasses and Glass Ceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Composites and Coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . HA/Silicone Composite Implant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioactive Glass/PE Composite Porous Implant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . HA-Coated Alumina Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . HA Implants Coated with Mesoporous Bioactive Glass . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary and Outlook . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

The applications of ceramics in the medical field are mainly related to the repair of hard tissues, like the bone and teeth, whereas their potential for the repair of other tissues has been often underestimated. A clinical area where porous ceramics are playing a key role is anophthalmic surgery, which deals with the removal of diseased eyes and their substitution by a (usually) spherical orbital implant replacing the ocular volume. Over the years, many bioceramics have been proposed for such an application, including glass, hydroxyapatite, and alumina.

F. Baino (*) Institute of Materials Physics and Engineering, Applied Science and Technology Department, Politecnico di Torino, Torino, Italy e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_60

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Recently, polymer/bioceramic composites have been also experimented in order to reduce the stiffness mismatch between the ceramic constituent of the orbital implant and the surrounding soft tissues. This chapter provides an overview of the currently used ceramic-based orbital implants, also exploring new research directions and highlighting the promises for the future disclosed by the recent advances in bioceramics science. Keywords

Hydroxyapatite • Alumina • Bioactive glass • Ceramic/polymer composites • Porosity • Fibrovascularization • Tissue ingrowth • Angiogenesis • Antibacterial properties • Eye orbit • Enucleation • Ocular surgery

Introduction The removal of an eye (enucleation) is one of the most serious and difficult decisions that a surgeon must consider in case of severe ocular trauma or potentially lifethreatening diseases to the patient [1]. This approach, however, leaves a partially or totally empty socket and creates significant cosmetic issues for the patient. Therefore, an orbital implant is then permanently placed in the orbit in order to restore the volume previously occupied by the ocular globe as well as to ensure adequate support to surrounding tissues (Fig. 1). Aesthetic outcome has been improved by

Fig. 1 Placement of an orbital implant in the anophthalmic socket following enucleation surgery: 1 orbital implant (a porous sphere is schematically depicted; wrapping of the ceramic implant within a sheet of soft and smooth material, such as a polymer, is recommended to facilitate its insertion into the anophthalmic socket and to avoid erosion of surrounding tissues by the outer irregular surface); 2 frontal peg (it is optional, and extra-surgery is needed for its placement after at least 6 weeks from primary surgery); 3 patient’s conjunctiva; 4 extraocular muscles attached to the implant; 5 orbital bone; 6 preserved eyelids; 7 ocular prosthesis; 8 seat to host the implant peg, in order to transmit a life-like motility to the artificial eye

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the development of ocular prostheses, also known as “artificial eyes,” that are typically polymeric inserts which the patient places behind the eyelids and in front of the conjunctiva. The outer surface of the ocular prosthesis is painted to almost exactly match the appearance of the contralateral eye; it can be temporarily removed for cleaning and when the patient goes to sleep. If the extraocular muscles are attached to the orbital implant, some movement may be consequently transmitted to the overlying ocular prosthesis, which is a highly desirable characteristic known as motility. Over the years, orbital implant design evolved from simple solid sphere (glass, silicone, poly(methyl methacrylate) (PMMA)) to more complex devices having specific sites for extraocular muscle attachment (the class of the so-called Allen implants made of PMMA) to porous spheres (hydroxyapatite (HA), polyethylene (PE), alumina) allowing better biointegration [2–4]. In this regard, bioceramics has been mainly employed to produce porous orbital implants, alone (as single-phase materials) or in the form of composites [5]. These implants incite minimal host response, characterized by the formation of a pseudocapsule around the implanted material. However, the success of these devices was hampered by the development of long-term complications, such as late migration and extrusion. An ideal orbital implant should display a number of characteristics, including biocompatibility (which should be a sort of precondition), adequate socket volume replacement, good motility transmitted to the ocular prosthesis, adequate support for the ocular prosthesis, low cost, easiness of implantation, non-degradability, and overall very low rate of complications. Many evidences seem to demonstrate that porous orbital implants lead to better performances and outcomes with respect to the other existing solutions. These implants are typically characterized by a 3-D network of interconnected pores, resembling – at least to some extent – the trabecular structure of human cancellous bone [6]. Fibrovascularization, defined as the ingrowth of viable vascular connective tissue, typically occurs within 4 weeks from implantation, and tissue reaction is often minimal (Fig. 2). Wrapping within a sheet of soft, smooth material of biological or synthetic origin (e.g., autologous sclera, polymeric mesh) at the time of surgery can be useful to reduce the risk of abrasion of the conjunctiva – and associated implant exposure – due to the irregular surface of ceramic porous implants. From a general viewpoint, the growth of tissue into a porous orbital implant offers several key advantages that can be summarized as follows [2, 4]: • Extrusion is generally very rare. • With sufficient time for fibrovascularization to occur, any exposures (due to the abrasion of the conjunctiva by the implant) might be able to heal spontaneously due to the good blood supply within the implant. • The presence of adequate blood supply within the implant can significantly reduce the risk of implant infection. • Porous implants can be pegged, thereby improving the motility and the life-like appearance of the ocular prosthesis that can more closely follow the movement of the contralateral eye.

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Fig. 2 Reconstruction by cone beam computed tomography (CBCT) of an orbital implant (synthetic HA porous sphere) implanted in the right orbit of a human patient where the red zones represent the areas of fibrovascularization that starts at the implant periphery; the vitreous body of the contralateral eye is represented as a gray sphere (Courtesy of Lukats et al. [6])

The last advantage deserves a short discussion for better understanding. Looking at the historical evolution of orbital implants, previous attempts to establish a direct mechanical connection between nonporous orbital implants (e.g., a simple polymeric sphere) and the ocular prosthesis had been invariably met with the development of infections (due to bacterial colonization of the implant region to be connected with the prosthesis), exposure, and other complications. Thanks to the fibrovascularization of porous implants, it is possible to drill a hole in the implant and to insert a peg between implant and ocular prosthesis (Fig. 1). The presence of a good blood supply allows the drilled area to re-epithelialize with conjunctiva; thus, the implant remains separated from the external environment by the living conjunctival layer but retains a direct connection to the prosthesis, thereby providing enhanced motility without exposure of the implant and the associated complications. This chapter provides an overview of ceramic-based orbital implants, including both the devices currently adopted in the clinical practice and those fallen in disuse; some indications for prospective research and future challenges are also given in the light of the recent advances in bioactive ceramic science. Table 1 provides a short glossary of the medical terms that are not explained directly in the text and which may be unclear or unknown to nonspecialist readers.

Essential Medical Background A short overview of the reasons why the removal of an eye can be necessary, as well as the related surgical procedures, is given in this section for the reader’s benefit. At present, the removal of a diseased eye can be carried out by one of three different

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Table 1 Medical glossary (terms listed alphabetically) Term Conjunctiva

Ectropion Endophthalmitis

Enophthalmos

Exposure

Extraocular muscles Migration Pegging

Pyogenic granuloma

Retinoblastoma

Sclera

Explanation Clear mucous membrane constituted by stratified columnar epithelium that covers the sclera and lines the inside of the eyelids. It contributes to eye lubrication by producing mucus and tears, although in a smaller amount with respect to lachrymal glands. In addition it prevents the entrance of pathogenic agents and foreign bodies into the eye Turning out of the eyelid (usually the lower eyelid) so that its inner surface is exposed Inflammatory condition of the intraocular cavities containing the aqueous/ vitreous humor usually caused by infection. Panophthalmitis is the inflammation of all coats of the eye including intraocular structures; endogenous endophthalmitis results from the hematogenous spread of organisms from a distant source of infection; exogenous endophthalmitis are due to direct inoculation of bacteria/fungi from the outside as a complication of ocular surgery, foreign bodies, or penetrating ocular trauma Recession of the eyeball or orbital implant inserted in the anophthalmic socket within the orbit; it may be a congenital anomaly or be acquired as a result of trauma (e.g., blowout fracture of the orbit) or else be related to postoperative complications of oculo-orbital surgery Break in the tissue overlying the orbital implant, which in severe cases may lead to extrusion of the entire implant. Poor surgical technique, excessively large implant size, and infection may all contribute to this postoperative complication Group of six muscles, attaching to the sclera/orbital implant, that control the movements of the eye/implant Change in position of the implant following placement within the anophthalmic socket Surgical procedure that can be optionally performed after some months from orbital implant placement in the anophthalmic socket (primary surgery). In this procedure, a hole is drilled into the front surface of the implant, and a polymeric or metal peg is inserted into this hole. The peg articulates with a cavity in the back surface of the prosthetic eye, thereby providing improved motility. Pegging is usually adopted only for porous implants, as fibrovascularization allows to decrease the risk of infections that might follow the pegging procedure Overgrowth of tissue due to irritation or physical trauma. Its appearance is usually a color ranging from red/pink to purple and can be smooth or lobulated. Younger lesions are more likely to be red because of the high number of blood vessels, whereas older lesions begin to change into a pink color. It can be painful, grow rapidly, and often bleed profusely with little or no trauma Rare type of eye cancer that affects the retina and usually develops in early childhood, typically before the age of 5 (it is typically diagnosed in children aged 1–2 years). Retinoblastoma is due to the mutation of RB1 gene, which can be inherited (in this case the tumor typically develops in both eyes) or occur in the early stages of fetal development Opaque, fibrous, protective, outer layer of the eye. Primarily constituted by collagen, it maintains the shape of the globe, offers resistance to internal and external forces, and provides an attachment for the extraocular muscle insertions. The thickness of the sclera varies from 1 mm at the posterior pole (continued)

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Table 1 (continued) Term

Sympathetic ophthalmia

Tenon’s capsule Wrapping

Explanation to 0.3 mm just behind the rectus muscle insertions. It is commonly referred to as the “white of the eye” Bilateral diffuse granulomatous uveitis (a kind of inflammation) of both eyes following ocular trauma. It is quite rare but can leave the patient completely blind; early symptoms (e.g., pain, photophobia) may develop from days to several years after a penetrating eye injury Sheet of connective tissue that lines the eyeball and aims to provide a smooth socket for allowing the free movement of the ocular globe Preoperative strategy involving the wrapping of an orbital implant within a sheet of a smooth material with the aim of facilitating its placement within the soft tissue of the eye socket, diminishing tissue drag, and helping precise fixation of the rectus muscles to the implant surface. Wrapping is particularly recommended for porous implants in order to provide a physical barrier over their slightly irregular porous surface. Suitable wrapping materials include scleral autografts and allografts, bovine pericardium, and synthetic polymer sheets

surgical approaches, i.e., evisceration, enucleation, and exenteration, according to the particular pathology and the medical history of each patient [1]. Evisceration involves the removal of the intraocular contents of the eye, while the sclera, Tenon’s capsule, conjunctiva, extraocular muscles, and optic nerve are left intact [7]. Enucleation is another option involving the removal of the globe from the orbital socket, together with the scleral envelope and a portion of the optic nerve, while, as with evisceration, the conjunctiva, Tenon’s capsule, and extraocular muscles are spared [1, 7]. In the final stage of surgery, an orbital implant is placed within the scleral envelope after evisceration or within the Tenon’s capsule after enucleation; an ocular prosthesis will be then worn by the patient to restore an appropriate aesthetic appearance, including life-like motility (Fig. 1). Removal of an eye can be necessary in the cases of intraocular malignancy (e.g., retinoblastoma, which can develop especially in children), blind painful eye, prevention of sympathetic ophthalmia in a blind eye, severe trauma, and infections not responsive to pharmaceutical therapy. From a general viewpoint, evisceration is less invasive and less surgically complex than enucleation, but in some cases the complication rate for evisceration, specifically implant extrusion, may be significantly higher [7]. Evisceration is indicated in the treatment of active, uncontrolled endophthalmitis and in all cases when there may be a danger of intraocular infection spreading back along a cut optic nerve sheath; enucleation may be indicated if the infection has spread to the sclera. Evisceration is also recommended in patients who have bleeding disorders since it is a faster, easier procedure and damages fewer blood vessels than enucleation. Evisceration is absolutely contraindicated in the presence of intraocular malignancy as it does not allow for eradication of tumor cells that have spread to the sclera. Enucleation is generally indicated for tumors that are confined to the ocular globe; if the malignancy has spread to the extraocular structures (i.e., the conjunctiva and

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eyelids, adjacent sinuses, cranial bones, face muscles, and skin), a radical procedure of exenteration, which involves the removal of the entire orbit and surrounding tissues, is recommended [8]. In this case, silicone or acrylic custom-made large prosthetic devices are attached to the orbit and the skin with various types of adhesives to provide good cosmetic results [9].

Historical Overview: A Century of Evolution (From the Late Nineteenth Century to 1980s) Since the ancient Roman age, a wide variety of materials, including wool, clay, gold, and silver, has been used to manufacture more or less rudimental orbital fillers, often painted or enameled to mimic the natural iris, with the aim of replacing the anophthalmic socket volume and restoring an acceptable aesthetic appearance to the patient’s face. The first clear description of a production process to make orbital implants dates back to the end of sixteenth century, when Venetian glassmakers began to fabricate prosthetic eyes of blown glass that however were brittle and had poor fit and little comfort [3]. The modern era of anophthalmic surgery started in 1885, when Mules first described in detail the surgical placement of a hollow glass sphere into an eviscerated globe [10]. Since the early 1900s, the use of orbital implants coupled with glass ocular prostheses began to be adopted in order to restore a better aesthetic appearance to the patient’s face; the prosthesis was a glass shell placed between the closed conjunctival surface covering the orbital implant (bulbar conjunctiva) and the eyelids (palpebral conjunctiva) (see also Fig. 2). Glass eyes had to be worn with caution as they were brittle and prone to implosion with acute changes in temperature; furthermore, they became etched from exposure to body secretions. It was not until the First World War that glass eyes were used by the general population; from 1920 to 1940, Germany became the main supplier of glass orbital implants due to the superior glass-blowing techniques as well as the improvements adopted in implant design (thicker glass shell to reduce the risk of fracture, the so-called Snellen implant) [11]. In a study involving 52 patients, Burch estimated failures in less than 10 % of cases using the Snellen implant in the mid 1940s [12]. The battle casualties of the Second World War caused a large demand of artificial eyes, but the wartime shortage of glass eyes imported from Germany led to the development of PMMA orbital devices. PMMA implants allowed to overcome the problems of glass implants (brittleness and chemical etching by body secretion), permitted custom fitting at a relatively low cost, and allowed better motility of the prosthesis due to a new design for improved muscle attachment [13–15]. In the last decades, the use of glass spheres as orbital implants has been almost totally abandoned, and only occasionally they were still implanted in selected cases [16, 17]. As an alternative to brittle glass implants, small spheres of natural ivory from elephant’s tusks, a biocomposite constituted by nano-sized HA-like crystals (about

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70 wt%) and organic matter (a complex network of type I collagen fibers) that is eliminated after the death of the animal, were experimented as orbital implants in the first half of the twentieth century [18]; after the Second World War, no other application of such type of biological apatite has been reported in the ophthalmic field. Animal-derived HA implants were also prepared by heating spheres of bovine cancellous bone to destroy all organic matter, leaving only the calcium phosphate mineral framework [19–21]. Interestingly, these porous HA implants derived from charred bone were widely used before the Second World War and recommended as “the most satisfactory of all orbital implants” [18]. Since the late 1940s, biologically inert nonporous polymeric spheres (silicone and PMMA) progressively displaced these early types of porous implants due to the easier material processing and low cost. However, the concept of a bone-derived HA orbital implant was resurrected by Molteno and coworkers in the 1980s and led to the development of the “M-Sphere,” which is still currently used in the clinical practice (see the section “Bone-derived HA”). In the late 1970s, a ceramic/polymer felt-like composite, called Proplast I, gained a certain popularity as an orbital implant material. Proplast I was constituted of polytetrafluoroethylene (PTFE) and carbon fibers and, when implanted, could be invaded by fibrous tissue, thereby overcoming the problem of extrusion and rejection [22]. Studies in human patients who received Proplast I hemispherical orbital implants were promising, with no cases of migration or extrusion after a 2-year follow-up [23]. In recent years, however, the popularity of Proplast I has declined because of long-term postoperative complications, primarily late infections, associated with its use in other medical applications [24]. One decade later, the implantation of porous enucleation implants made of Proplast II, an evolution of Proplast I, was reported. This new device was different from its predecessor in the composition (Proplast II was an alumina/PTFE composite) and in having a siliconized nonporous posterior surface to allow smoother movements, together with a porous anterior portion to ideally facilitate fibrovascular ingrowth [25, 26]. Proplast implant II had a nipple on its anterior surface (lined by the patient’s conjunctiva) that could integrate with a depression on the posterior surface of the ocular prosthesis. Several Proplast implants II required subsequent removal because of poor motility and, over histopathological examination, were found to be completely avascular and surrounded by a thick pseudocapsule [27]; therefore, since then the use of Proplast II to fabricate orbital implants was abandoned. However, in the mid-1990s, Proplast II was successfully adopted to manufacture subperiosteal implants for the correction of anophthalmic enophthalmos in 34 patients having poor orbital volume replacement despite the prior insertion of an adequately sized spherical implant within the orbital socket [28]. At present, all the ceramic-based orbital implants mentioned above are fallen in disuse; they are also collected in Table 2. Maybe their use might be reconsidered in the future, in the light of new advances of biomedical materials research.

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Table 2 Currently abandoned ceramic-based orbital implants Material Glass (noncrystalline ceramic)

Type of implant Hollow sphere (blown glass)

Recipient Human

Tusk-derived HA (ivory) HA derived from heattreated bovine bone

Solid sphere

Human

Porous sphere of charred cancellous bone

Human

Carbon/PTFE composite (Proplast I)

Hemispherical implants

Human

Alumina/PTFE composite (Proplast II)

Porous implant having a siliconized nonporous posterior surface to allow smoother movements

Human

Remarks It was the standard orbital implant from the late 1800s to the Second World War. In recent years it has been almost totally abandoned It was used till the Second World War It was used till the Second World War and considered an excellent alternative to glass orbital implants Despite the fibrovascular ingrowth and generally good outcomes, it was abandoned in the 1980s due to the risk of late infections It was abandoned due to poor motility and absence of fibrovascular ingrowth

References [10–12, 16, 17]

[18] [18–21]

[22–24]

[25–28]

Currently Used Ceramic Porous Implants This section mainly focuses on HA-derived and alumina orbital implants that are commonly used in the current clinical practice; their main characteristics are summarized in Tables 3 and 4. Furthermore, the results of recent studies involving the experimental use of bioactive glass orbital implants in humans are also mentioned.

HA HA formally belongs to the class of calcium orthophosphates and, especially in the form of coralline and synthetic HA, has been widely used since more than 50 years in orthopedics and dentistry, thanks to its chemical and compositional similarity to biological apatite of hard tissues [29]. Porous HA is widely used as an orbital implant material due to its biocompatibility, non-absorbability, and as it allows biointegration through fibrovascular ingrowth.

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Table 3 Currently used ceramic porous orbital implants and comparison with porous PE devices Material Bovine bonederived HA

Type of implant Porous sphere

Recipient Human

Coralline HA

Porous sphere, egg-shaped porous implants Porous sphere, egg-shaped porous implants

Human

Alumina

Porous sphere

Human

Porous PE

Porous sphere, egg-shaped and conical porous implants, other variations with gradients of porosity

Human

Synthetic HA

Human

Remarksa Commercial product, M-Sphere; cost, around 500€. Problems of brittleness can occur Commercial product, Bio-Eye ®, cost, around 600 € Mostly used commercial products, FCI3; cost, around 450 €. Other less expensive implants are available (with problems associated with low purity) Commercial product, bioceramic implant; cost, around 450€ Commercial product, Medpor®; cost, around 150€

References [30–35]

[36, 37, 44–54] [38–43]

[56–65]

[58, 76]

a

Prices are indicative only and refer to unwrapped implants; they may vary depending on country, hospital, and number of implants ordered

Table 4 Microstructural characteristics of the ceramic porous orbital implants that are in common use and comparison with porous PE devices Marketed product M-Sphere

Material Bovine bonederived HA Coralline HA

Bio-Eye ®

Synthetic HA

FCI3

Alumina

Bioceramic implant

PE

Medpor ® spherical implants

Porosity (vol.%) Around 80 Around 65 Around 50 Around 75 30–70

Pore size (μm) above 300 300–700 300–500 500

100–1000

Surface structure Ultramicroscopic crystals typical of bone HA Irregular microcrystals of HA with size around 2 μm Hexagonal microcrystals of HA with size within 1–5 μm Cobblestone pattern of microcrystals with size within 0.4–1.1 μm Woven texture; presence of surface irregularities

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Bone-Derived HA In the 1970s, the research group coordinated by Dr. Molteno carefully reviewed the available literature on orbital implants and noted that the postoperative exposures of bone-derived HA spheres used till the Second World War were generally rare, small, and frequently tended to heal spontaneously [18]. This behavior, which was quite unlike that observed with smooth-surfaced polymeric orbital implants, suggested that the biodegradable nanocrystalline HA matrix of the bone would constitute a superior orbital implant since, once organized by host connective tissue, it would not migrate through the tissues, while any small exposures would heal spontaneously. Furthermore, the mass of host connective tissue incorporating the bone mineral implant would likely persist unchanged for the patient’s whole life. The early trials of this type of implant (the so-called M-Sphere) involved the use of deproteinized (antigen free) bone of calf fibulae and confirmed that the mineral matrix of cancellous bone was readily incorporated into the tissues and that small exposures were followed by spontaneous crumbling of the exposed bone with healing of the overlying conjunctiva [30, 31]. Other 52 cases with up to 10-year follow-up were reported in 1991 [32], and the long-term successful outcomes of 120 M-Sphere orbital implants inserted after enucleation between 1977 and 2000 were more recently documented [33]. This implant is significantly more porous (around 80 vol.%) than other available HA orbital implants (50–65 vol.%); the use of a lighter device is an advantage leading to decreased stress on the lower lid and associated ectropion formation. However, due to the high porosity, the M-Sphere orbital implant is fragile and may be unable to support a peg [34, 35]. This drawback associated to its relatively high cost (around 500€, lower than coralline HA but significantly higher than porous PE implants, see Table 3) may have contributed to its limited diffusion. Coralline HA Porous orbital implants spread worldwide since the 1990s after the introduction of modern HA orbital implants that are not based on treated bone deriving from animal sources. Perry first experimentally introduced the coralline porous HA sphere (Bio-Eye ®) in anophthalmic surgery with excellent postoperative outcomes [36]. The interconnected porous structure of the HA implant allowed host fibrovascular ingrowth, which potentially reduces the risk of migration, extrusion, and infection [37]. Apart from discouraging bacterial colonization of implant surface, vascularization also allows the treatment of ocular infection by antibiotic therapy. After some months from primary surgery, the frontal region of the HA implant can be drilled to place a peg that can be subsequently coupled to the posterior surface of the ocular prosthesis (Fig. 1), so that a wide range of artificial eye movements (especially along the horizontal axis) as well as fine darting eye movements (commonly seen during close conversational speech) can be achieved, thereby imparting a more life-like quality to the prosthetic eye. However, the use of coralline porous HA implants is associated to two peculiar drawbacks [2, 4]. The first problem is ecological, as the manufacture of such implants involves damage to marine ecosystems due to the harvesting of natural

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Fig. 3 Synthetic HA orbital implant: (a) appearance of the porous sphere; (b) sphere wrapped in a smooth, soft mesh of polyglactin 910 prior to implantation; (c) microstructure of the HA implant (planar reconstruction by micro-CT) (Courtesy of Lukats et al. [6])

corals; the second issue is related to the high cost (around 600€) compared to other options (e.g., porous PE; see Table 3, or nonporous silicone, and PMMA spheres whose price is around 50€). Primarily in order to reduce the cost of the device, other forms of HA have been proposed as potentially suitable and less expensive materials for orbital implant fabrication.

Synthetic HA Chemically synthesized HA implants (the reference implant is the so-called FCI3 sphere) (Fig. 3) have an identical chemical composition to that of the Bio-Eye ® [38], although scanning electron microscopy (SEM) investigations revealed some architectural differences (lower porosity: 50 vs. 65 vol.%; decreased pore size uniformity; presence of blind pouches and closed pores) [39]. Implant interconnectivity (Fig. 3c)

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is still sufficient to allow central implant fibrovascularization [40]. FCI3 implant has gained increasing popularity over the past 10 years especially as it is significantly less expensive than the Bio-Eye ® (see Table 3) and easier to drill for peg placement. Low-cost versions of synthetic HA orbital implants have been developed and are currently in use especially in some emerging countries; however, they exhibit a number of drawbacks that strongly limit their economic advantage over the other available options. Some of these implants have been reported to contain CaO impurities that, after hydration in host tissues, may form Ca(OH)2, which is caustic [41]. Other implants have higher weight, lower porosity (below 50 vol.%), and lower pore interconnectivity than FCI3 devices, with consequent limited fibrovascularization and enhanced risk of implant migration [42]. A new type of synthetic HA orbital implants (75 vol.% porosity, pore sizes ranging from 100 to 300 μm) was recently experimented by Kundu et al. [43] with apparently good outcomes, but their study is still too limited (25 patients, 2.5 years of follow-up) to draw definite conclusions on their long-term safety and efficacy.

Shortcomings of HA Orbital Implants Among porous orbital implants, coralline and synthetic HA devices are still the preferred choice by the majority of ophthalmic surgeons worldwide. However, despite the relatively good overall biocompatibility profile, HA generally exhibits certain drawbacks for use in orbital implants. Being a porous ceramic, its brittle nature precludes suturing the extraocular muscles directly to the implant [1, 7]; thus, preoperative placement of the HA implant within a sheet of soft material is necessary for muscles attachment. Furthermore, there are convincing evidences that the rough surface of HA implants may contribute to the development of late exposure due to the abrasion of the relatively thin conjunctiva and Tenon’s capsule as the implant moves. Therefore, also for this reason it is generally recommended that HA implants are placed within a sheet of wrapping material (Fig. 3b) before introduction into the orbit [44, 45]. It was also shown that the majority of exposed HA implants can be successfully treated by using patch grafts of different origin (e.g., scleral graft, dermis graft, oral mucosa graft) without the need for implant removal [46–48]. In case of orbital implant infections associated to exposures, administration of systemic antibiotics and topical eye drops can solve the problem, but if no symptoms improvement is noticed, implant removal should be considered [49]. Other reported complications include conjunctival thinning (followed or not by exposure), socket discharge, pyogenic granuloma formation, midterm to chronic infection of the implant, persistent pain or discomfort, and peg extrusion from drilled HA implants [50–54]. In the search for an “ideal” porous orbital implants with a reduced complication profile and diminished surgical and postoperative costs, alternative materials such as porous PE (not treated in the present chapter; the interested reader is addressed to some recently published articles [2, 4]) and alumina have been investigated over the last two decades.

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Alumina Aluminum oxide (Al2O3), commonly termed alumina, has been used for decades in orthopedics, thanks to its attractive mechanical properties (high hardness and compressive strength, excellent resistance to wear), biocompatibility, and bio-inertness. For instance, the introduction of alumina and, later, alumina-based ceramic composites for manufacturing prosthetic femur heads had a significant impact in the field of hip joint replacement, leading to an improvement of prosthesis duration and performance as well as of patient’s life quality [55]. Since the late 1990s, alumina began to be introduced as an alternative to HA for the fabrication of porous orbital implants; this type of device was then approved by Food and Drug Administration in 2000 and is currently marketed under the commercial name of “bioceramic implant.” There are convincing evidences showing that alumina implants are a reliable and even more effective alternative to the available porous HA devices under many viewpoints, including cost (these implants are significantly less expensive than Bio-Eye ®; see Table 3), low tendency to exposure/extrusion (smoother surface compared to HA implants; see Table 4), and overall postoperative performances. The first in vivo study was reported in 1998 by Morel et al. [56], who evaluated the clinical tolerance of porous alumina implants in 16 eviscerated rabbits: only one infection was observed without conjunctival breakdown, and fibrovascular ingrowth started 15 days postoperatively and was complete after 1 month. These promising results were confirmed 2 years later by Jordan et al. [57], who compared the performance of alumina and HA implants in rabbits and highlighted that the new alumina implant was as biocompatible as HA and less expensive, and its manufacturing did not involve any damage to marine life ecosystems as may occur in the harvesting of coral for coralline HA devices. An accurate comparison about the proliferation of orbital fibroblasts in vitro after exposure to bioceramic implant and other three implants made of different materials (coralline HA, synthetic HA, porous PE) was reported by Mawn et al. [58]. The proliferation of fibroblasts differed on the various studied implants and, specifically, was maximum on the bioceramic implant; furthermore, the fibroblasts growing on the Bio-Eye®, synthetic HA, and PE implants all had debris associated with them, whereas the alumina implant was free of these debris, which was mainly attributed to its finely crystalline microstructure (see also Table 4). Promising results were also published in 2002 by Akichica et al. [59], who implanted pieces of alumina with 75 vol.% porosity in the eye sockets of albino rabbits. There were no signs of implant rejection or prolapse of the implanted material over an 8-week follow-up; at 4 weeks after implantation, fibroblast proliferation and vascular invasion were noted, followed by tissue ingrowth by the 8th week. The first outcomes of bioceramic implant in humans (107 patients over a 3-year follow-up) were reported by Jordan et al. in 2003 [60]. Postoperative problems encountered with its use were substantially similar to those observed with coralline HA orbital implants (Bio-Eye ®) but appeared to occur rarely; the incidence of exposure associated with the bioceramic implant was significantly less than that

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reported for the HA ones, and infection did not occur in any patient. In a following study, Jordan and coworkers further confirmed that alumina implant infections are rare [61] and, after reviewing a clinical case series of 419 patients who received a bioceramic orbital implant, estimated an implant exposure rate of 9.1 % with the majority of the exposures occurring after a 3-month follow-up period [62]. Wang et al. [63] reported that exposures of bioceramic implants occurred only after longterm follow-up and were preferentially associated with evisceration, pegging, and prior ocular surgeries, whereas no late side effects were found in enucleated eyes. Implant wrapping was demonstrated to be an excellent strategy to avoid exposure and related complications [64]. In case of exposure, successfully treatment (without the need for implant removal) can be performed by covering the exposed area with appropriate patches of biological origin (e.g., retroauricular myoperiosteal graft containing myofibrovascularized tissue) [65].

Bioactive Glasses and Glass Ceramics Glasses are noncrystalline ceramics that, due to their attractive optical properties, have been used for centuries to correct visual deficiencies, like myopia or other refractive diseases. As first demonstrated by Hench and coworkers in the early 1970s [66], a special subset of biocompatible glasses, referred to as bioactive glasses, exhibit the unique property to bond to host bone stimulating the growth of new tissue. Therefore, bioactive glasses are recognized as ideal materials for bone substitution and have been widely investigated over the years in the form of dense implants, fine particulates, coatings on metal prostheses, and 3-D porous scaffolds mainly for orthopedic and dental applications [67, 68]. Few bioactive silicate glass and glass-ceramic formulations have been occasionally proposed in ophthalmology for the manufacture of porous orbital implants in the form of single-phase materials or as a second phase added to a polymeric matrix (composites; see the section “HACoated Alumina Implants”). There is a general paucity of scientific literature on this topic, although it has been recently highlighted that bioactive glass can have a great potential for application in ocular surgery [5]. In the late 1990s, Xu and coworkers [69] implanted bioactive glass-ceramic porous orbital implants in enucleated rabbits and observed no rejection during a 6-month postoperative follow-up; ultrasound examination revealed a venous-flowlike spectra in the implants after 3 months, and histological analysis showed that around 90 % of the implant pores were filled by fibrovascular tissue after 6 months from operation. Encouraged by these promising results, the same authors implanted glass-ceramic orbital devices in 102 human patients, declaring a success rate of 96.1 % (98 cases) [70]. In four cases the conjunctiva was torn partly when suture stitches were taken out of the wound, and one patient needed the implant removal. There were no reported complications after a follow-up of 6 months to 2 years, and all patients were satisfied with their cosmetic appearance, although implant drilling and placement of the motility peg to connect the implant to the ocular prosthesis were never performed as a secondary procedure.

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The use of bioactive glass to fabricate orbital implants was also claimed in a recent patent by Richter et al. [71], but no manufacturing or clinical studies have been reported yet in the literature on this type of implant. Another interesting application involves the use of bioactive glass as a “contingency plan” to fill old peg tracts and to permit repegging in porous HA orbital implants, if the initial drilled tunnel was not perpendicular and central to the implant surface [72]. This approach has been reported in a study on three patients who had persistent problems with their pegged HA orbital implants and no longer responded to conservative treatment. After removal of the old peg, the hole was partially filled with bioactive glass, and after 2 months, two patients also underwent successful redrilling of the implant followed by insertion of a new titanium peg, with satisfactory connection to the ocular prosthesis and absence of complications over a 3-year follow-up.

Composites and Coatings This section provides an overview of the polymer/ceramic composites and bioceramic coatings that have been proposed in recent years for the manufacture of orbital implants; their main characteristics are summarized in Table 5.

HA/Silicone Composite Implant In the early 1990s, Guthoff and coworkers [73] developed a composite orbital implant comprising a hemispherical anterior part made of synthetic porous HA to guarantee tissue integration and joined to a posterior part that was manufactured using silicone rubber; the horizontal and vertical eye muscles were sutured crosswise in front of the implant to ensure better stability and motility. Overall implant biocompatibility was excellent, and the transmission of the motility to the ocular prosthesis was generally acceptable [74, 75]. At present this implant is mainly employed in Europe; its diffusion is limited due to the high cost and complex surgical procedures needed for its implantation compared to “standard” spherical porous implants made of HA, alumina, or PE.

Bioactive Glass/PE Composite Porous Implant A couple of studies reported by two different groups of Chinese researchers are currently available on this type of implant. In 2006, the effects of the incorporation of bioactive glass powder on the fibrovascular ingrowth occurring in porous PE orbital implants were investigated in rabbits [76]. Forty-eight rabbits were divided into four equally sized groups, according to the different surgical techniques and implanted materials used: groups 1 and 2 received porous PE after enucleation or evisceration, respectively (reference groups), whereas groups 3 and 4 were implanted with

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Table 5 Ceramic-based composites and coatings used in the fabrication of orbital implants Type of composite or coating HA/silicone

Type of implant Implant comprising a hemispherical anterior part made of synthetic porous HA and a posterior part made of silicone rubber

Recipient Human

Bioactive glass/PE

Porous sphere

Human

HA-coated alumina

Porous sphere

Rabbit

HA coated with Cu-containing mesoporous bioactive glass

Porous sphere

In vitro tests

Remarks It is commonly known as “Guthoff implant.” It exhibits good postoperative outcomes but has high cost and requires complex surgical procedures of implantation Absence of a clear improvement in implant fibrovascularization with respect to porous PE. Studies involving a higher number of patients would be necessary for a more exhaustive assessment Absence of a clear advantage over bare alumina orbital implants Encouraging antibacterial results against S. Aureus and E. Coli

References [73–75]

[76, 77]

[78–81]

[82]

bioactive glass/PE composite porous implants after enucleation or evisceration, respectively. Histological examinations revealed that there was no statistically significant difference with regard to fibrovascular ingrowth among the four groups up to 8 weeks of postoperative follow-up. Apparently, the inclusion of bioactive glass particulate did not significantly promote the rate of fibrovascular ingrowth into porous PE orbital implants. In 2011, the clinical outcomes of 170 enucleated patients receiving bioactive glass/PE composite porous orbital implants were reviewed by Ma and coworkers [77]. The majority of patients experienced no complications (161 cases) and had comfortable socket characterized by good implant motility, without conjunctival thinning or inflammation; excessive discharge and implant exposure occurred in two and seven cases, respectively. All exposures were successfully treated with antibiotics or additional surgery; secondary surgeries were required by some patients but were not due to implant-related complications (ectropion repair in five patients and volume augmentation in three patients). These early results suggest that bioactive glass/PE composite implants may be an interesting option for orbital reconstruction,

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but comparative studies are necessary to definitely estimate their performance with respect to the other available – and routinely used – implants.

HA-Coated Alumina Implants The clinical effects of a synthetic HA coating on the struts of a porous alumina implant were first investigated by a group of Korean researchers in the early 2000s. This experimental implant was fabricated by the polymer sponge replication method: the porous alumina skeleton acted as load-bearing structure, whereas a 20-μm-thick HA layer deposited on it was advocated to provide superior biocompatibility and better long-term stability in the eye [78]. Animal studies in eviscerated rabbits receiving 12-mm-sized HA-coated alumina spheres with different pore sizes (300, 500, and 800 μm) revealed peripheral fibrovascularization of the implant in all groups after 15 postoperative days and also at the center of the implant after 28 days; fibrovascularization was more predominant in the group of implants having 500-μm pores compared to the other two types [79]. In 2002 Jordan et al. [80] reported a comparative study on the implantation of experimental alumina implants coated with HA or calcium metaphosphate in rabbits. Both types of implant had multiple interconnected pores and, in comparison to the uncoated device, the coatings increased the size of the trabeculae from 150 to 300 μm; therefore, the pores appeared smaller but still ranged in the 300–750-μm range. There was no clinical difference in the socket response between coated and uncoated implants, and fibrovascularization occurred uniformly throughout each implant at 4, 8, and 12 weeks after implantation. A few years later, Chung et al. [81] investigated the fibrovascular ingrowth and fibrovascular tissue maturation of HA-coated porous alumina implants in comparison with commercial HA spheres in enucleated rabbits over a 24-month follow-up and found no significant difference between the two groups. No other studies about HA-coated implants have been published; probably, the absence of a clear advantage from a clinical viewpoint (HA coatings did not appear to facilitate or inhibit fibrovascular ingrowth with respect to uncoated implants) and the presence of significant amounts of CaO as a contaminant (related to the coating manufacturing) [80] discouraged the researchers from performing further investigations in this direction.

HA Implants Coated with Mesoporous Bioactive Glass Very recently, Ye et al. [82] coated macroporous HA orbital implants with a thin layer of CuO-containing mesoporous bioactive glass (Cu-MBG): the aim of this research was to combine the antibacterial effect (copper shows potent antibacterial activity in suppressing a range of bacterial pathogens involved in hospital-acquired infections [83]) and drug delivery capacity (drug molecules can be hosted within the glass mesopores [84]) of the Cu-MBG coating to improve the final outcomes of

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anophthalmic socket surgery. Cu-MBG coatings with 0, 2, or 5 mol% of CuO were prepared by dipping the porous HA implants into the sol precursor of the mesoporous glass, followed by evaporation, aging, and calcination. With the peculiarity of releasing antibacterial ions as the Cu-MBG degrades (viability of S. Aureus and E. Coli was inhibited) and good drug uptake/delivery ability (in this study ofloxacin), Cu-MBG coating could be a promising, multifunctional tool in the prevention of implant-related infections.

Summary and Outlook The available literature seems to demonstrate that the performances of porous orbital implants, including the commercial ceramic ones (porous HA and alumina), are superior to those of nonporous devices, such as solid polymeric spheres (silicone or PMMA) and Allen implants (PMMA); the interested reader is addressed to some recent publications for further details [2, 4]. In summary, there is convincing evidence that the exposures occurring in porous implants are more amenable to conservative management without a second operative procedure; in contrast, exposures in nonporous implants, unless very limited, almost always require implant removal [2]. This peculiarity is possible due to the fibrovascularization of the porous implant in vivo, as vascular ingrowth not only helps to anchor the implant in situ but also discourages bacterial colonization of the surface. In the case of ceramic orbital implants, wrapping is useful to decrease the risk of exposure, since the smooth wrapping material acts as a barrier between the overlying soft tissue and the micro-/ macro-rough surface of the implant [44, 85]. Furthermore, a frontal peg can be inserted in porous orbital implants to improve the motility transmitted to the ocular prosthesis. The use of porous devices is discouraged in children as a subsequent implant exchange will be necessary later since the patient is growing, but implant removal is difficult due to fibrovascularization [86]; at present, nonporous polymeric implants (e.g., solid silicone and PMMA spheres) remain the preferred choice by surgeons for the pediatric population [87]. It is instructive to report part of the results of a recent questionnaire addressed to UK ophthalmologists to evaluate current clinical practice in the management of the anophthalmic socket [88]. The surgeons’ responses indicated that 56 % used porous orbital implants (HA, alumina, or PE) as their first choice, and HA (coralline or synthetic) was the preferred option; most porous implants were spherical (diameter 18–20 mm), with only a minority being egg shaped or conical; the majority of surgeons wrapped the implant after enucleation using salvaged autogenous sclera (20 %), donor sclera (28 %), or synthetic polymeric mesh (42 %); only 7 % placed motility pegs in selected cases, usually as a secondary procedure; postoperative exposure occurred in 14 % of cases, and extrusion was reported by 4 % after enucleation and 3 % after evisceration. In summary, this survey highlights that most UK surgeons use porous orbital implants with a synthetic wrap after enucleation, but only a few perform motility pegging. The validity of these results may be

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reasonably extended to the whole European area; however, in other areas of the world, different options may be preferred. From a general viewpoint, the choice of an “optimal” orbital implant is deeply influenced by many “extra-material” factors, including the overall cost and economic viability of the patient, the specific characteristics of the injury/disease, the experience/opinion of the surgeon, and the patient’s clinical history and age. The clinical performance, however, is strictly related to the physicochemical, mechanical, and biological properties of the implanted material; in this regard, it is useful to summarize few key concepts and to provide some highlights for material improvement and future research directions. Looking at the microstructural characteristics of ceramic orbital implants, there is a paucity of relevant studies in the literature since the majority of reports focused on the in vivo biological compatibility and postoperative performances, giving less importance to the assessment and understanding of the basic properties of the materials. From a general viewpoint, it is known that cellsubstrate interactions at the micro- and nanoscale can be regarded as one of the major factors ultimately determining the long-term performance of a biomaterial/implant in situ [89]. Mawn et al. in the late 1990s [39] and more recently Choi et al. [90] investigated by SEM the microstructural and architectural features of coralline HA, synthetic HA, and alumina porous orbital implants, also providing a comparison with other polymeric devices available on the marketplace. As shown in Table 4, there were marked variations of crystal size/shape and surface topography among the analyzed orbital implants. The authors of these studies suggested that surface roughness could influence the inflammatory response after implantation, and crystal size could determine the material-induced phagocytic response: in fact, bioceramics with crystal size around 2–3 μm (coralline and synthetic HA) showed greater tissue reaction in comparison to implants with finer grain (alumina), which was probably due to increased phagocytic activation by crystals of larger size. In this regard, Nagase et al. [91] showed that smooth HA crystals have been associated with less inflammation than sharp-edged crystals. From these results it is still impossible to unequivocally claim that one porous material is clearly superior to the others, even though alumina, exhibiting excellent biocompatibility and favorable microstructural features, seems a promising candidate. An additional issue to be considered is the effect of micro- and nanoscale topography on bacteria, since cells may have to compete with pathogens in the ocular environment. In a fascinating scenario, the surface topography of ceramic orbital implants could be purposely designed (e.g., by micro- and nano-fabrication techniques or by the optimization of the sintering temperature and time to develop crystals with a specific size thus creating a customized surface roughness) to encourage cells to colonize while limiting bacterial adhesion and risk of infection [89]. Pore size, interconnectivity, and overall macroscale architecture also influence the success of an orbital implant. The presence of an interconnected porous network is of crucial importance to promote fibrovascularization, as blood vessel access favors immune surveillance and permits treatment of infections (often following implant

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exposure) via systemic antibiotics. Rubin et al. [92] studied the vascularization in porous HA orbital implants with various pore size and suggested that the optimal pore diameter to encourage fibrovascular tissue ingrowth should be around 400 μm. Other issues deserving careful attention concern the material surface chemistry and response to biological fluids. It has been demonstrated that the presence of an interconnected 3-D network of macropores is a sufficient condition per se to encourage fibrovascularization in orbital implants (like in the case of porous HA and alumina); the use of bioactive glasses in implant manufacturing could carry the significant added value of eliciting specific, desired responses via the release of appropriate ions [93]. In this regard, it is instructive to underline the importance of the study by Ye et al. [82], who coated a porous HA orbital implant with a layer of Cu-containing MBG: although these authors focused on the antibacterial effect of released Cu2+, it was also demonstrated that Cu2+ induces migration and proliferation of endothelial cells during in vitro culture, which could lead to an improved fibrovascularization of the orbital implant in vivo. Looking at the future, appropriate design of bioactive glass composition could impart angiogenetic properties to the material for enhancing the fibrovascular ingrowth, with consequent improvement of clinical performances. Indeed, these smart bioactive glasses should be characterized by high stability once implanted in the ocular environment, as orbital implants are intended as permanent devices, i.e., they must remain in situ indefinitely during the patient’s whole life without undergoing degradation to ensure an adequate socket volume replacement. The incorporation of zirconia (ZrO2) in a bioactive silicate glass has been recently proposed, intended to act as a radiopaque phase in composite bone cements for better visualization under radiographic imaging [94]; a similar approach could be useful also in the field of orbital implants for better assessment of implant position and to detect problems of undesired postoperative migration. Bioactive glasses also have other two important advantages with respect to HA and alumina, i.e., they can be processed at lower sintering temperatures and have a lower density. The latter could be an important added value to reduce the risk of migration downward with possible ectropion and incorrect replacement of orbital socket volume. In order to limit the issues related to bacterial colonization of the implant, deposition of an antiseptic Ag nanoclusters/silica composite layer on the surface of ceramic orbital implants could be a valuable strategy, as recently suggested by Baino et al. [95]. This coating, whose thickness can be properly modulated in the 1–1000nm range, is produced by co-sputtering from silver and silica targets, exhibits a good adhesion on a wide range of substrates (glasses, crystalline ceramics, polymers), and has a good durability in biological environment [96, 97]. A final interesting issue is related to the mechanical properties of orbital implants that, especially if made of ceramic materials (HA, alumina), are remarkably stiffer than the ocular globe as well as the surrounding tissues like orbital fat. Indeed, the use of stiff biomaterials has a number of advantages from an operative viewpoint – e.g., the surgeon can easily handle and place the implant within the orbit with a great control over its position – but compliance mismatch between implant and overlying

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conjunctiva/soft tissues, in combination with repetitive movement of the implant by the extraocular muscles, might contribute to inflammation and soft tissue necrosis, leading to implant exposure. Therefore, future research directions toward an ideal orbital implant might consider the use of novel polymer/ceramic composites to obtain more compliant biomaterials; potential options for the polymeric phase might be adapted, for instance, from the field of experimental vitreous substitutes, such as hydrogels that are biocompatible, porous, and able to absorb water and thus have similar physico-mechanical properties to living ocular tissues [98–100].

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Current Implants Used in Cranioplasty

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Dumitru Mohan, Aurel Mohan, Iulian Vasile Antoniac, and Alexandru Vlad Ciurea

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Applications of Biomaterials in Cranioplasty . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . General Aspects About Neurosurgery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Specific Aspects About Cranioplasty . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Available Biomaterials for Cranioplasty . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Requirements for Biomaterials Selection . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Metals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Bioceramics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biopolymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biocomposites . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Recent Developments and Future Approaches . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Modern Manufacturing Techniques for Personalized Implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . Case Report: Cranioplasty with PEEK Personalized Implant . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Case Details . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Clinical Case Follow-Up . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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D. Mohan (*) • A. Mohan University of Oradea, Oradea, Romania e-mail: [email protected]; [email protected] I.V. Antoniac University Politehnica of Bucharest, Bucharest, Romania e-mail: [email protected] A.V. Ciurea University of Medicine and Pharmacy “Carol Davila” Bucharest, Bucharest, Romania e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_59

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Abstract

The loss of cranial bone integrity due to a trauma, a surgical intervention, or the natural aging process is a contemporary example of tissue failure, which usually requires the permanent or temporary implantation of a bone substituent and may become challenging in case of large defects. Additionally, the neurosurgical procedures developed for restoring cranium defects require an optimum aesthetic outcome and avoidance of artifacts during imagistic investigations. The current research includes both polymeric materials and ceramics, due to the need for different material delivery forms. Small-sized cranial reconstructions are usually performed with cements, for which the calcium phosphates and other ceramics seem to be a reliable solution, as well as polymethyl methacrylate (PMMA). However, the larger-sized bone defects are the more challenging applications in cranioplasty. Nowadays, the polyether ether ketone (PEEK) seems to be a reasonably material for large-sized implants, especially for its physical and mechanical characteristics, as well as for its suitability for tridimensional printing techniques, which may allow for better aesthetic outcomes. In the next step, the use of biocomposite materials will also improve a wide set of biomedical applications, due to the large number of properties that may be acquired through proper manufacturing control. Both the structure and the biological behavior may be improved through phases’ variation and filler distribution. Keywords

Cranioplasty • Neurosurgery • Injury • Personalized implants • PEEK • Case report • Clinical follow-up

Introduction The loss of cranial bone integrity due to a trauma, a surgical intervention, or the natural aging process is a contemporary example of tissue failure, which usually requires the permanent or temporary implantation of a bone substituent and may become challenging in case of large defects. Additionally, the neurosurgical procedures developed for restoring cranium defects require an optimum aesthetic outcome and avoidance of artifacts during imagistic investigations. The main objective of a bone graft is the osseointegration, which strongly depends on factors like extracellular processes, resorption of the implanted material, tissue vascularization, mechanical support, sterilization procedures, or inflammation-related mechanisms. Until now, the autologous bone is the standard material used in cranioplasties due to its biocompatibility, osteoinductive and osteoconductive properties, and chemical composition. However, autologous bone’s main disadvantage is the limited supply (because only a few body parts can provide an optimal bone graft), as well as the multiple surgical procedures, which will lead to anesthetic scars and will increase the healing time [1–3].

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The use of xenografts (e.g., bovine bone tissue fragments) as bone replacement may increase the risk of biological contamination, infections, and faster resorption (before the complete bone regeneration). Other solutions, like allografts (human bone grafts, sampled from another individual), are associated with difficult harvesting and storage procedures, as well as high risks of tumoral cell transmission or pathogenic agents, and sometimes with host incompatibility [1]. Since each of the previously mentioned solutions has its own limitation and the requirements for bone substitution materials are rapidly increasing, the research and development of alloplastic materials became an important objective of biomaterials science. The synthetic materials have an increased availability and may easily fulfill the various requirements imposed by their use in a large series of medical devices. Furthermore, these materials have a lower risk of biological contamination or immunological incompatibility [1, 4]. Various alloplastic materials may be currently used as bone tissue replacements. An ideal bone substitute will have to mimic the structure and function of the original tissue; to be osteoinductive, osteoconductive, and biodegradable; and to have an appropriate chemical composition and porosity which will allow for tissue vascularization and new bone formation. Also, the ideal bone substitute should ensure the mechanical support for the newly formed bone, until its complete regeneration is fulfilled. Finally, sterilization resistance, ease of storage, and economical processing methods are strongly needed [4–8]. The first materials used in primitive cranial bone reconstructions were the precious metals, along with gourds and coconuts crusts. Cranioplasty was not mentioned until the sixteenth century, when the material of choice was gold, and became efficient only after the beginning of antiseptic practices. In the beginning of the twentieth century, silver was introduced as a cranial bone reconstruction material and as a more economical and easier to shape alternative to gold, but the new material could not provide the adequate mechanical strength nor the biocompatibility required by this type of surgical procedure. The research for an optimum cranioplasty material continued with platinum, lead, and aluminum, and soon after this step, the alloys were also considered viable solutions. Vitallium and multiple classes of steel were tested, but currently only titanium is used in cranial reconstruction. An important development in cranioplasty materials began with the introduction of bone substitutes in the current practice. Celluloid was first used because it does not interact with the dura, and later, acrylic resins and tantalum were considered better solutions. Both are still used, especially acrylics (polymethyl methacrylate is still the first option as a material for cranioplasty). The current research includes both polymeric materials and ceramics, due to the need for different material delivery forms. Small-sized cranial reconstructions are usually performed with cements, for which the calcium phosphates and other ceramics seem to be a reliable solution, as well as polymethyl methacrylate (PMMA). However, the larger-sized bone defects are the more challenging applications in cranioplasty. Nowadays, the polyether ether ketone (PEEK) seems to be a reasonably material for large-sized implants, especially for its physical and

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mechanical characteristics, as well as for its suitability for tridimensional printing techniques, which may allow for better aesthetic outcomes. However, PEEK implants are still expensive, and their lack of bioactive properties may lead to displacement and infection-related complications. The current step in cranioplasty materials development is testing of composite materials, especially as biodegradable implants. Many composite formulas, constantly improved due to tissue engineering contribution, have already been tested with good results, both in vitro and in vivo. Alloplastic materials will continue to be upgraded for providing optimal solutions when autologous bone is not available for cranioplasty [9]. This chapter describes various alloplastic materials that are currently used or tested as a cranioplasty material as well as some of the future approaches from this area.

Applications of Biomaterials in Cranioplasty General Aspects About Neurosurgery Neurosurgery is defined as “an area of medicine that includes diagnosis and treatment of tumours, infections, hematomas, degenerative disorders or others legally treatable entities by surgical point of view, identified in the central, peripheral or autonomic nervous system” [10]. The main divisions of neurosurgery are (a) The vascular and endovascular neurosurgery, which deals with the diagnosis and treatment of various diseases of the central nervous system by using catheters and radiological techniques (b) Stereotactic, functional, and epileptic neurosurgery, a branch targeting patients with neuromotor disorders, epilepsy, and obsessive-compulsive disorder and uses nerve stimulation techniques mainly guided by microelectrodes (c) Oncologic neurosurgery responsible for treating tumors located in the brain and spinal cord (d) The spine neurosurgery which refers to treating spine conditions located in the meninges or spinal cord (e) The cranial neurosurgery that diagnoses and treats various lesions of the skull, located both at bone tissue level as well as at the level of the central nervous system components (f) Peripheral nerve neurosurgery that deals with restoring the peripheral nervous system components A field of neurosurgery that is being treated separately is pediatric neurosurgery, which requires multidisciplinary approaches and uses methods specific to each area of neurosurgery, where the use of biomaterials has been documented separately [11–13].

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Both spine surgery and cranial surgery intensively use biomaterials. For spine surgery, the use of biomaterials began with the use of metallic biomaterials for improving the mechanical properties of medical devices like stabilization systems which use screws and pins. Currently, the surgical interventions for spine bone reconstruction are applicable both in the case of degenerative diseases as well as traumatic pathologies. Most medical devices designed for such applications use autologous bone, but its limitations may affect the performance of the intervention. Biocomposite or bioceramic materials are used mainly for the healing of bone defects caused by osteoporotic spinal fractures and bone infections caused by osteomyelitis, as well as for replacing bone tissue fragments that are removed to facilitate the extirpation of cysts or tumors [14]. In cranial neurosurgery, a wide variety of biomaterials are used either as [15]: – Carriers of drugs for treating neurological disorders – In the form of neuron electrodes for the recovery of neurological function – Grafts, to restore the nervous and bone tissues’ integrity The nervous tissue is superior to bone tissue both structural and physiologically. It needs to receive, decode, and transmit information throughout the entire body, and its functional unit, the neuron, has lost its ability to divide. The neuron is however capable of cellular division and may be involved in tissue regeneration. The use of biomaterials that support the nervous tissue restoration often involves the use of additional growth factors or cells. By having a much more limited regeneration capacity, the central nervous system challenged tissue engineering to provide an optimum substrate which allows for the neuronal infiltration and proliferation without compromising tissue vascularization or generating a host response. Some hydrogels based on hyaluronic acid and laminine already proved their capacity to support the nerve tissue restoration, while the biocompatible single-walled carbon nanotubes are intensively tested for the same reason [15]. Finally, regarding cranial bone defects, it is well known that spontaneous skull ossification occurs only until the age of two, so after this age the use of bone substitutes becomes necessary. Different surgical techniques may be used with a common purpose of craniofacial reconstruction. However, each method and material has limitations and may lead to poor resorption, release of toxic products generated by the material’s biodegradation, aesthetic irregularities, biological contamination, or infection development.

Specific Aspects About Cranioplasty Cranioplasty is a surgical procedure performed in order to treat the cranial deformations or defects [4, 16, 17] and aims to restore the tissue subjected to injury, trauma, or tumor removal. Also, the infections and congenital defects could be mentioned. The main indications for cranioplasty are decompressive craniectomies, but a cranial

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reconstruction may also be recommended after tumor removal, to restore a posttraumatic cranial defect, or in case of an infectious osteitis [4, 18]. Indications generally accepted for cranioplasty are brain protection to the underlying bone defect produced by direct mechanical trauma and aesthetic purposes, especially when bone defect is situated in frontal region. Cranioplasty is also indicated with less sustainable arguments for: • • • •

Combating posttraumatic epilepsy Brain protection against direct atmospheric pressure Preventing structures deviation of median line Combating “Burr syndrome” (symptomatic spectrum that occurs in patients with bone defects: headache, dizziness, fatigue, insomnia, and depression) • Neurological improvement by improving cerebral blood flow underlying bone defect • Normalization of the intracranial pressure The favorable outcome of this procedure mostly depends on the technical skills of the neurosurgeon but also on the adjacent soft tissues’ condition and, obviously, on the type and size of the bone defect [16, 19], which may be congenital or gained [20]. In pediatric neurosurgery, the cranioplasty also supports the development of the nervous system [16]. Bone defects smaller than 2–3 cm, when they are located in regions covered by hair or on temporal region, where the temporal muscle disguises and protects aesthetic, do not require cranioplasty. Optimal time for practicing a cranioplasty is between 6 months and 1 year after surgery that caused the defect. Post-surgery replacement of the bone tissue with a new material has both protective and aesthetic roles. Furthermore, the cranioplasty may reduce some neurological symptoms like depression or headache [4, 18, 19, 21]. However, excessive inflammation, infection, displacement, bone resorption, seizures, intracranial hematoma or hemorrhage, and wound breakdown are sometimes identified as complications of cranioplasty [17, 22, 23]. Still, many of these complications are strongly related to the defect’s dimension and location rather than the material used for reconstruction. For most of the neurosurgeons, autologous bone is the recommended option as a cranioplasty biomaterial. However, bone resorption is one of the major disadvantages of most autologous bone cranioplasties, but this strongly depends on the cranial defect dimensions and neither on its cause nor location. Hence, the number of associated secondary surgical interventions may be twice as high in comparison with some alloplastic materials like hydroxyapatite or polymethyl methacrylate. The results are yet contradictory and require further research [19]. On the other hand, infection may be induced by the low blood supply at the implantation site, which is influenced by radiation treatment or repeated surgical procedures. The risk of infection is also associated with the reconstruction location [19]. The infection derived from cranioplasty or vertebral reconstruction affects both the patient and the implanted material: the patient will suffer severe pain while the implanted material will degrade and eventually fail.

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Clinically, two forms of infection are significant and easy to distinguish: the first one is the overt infection (which is revealed through an increase of body’s temperature) and the second one is the inflammatory response (characterized by localized warmth, pain, and sometimes redness and swelling). Some less severe forms of infection located near the implanted area may be difficult to detect because the same bacteria characteristic to biomaterial-related infections may produce specific substances that allows them to avoid the host’s immune system. The first steps in infection’s treatment are the removal of the implant and the elimination of the infectious agent through proper medication. Currently, the cements doped with antibiotics (gentamicin) are an efficient solution for diminishing the infection-related risk. However, a possible disadvantage arises from the fact that almost 80 % of the antibiotic will be fixed in the cement’s internal structure, and it may be released after months from implantation through microcracks [24].

Available Biomaterials for Cranioplasty Clinical Requirements for Biomaterials Selection The diagnostic and treatment methods for cranial defects or deformations are considerably developed in the last years, and many of these developments have been influenced by [3]: • The use of modern imagistic techniques like CT or MRT, for tridimensional localization of bone and soft tissues. These techniques have a superior accuracy (only a few millimeters) and are also used, along with modeling software, to localize or to stimulate bone defects. • The use of modeling techniques for the simulated defect restoration and pre-shaped implant manufacture, based on a virtual model (generated by a computer). • The biomaterials used for hard tissue reconstruction. Since the cranial reconstruction has vital, functional, and cosmetic outcomes, the choice of biomaterials must be preliminary to the surgical intervention. Among these, the biomaterial choice for cranial reconstruction is a very important step for performing a successful surgical intervention. Besides material characteristics, other aspects may be considered [3, 19]: • Indication for the procedure (if a tumor disease is involved or not) – if the hard tissue reconstruction is necessary after a tumor removal, it is very important for the material to allow for future imagistic investigations. • Location and size of the implant – some locations (frontal sinuses, ear, etc.) might be subjected to failure due to anatomical factors, reduced soft tissue coverage, or the presence of bacteria flora [19]. In this case, an inflammatory

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response is expected both intraoperatively and long time after the surgery, so the antibiotic prophylaxis is recommended [3]. Large-sized implants have to be covered by the adjacent tissues and a sized implant will affect the local vascularization, which may lead to complications. Currently, large-sized implants have perforations to allow for tissue vascularization, but long-term stabilization may become an issue for these applications [25]. For the larger cranial defects, material’s aesthetic outcome and mechanical support are the main requirements. • The complexity of the procedure – the complexity of the procedure will increase if both dura and cranial bone must be replaced. The material used for the hard tissue reconstruction should favor the alloplastic dura integration or at least to induce a minimal influence. The easiest method for performing a cranioplasty is the reattachment of the initial bone flap, because it will ensure a perfect fit, a good biocompatibility, and will support bone regeneration [16, 19]. The bone flaps’ storage methods influence biomaterial’s future behavior. Autologous bone may be stored inside patient’s body with the advantage of maintaining the bone cells viability while providing a reduced infection risk but may be uncomfortable for the patient and will induce additional scars. When the storage is performed outside patient’s body, the deproteinization of bone tissue is usually required, and this will affect the cellular regeneration while increasing the risk of bone resorption [16, 18, 26]. Besides the disadvantages induced by the storage methods, autologous bone is associated with a high risk of infection, resorption, poor mechanical behavior, and difficult intraoperative modeling which will affect the aesthetic appearance, particularly in case of large defects [4, 19, 27, 28]. All these downsides led to the development of alloplastic biomaterials, but the choice between them must be evaluated for each individual [19]. Nowadays, the second option for cranioplasty materials, after the autologous bone, is the polymethyl methacrylate, used especially as cement. Although current cranioplasty alloplastic materials are parts of all material classes (metals, polymers, ceramics, and composites), neither one of them represents the optimal solution for human skull reconstruction [17]. When an alloplastic material is chosen for cranial reconstruction, its manufacture technique becomes one of the aspects that need to be considered, besides the one mentioned earlier in this chapter. A novel domain, called “biomanufacturing” and defined as “the use of additive technologies, biodegradable and biocompatible materials, cells and growth factors to produce biological structures for tissue engineering” aims to deliver high-quality devices for medical applications. The final products used as medical devices are available in multiple delivery forms. Among these, implants (biodegradable or nondegradable) and cements have significant importance in cranioplasty. Implants are medical devices designed for replacing, assisting, or enhancing the functionality of bone tissue. Besides the ones replacing the biological structures, many other types of implants are available for

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drug transportation or body function observation [29]. In the specific case of cranioplasty, the implants are conventionally manufactured with respect to the location and size of the cranial defect, by using a plaster impression achieved through the overlying skin, so the general accuracy of the impression will be poor. Furthermore, the presence of edema, hemorrhage, swelling, and overlying muscles will induce further dimensional errors; therefore, additional steps may be required for improving implant’s shape [17]. Currently CAD/CAM and solid free-form fabrication techniques may be used with imagistic investigation methods for developing personalized implants. In cranioplasty these techniques are intensively used for titanium plates, hydroxyapatite, or polymethyl methacrylate pre-shaped implants [17, 30, 31]. Biodegradable implants or scaffolds are porous structures which are implanted for enhancing bone regeneration. These devices also ensure the support required for cell attachment, differentiation, and proliferation [29]. The scaffolds used for cranioplasties must fulfill the same mechanical requirements as the natural bone and should take in the mechanical loads of the newly formed tissue until the bone regeneration is complete. The composites with polymeric matrix and ceramic fillers are the appropriate choice for cranioplasty scaffolds due to their similarity with the bone tissue, both chemically and structurally. Nowadays, the development of this research area led to the use of nanotechnology for improving material performance [17].

Metals Metallic biomaterials are used intensively as load-bearing implants, due to their superior mechanical properties. The most popular applications are the hip and knee prosthesis, as well as a series of wires, screws, or rods used for fixation. The metallic biomaterials are regarded as “bioinerts” because they induce minimal response when they interact with a living body. The main risk associated with these materials is their corrosion in the physiological environment which may lead to the release of potentially harmful ions and to the mechanical damage of the implant. Some metals are widely known for their use in medicine: titanium and its alloys, the austenitic stainless steels, cobalt-chromium alloys, and the precious metals’ alloys. Among these, titanium is mostly used in neurosurgery due to its excellent biocompatibility (the material forms a protective oxide layer at its surface) and to the influence upon the osseointegration process, which allows for a structural and functional connection with the surrounding bone tissue [29]. Titanium is often used with other synthetic materials in neurosurgery [4], mostly because its biocompatibility may be improved by applying a bioceramic coating on the outer surfaces that come in direct contact with the host tissue [1]. Although the use of titanium is associated with high complication rates, a recent study on titanium cranioplasties revealed that the main associated complications are infection and plate removal, which are mostly influenced by the size of defect, the timing of cranioplasty, and location of the implant [32]. The overall complication

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rate may be improved if prefabricated implants are used in the procedure [33]. Also, very recently, titanium meshes were used with bioceramic cements (brushite, hydroxyapatite, and monetite) for large cranial defects’ treatment with very promising results [34].

Bioceramics A bioceramic material is defined as “any ceramic, glass or glass-ceramic that is used as a biomaterial” [35]. These materials are used especially for repairing the locomotive system’s components and have direct contact with both the bone and with the surrounding tissues [36]. Bioceramics are currently used for a wide area of medical applications like lenses, thermometers, or other glass-derived products, optical fibers for endoscopy, and drug delivery agents, as restoration materials used in stomatology or as coatings for a biocompatibility improvement of metallic biomaterials [14, 15, 37–39]. The bioceramics are used for reducing the pain or restoring the impaired, lost functions of the calcified tissues (bones and teeth) existent in human or animal organisms, and their final goal is replacing a tissue that is either aged or degraded, with a material that may ensure the functionality of the initial component for an adequate amount of time (the patient’s whole life or until the complete tissue regeneration), in a non-friendly environment: biological environment (saline and corrosive), at approximate 37  C, under the influence of mechanical loads with different intensities and distributions [39]. In neurosurgery, the bioceramics are mostly used as implants for tissue reconstruction subsequent to trauma or other medical interventions performed on the neural components. The clinical applications exploit the bioceramics in different forms: films, powders, or solid blocks (which are either dense or porous). The bone substituents are generally prepared from porous calcium phosphates. The same materials may be chosen for repairing large bone defects which are not subdued to significant mechanical loads. The dense bioceramics, like alumina or zirconia, are used as components of hip prostheses, while the thin hydroxyapatite films are used as coatings for artificial teeth or metallic components of hip prostheses. All these forms of bioceramics are prepared through application-specific methods and manufacture processes. In the special case of hard bioceramics with a fine microstructure, the strict limitation of the impurities concentration is essential due to the catastrophic effects which they may have regarding the mechanical behavior after the implantation [40]. The use of bioceramics as scaffolds (degradable implants) provides the mechanical support needed during bone regeneration [38, 39] but requires the understanding of the interactions which take place at the bioceramics–tissue interface after implantation. In this case the bioceramics reactivity is an important material feature which allows for their classification and a thorough control of their applications. After implantation, some chemical reactions occur at the material–host tissue interface. After implantation, these reactions will modify the biomaterial surface characteristics and the behavior of the host tissue [39].

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According to Hench et al. [36, 37], four types of implant tissue interactions may be described: toxic, inert, bioactive, and bioresorbable. This information led to the following classification of bioceramics [7, 36–38, 41]: • Inert bioceramics (dense): are those nonporous bioceramics with superior mechanical properties, like alumina or zirconia. For the field of medical applications, the term “inert” is unspecific [35] and should be understood as “a material which will induce a minimal response from the host tissue and will lead to the formation of a thin, fibrous and non-adherent layer on the surface of the medical device” [7], because any material which is implanted will generate a body reaction. The attachment of these materials to the host tissue is accomplished through pressing or cementation and is called “morphological fixation.” • Inert bioceramics (porous): although these materials induce the same type of reaction at the material–tissue interface, they constitute separate category due to their increased porosity which allows a mechanical attachment to the tissue, which is called “biological fixation.” • Active bioceramics: are nonporous materials enabling the chemical bonding with the surrounding tissue (“bioactive fixation”). The most popular materials from this category are the bioglasses, the glass-ceramics, and the hydroxyapatite. • Resorbable bioceramics: these materials have different porosity levels and are slowly replaced by the osseous tissue, while becoming directly involved in the metabolic processes of the body. Calcium sulfate and tricalcium phosphate are illustrative for this class of materials. Another bioceramic classification may be performed based on the regeneration processes (“osteogenesis”) which begin after implantation. Therefore, the first group of materials (class A materials) increase both the osteoconduction (“process of passively allowing bone to grow and remodel over a surface” [35]) and the osteoinduction (“act or process of stimulating osteogenesis” [35]), while the second group (class B materials) will only encourage the osteoconduction [7, 37, 42, 43]. Finally, the bioceramics may be categorized by their chemical composition [40]. This classification is presented in detail in the following pages, including their properties, their specific applications, and the presentation of some composite materials prepared by matching each bioceramic class with other suitable material types. Alumina or aluminum oxide (Al2O3) is one of the most popular ceramics in engineering and medicine. Among its most known uses are the ones of abrasive material, for the manufacture of cutting instruments for laboratories, in the textile industry and paper industry. In the medical field, alumina and zirconia bioceramics are mostly used in dental and orthopedic applications. Aluminum oxide with a purity of 99.99 % was designed as an alternative to metal biomaterials, due to its roughness (the ionic and covalent bonds in the material block the movement of dislocation in the crystal lattice), its reduced friction number, and its good resistance to corrosion and fatigue, properties that recommend its use in medical applications specific to orthopedics. In order to maintain its stability in time,

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alumina needs a reduced porosity, a reduced quantity of additives, and a fine and homogenous microstructure. Decreasing the additives content is necessary to avoid their concentration at the grain boundary, because this will also decrease the fatigue resistance, especially in corrosive environments, while a fine and homogenous microstructure inhibits static stress that interferes with the loaded material [37, 40]. A major drawback of alumina bioceramics is given by their high modulus of elasticity (it is 10–50 times higher than the one of the bone tissue), which can additionally load the bone tissue during its regeneration process. Therefore, the use of aluminum oxide in these applications should be performed with respect to the age of the patient, the type of disease, and the biomechanics specific to the application [37]. However, the aluminum oxide has proven its non-cytotoxicity after being tested in cell cultures. After being tested through the implantation of some biomaterial samples in rabbits orbit, alumina samples have proven good acceptance of the material from the host tissue, cellular and vascular proliferation, and bone growth [41]. In biomedical applications, the aluminum oxide is used especially in orthopedics and dental applications: in knee prosthesis, maxillofacial reconstruction, or as a substitute for bone or in dental implants. In neurosurgery, the use of alumina has drawn the specialists’ attention due to its mechanical resistance and the aesthetic benefits which resemble the acrylic materials. The experimental studies conducted on dogs have shown that the bone tissue can regenerate in the presence of multiple ceramic implants, but the aluminum oxide had a superior behavior compared to the titanium, with respect to mechanical properties and the compatibility with the bone tissue [44]. Furthermore, the addition of yttrium to the aluminum oxide offers radioopaque properties [4]. Its usage in cranioplasty is documented in an article from 1987 [45, 46]. The zirconium oxide (ZrO2) is another material with an enormous potential as a bioceramic material due to its excellent mechanical properties. Nowadays, it is used in hip prostheses and in dentures, but its development as a biomaterial is possible in other medical devices as well. Until now, the cytotoxicity tests performed on polycrystalline zirconia have proven the nontoxic character of the material. On the other hand, a possible risk agent in the case of using the zirconium oxide is given by the other radioactive elements presents in the chemical composition as residues. The effects of the concentration of radioelements on the tissues and organs have already been characterized through measuring the activity of femoral implants [41]. The glass and the bioactive glass-ceramics are made of high-purity materials through a series of standardized production techniques which ensure high biocompatibility and good osteoconduction properties so the material allows the formation of direct bonds with the bone without the formation of a fibrous capsule [1]. In time, various chemical compositions of glass and bioactive glass-ceramics have proven their functionality in medical applications, especially in orthopedics, thanks to appropriate mechanical properties [37]. By changing the chemical composition of the bioactive glass, one can enhance significantly the behavior of this

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material: new series of bioactive glasses, based on boron oxide and phosphorus oxide, have proven their biocompatibility and their ability to promote the bone tissue regeneration. Porous glasses based on boron oxide are both biodegradable and can be designed with a degradation matched with the regenerating speed of the bone tissue [47]. Their biological properties have been intensely studied and documented, both in in vitro and in vivo [37, 41]. Briefly, the studies made to evaluate the biocompatibility of these materials have proven the absence of toxicity. Moreover, the class A bioactive glass-ceramics have interfered in tissue regeneration by creating chemical bonds with specific actions on the cells and tissues. These phenomena are directly dependent on the resorption speed and the ion quantities released after implantation [48, 49]. Special attention was pointed toward the study of the adherent layer formed on the bioglass–tissue interface, whose chemical composition was estimated as being rich in calcium and phosphorus [41], which are the main chemical elements in bone tissue, after carbon and oxygen [50, 51]. The development of the adherent layer of ceramic glasses is activated in saline or sangvine environments. The chemical bonds between silicon and oxygen are broken with release of salicylic acid, whose condensed form is a negatively charged gel situated on the biomaterial surface. The gel later crystallizes in a calcium phosphate and produces a new layer of apatite at the tissue–implant interface. The initiation of the bioactivity takes place when the apatite layer interacts with bone’s collagen, polysaccharides, and glycoproteins, which incorporate the apatite layers in their structure. These chemical bonds represent the beginning of tissue regeneration [52]. Although glass and bioactive glass-ceramics are characterized by the good biocompatibility, the large-sized bone substitutions made with these materials are long time implanted in the body and affect bone vascularization [5]. Besides this, the major disadvantage of glasses is their poor mechanical behavior: they have low-impact resistance due to their bidimensional amorphous lattices and high risk of crevicing at implantation, due to their fragile character [53]. The applications of bioglasses and bioactive glass-ceramics in orthopedics and dentistry have been documented since the end of the 1980s [6, 40, 45, 54–57]. A series of bioglasses or bioactive glass-ceramics products have been developed to replace bones in the middle ear, to repair dental defects, or to cover the implants used in different biomedical applications [37]. Until now, the ceramic glasses fulfilled all neurosurgical requirements regarding biological behavior, but the classical manufacturing process could not meet the necessary aesthetic appearance characteristic to personalized implants. Currently, the modern tridimensional printing techniques are able to produce personalized implants made of ceramic glass with minimal chemical composition changes, without altering the bioactive properties of the material. In addition, some bioactive glass composites were already used for craniomaxillofacial reconstruction. Thus, a biocomposite made of bioactive glass and bone sampled from the iliac crest had an improved healing time when compared with autografts, and a similar composite, with bone particles sampled upon a trephination, was used for skull reconstruction without side effects or further surgical interventions [52].

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Another development aims for reducing their disadvantages by the addition of new ingredients, for preparing composite materials. An example is the glass-ceramic and titanium oxide composite which has been examined after implantation in patients’ spinal cords. The material presents a good biocompatibility and long-term stability without forming a fibrous capsule on the interface with the body and without artifacts in the MRT investigations [33]. An apatite–wollastonite bioactive glass-ceramic has also been used in spine surgery [58, 59]. Calcium sulfate known as “gypsum” has been used since ancient times as a construction material. Products made of gypsum can be found in a great variety of shapes and purities, and their hydration degree can be adjusted through controlled heating. To obtain the hemihydrated form of calcium sulfate (Ca(SO4)2  ½ H2O) mostly used in biomedical applications, various hydrated forms are being processed and chemically treated in controlled conditions. The calcium sulfate acts as an osteoconduction material and allows the vascularization of the bone tissue and the migration of the osteoblast if it is implanted in direct contact with the periosteum. Its usage increases the adhesion of bone cells with reduced or no inflammation at the implant site. The most important feature of the calcium sulfate is its high absorption speed in a physiologic environment (from several weeks to several months) [20, 40]. Although the material has been documented in literature, its usage is still limited. Applications in the pharmaceutical field have established a precedent in its usage as a biomaterial for orthopedics, and in the last decades calcium sulfate has been used as a bone substitute material, a binder for hydroxyapatite ceramic particles, or in different formulas for composite cements [40]. To avoid the fast absorption of calcium sulfate, the hydroxyapatite was introduced in a composite material designed to ensure the presence of non-resorbable hydroxyapatite particles at implantation site for enhancing bone regeneration. Two mixtures of this kind have been successfully tested on cats with frontal and parietal cranial defects. The study has proven that the material allows the regeneration of the bone tissue, is easily molded, and does not produce infections in the frontal sinus area [43]. Other composites, based on calcium sulfate and calcium carbonate, were tested on rats, through the implantation of some cylindrical samples in the femoral condyle, along with in vitro testing. The study showed that calcium carbonate incorporation in calcium sulfate decreases the biomaterial’s degradation rate while enhancing the bone regeneration [60]. Besides the testing of biocomposites based on calcium sulfate, some studies are being developed to evaluate its use as drugs or growth factor delivery agents or if it has hemostatic or angiogenic properties [40]. A porous material formed of almost equal percentages of hydroxyapatite and calcium sulfate has already been approved to be used as antibiotic delivery system in osteomyelitis [4]. Calcium phosphates are ceramic materials with an excellent biocompatibility, a fact proven by their presence in the human body (bones, teeth, and tendons) [61]. Currently, the calcium phosphates are being used as bone substituents due to their chemical similarity with the inorganic component of bone. These materials have the ability to stimulate the regeneration of the bone tissue, by actively

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interfering in the bone remodeling processes. However, calcium orthophosphates are difficult to use as implant materials or for large-size applications, due to their mechanical properties. Calcium phosphates are currently used in orthopedics and neurosurgery in the treatment of bone defects and fractures [6], in spinal cord surgery [62–67], and in cranio-maxillo-facial reconstruction [54, 68–73]. Nevertheless, just a limited variety of calcium phosphates can be used in biomedical applications. For example, the compounds with a Ca/P report smaller than one are acids and are extremely soluble, while others, like tetracalcium phosphate (TTCP), are too basic for the needs of the living body. These materials can be used in medicine just combined with other calcium phosphates or other chemical substances [39], like the bone cement obtained from mixing tetracalcium phosphate (TTCP) and dicalcic phosphate (DCP), in the presence of water. Their reaction is isothermal and leads to the formation of a resorbable dense paste. If a buffer solution based on calcium phosphates is used during preparation, the reaction may be accelerated. Various studies conducted on cats have shown that the calcium phosphate bone cement is easy to shape, and its postsurgical aesthetic appearance was evaluated as “excellent” [32, 16]. Authors also draw attention on bone cement complications, especially on infection and the displacement of large defects restorations. Moreover, cases of mechanical failure have appeared even after trauma with reduced intensity. Apparently, a secondary restoration also accentuates the reabsorption of the material, which may eventually lead to infection. This drawback can be avoided if supplementary stabilization techniques are used [33]. Chemically, most bioceramics based on calcium phosphates are either in hydroxyapatite’s (HA) or tricalcium phosphate’s (TCP) class. Recently, some biphasic compounds HA + β-TCP or HA + α-TCP have risen interest in the scientific literature, and their preparation and biological behavior are being intensively tested [74]. In comparison, the hydroxyapatite has a better stability in a physiological environment and it is reabsorbed in an increased amount of time than tricalcium phosphate. Hence, the success of a type BCP product will be determined by the report between the stable phase (HA) and the resorbable phase (TCP) [39]. BCP scaffolds’ healing properties have been tested with good results in canine mandibular defects [20] The in vivo behavior of a BCP formula was also recently tested in rat cranial defects with bone formation similar to β-TCP [75, 76]. Hydroxyapatite (HA) is a calcium phosphate very stable in aqueous environments, with the chemical formula Ca10(PO4)6(OH)2. HA contains approximately 40 % calcium and 18.5 % phosphorus (in mass percentages) and has a hexagonal crystal lattice. The Ca/P report of the stoichiometric hydroxyapatite is 1.667 and is one of the most important indicators used to evaluate the different processes of obtaining this kind of material [6, 39, 54, 77]. Since approximately 70 % of the human bone is made of a nonstoichiometric form of hydroxyapatite, this material may be described as an ideal bone substitute. Different forms of porous hydroxyapatite lattices are available nowadays to repair bone defects or to support the tissue regeneration in almost the entire body. However, the geometry, porosity, and lattice substitutions are important characteristics which

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affect the material’s capacity to heal bone defects. These properties have already been tested in animal models [8, 20, 78]. The results from the cell cultures conducted to evaluate the biological properties of the nonstoichiometric form of hydroxyapatite have been evaluated with respect to their chemical composition [40]. The study revealed that hydroxyapatite with a higher content of carbonate groups increased the activity of the osteoclasts, and this result suggests that bone reabsorption (the phenomenon underlined by the activity of the osteoclasts) is directly influenced by the functional groups incorporated in hydroxyapatite crystal lattice. Another conclusion targets the fluoride ions substitutions, which stimulate the cell proliferation. Still, a major drawback of hydroxyapatite is its brittle character (specific to bioceramics), the reduced tensile stress, and a high risk of infection after implantation. The larger bone defects may be difficult to repair with hydroxyapatite due to the reduced osseointegration and the structural transformation in contact with cerebrospinal fluid [4]. Various hydroxyapatite implants were evaluated in 2013 in a study that analyzed the clinical data for 1549 cases of patients who have had cranioplasty with personalized devices [63]. The results shown the material’s compliance with the requirements assessed by this kind of surgical intervention and how well it is tolerated by both children and adults. Its brittle character may be balanced by increasing medical devices’ depth and also by permanently improving the surgical technique. In comparison with the polymethyl methacrylate (PMMA), which does not allow for the expansion of the cranium in the growing, developing, or regenerating processes, the hydroxyapatite may be successfully used in pediatric neurosurgery applications. Also, the hydroxyapatite does not induce rejection reactions coming from the organism and ensures a relatively good attachment on the bone tissue [4]. Due to its various delivery forms, hydroxyapatite’s bioresorbable character was evaluated in a comparative study of ceramic blocks and cements, conducted on sheep. The results have shown that differences between the reabsorption of the materials are not significant even after a year from the implantation, but tissue regeneration has been accelerated only for hydroxyapatite blocks, and not for the cement, as it has been expected [32, 16]. Hence, hydroxyapatite cement can be used in cranioplasty, because it does not induce bone growth and is not reabsorbed in the human body. Since hydroxyapatite is also available as cement, this form is used in cranioplasty because it can be easily shaped [6]. A series of porous ceramic blocks or titanium may be added to enhance the mechanical properties of hydroxyapatite-based cement. This may avoid the thermal shock induced by the exothermal polymerization of polymethyl methacrylate. Moreover, the hydroxyapatite-based cement is quickly fixed, after approx. 5 min [33]. Some frequent side effects of cranioplasties made with bone cements based on hydroxyapatite are the accumulation of serum [32] and the infections, especially in the frontal sinus area [62]. The complications rate varies between 0 % and 20 % in the retrospective studies reviewed by Neovius and Engstrand [19]. Excellent results, without significant side effects, have been obtained in the case of using grains of

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hydroxyapatite previously mixed with blood to form a paste with adhesive consistency [32]. Tricalcium phosphate (TCP) is rapidly diluted in physiological environment, which allows for bone tissue development after implantation. The TCP is found in four states, among which the most known are the α and β [6]. The β-TCP (β-tricalcium phosphate, β-Ca3(PO4)2) can be prepared through thermal decomposition of the calcium-deficient hydroxyapatite, at temperatures over 800  C, and through bone calcination. Pure β-TCP is not normally found in calcified tissue in human or animal bodies, but a magnesium substituted form, named whitlockite, may be found in renal calculus or arthritic cartilages [51]. This biomaterial has proven its biocompatibility with some results comparable with autografts’ behavior [78]. The α-TCP (α-tricalcium phosphate, α-Ca3(PO4)2 is a state of the tricalcium phosphate obtained by heat treating at approx. 1100  C, but the addition of silicates in its structure has allowed the stabilization at temperatures lower that 1000  C. The α-TCP and the β-TCP have the same chemical composition but are fundamentally different in regard with the crystalline structure and the dissolution behavior: α-TCP has an emphasized reactivity in aqueous environments and is able to be hydrolyzed along with other calcium-based phosphates. Nowadays, the α-TCP is used in bone cements [51]. Both known states of the tricalcium phosphates have a superior solubility in comparison with the hydroxyapatite, and they are reabsorbed faster after implantation in the biological environment. This is why the α-TCP and the β-TCP are used to obtain degradable biocomposite materials. Among these, a composite based on gelatin and β-TCP showed good results in the in vivo testing. The material is biocompatible, osteoconductive, and biodegradable and does not require a new surgical intervention to remove the device [1, 24]. Although the clinical applications of β-TCP are still limited, a bone substitute prepared from beta-tricalcium phosphate was already used in pediatric surgery to evaluate the healing rate of cranial bone defects in 23 patients. The healing was most effective for defects smaller than 40 cm, without graft-associated side effects [79].

Biopolymers The polymers are organic materials which incorporate large molecules in their entities called “monomers.” The monomers’ chemical bounds contribute to a giant chain formation and produce high ductile materials. The polymeric materials are similar to lipids, proteins, and polysaccharides from the biological environment, and the variations among them refer to chemical composition, molecular mass, crystallinity, solubility, and thermal properties. All these characteristics may be “programmed” by choosing the polymer type, its chain length, or by combining two polymers through copolymerization. The polymers can be easily dispersed in complex forms (like gels, sponges, and materials with complex porosity systems) and are nonmagnetic and transparent to

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X-rays, hence compatible with the modern medical imagistic methods. However, most of the polymers do not react favorably in the biological environment because they can generate either an inflammatory chronic response or a series of toxic degradation products. Polymers which however fulfill these biocompatibility requirements (named “biopolymers”) lack the necessary rigidity, ductility, or mechanical properties to resist to major loads or may be influenced by sterilization procedures. The polymers may fulfill the matrix role in biocomposite materials with different properties, including biodegradation, and polymeric materials involved in preparation may be either natural or synthetic, the natural biopolymers being a better choice for biocomposites used in bone substitution due to their excellent biocompatibility and degradation properties [1, 80]. Natural organic materials were studied in order to identify their role in bone tissue functionality, in the same manner in which calcium phosphates were studied for their similarity with bone mineral component [81]. Polysaccharides, proteins, and biofibers were found to be appropriate biopolymers for composite materials because they possess a material structure able to guide the cells for biasing the time- and space-dependent local growth, simultaneously with stimulation of an immune response. In general, a biopolymer can be designed with a certain degradation rate which will allow for a gradual transfer of the mechanical load from the implanted material to the newly formed bone. The requirements for these materials include mechanical strength and lack of toxic degradation products, along with degradation and absorption rates comparable with the bone tissue healing rate. Nowadays, the polymeric materials are especially used for soft tissue replacements or as matrixes for biocomposites. Different types of polymers, like polymethyl methacrylate (PMMA), polyether ether ketone (PEEK), or porous ethylene, are however used in neurosurgical applications [4, 24, 26, 52, 82, 83]. One of the most popular natural biopolymer is the collagen, which joins under this name several types of proteins that have an abundant contribution for the human body. These are indexed with Roman figures, considering the discovery date, and constitute structural unities that are capable to fulfill several connective tissue functions. Different types of collagen can be found in bone tissue, cartilages, or just partially in highly specialized tissues. The type I collagen is present in both soft and hard tissues, the type II collagen is found in bone epiphyses, and osteoblasts secrete a collagen type III fiber matrix along the periosteal surface. The collagen matrixes used in composite materials can suffer some improvements by adding cross-linking agents or by physical treatments like heating, irradiation, or copolymerization [81]. Another biopolymer which may be successfully used as a biocomposite matrix is chitosan, because of its ability to degrade simultaneously with bone tissue regeneration, without any toxic product release or other major inflammatory responses. The use of hydroxyapatite as a dispersed phase enhances the biocompatibility of the newly designed material, especially when HA is incorporated as nanometric particles. Fibrin also presents mechanical properties close to those of the bone tissue. Besides this, its incorporation in a hydroxyapatite composite

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material was tested on rat cranial defects and proved to stimulate bone regeneration [1, 80]. PMMA is the polymerized form of methyl methacrylate, which is a liquid material. During polymerization, the material solidifies over a known amount of time, which makes it suitable for the use in surgical procedures. In chemical applications, the monomer is mixed with PMMA powder. The powder will dissolve and will initiate a polymerization reaction defined by viscosity increase followed by solidification [24]. Additionally, zirconium dioxide may be incorporated in the material for improving the results of X-ray analyses or CT [53]. The main advantages of PMMA are the transparency, the ease of preparation, the mechanical properties, and the low price [53]. Also, PMMA is well tolerated by the body. After implantation a fibrous tissue will evolve at the interface between the material and the surrounding tissue, and the osseointegration will begin with a host reaction. The immune response expresses both locally and systemic and is distinguished by the macrophage activation. The PMMA is used in neurosurgery in vertebra stabilizations and replacements, as well as in cranioplasties, either as cement or as a pre-shaped solid implant. For small-sized defects in vertebras or in the skull, the cement is prepared intraoperatively and then shaped and fixed by the surgeon, while the pre-shaped, solid implant is rigorously dimensioned and designed using patient’s imagistic results, for restoration of larger defects [24]. The use of PMMA cement is limited by its temperature rise during polymerization, the toxicity of the liquid monomer, and by the reduced bone vascularization after implantation. In spine surgery all these factors may lead to bone resorption followed by mechanical failure, while in the cranial neurosurgery the increased temperature may affect the surrounding tissues and lead to disfiguration or may damage the nervous system [24]. Also, loosening of the implant (with both mechanical and biological origins) is one of the high-risk disadvantages in spine surgery because PMMA implants are fixed with metallic rods and screws. This risk has been overcome in cranioplasty which allowed for the development of better fixation systems [24, 53], but the complications rate still varies between 0 % and 23 %, most of them due to the location and size of the cranial defect and to postoperative radiation treatment [19, 84]. Polyether ether ketone (PEEK) is a polymeric semicrystalline material with mechanical properties similar with those of the autologous bone [4, 25, 85, 86] and good resistance to sterilization techniques, without inducing CT or MRI artifacts [4]. PEEK implants are superior to metallic implants in terms of thermal and biological properties. Also, they may be easily shaped and can be manufactured through 3D printing techniques [4, 53]. Until recently, PEEK has been intensively used in spinal surgery [53, 87, 88], but the appropriate properties lead to its usage for the reconstruction of cranial bone defects caused by tumors, traumatic injuries, or infectious lesions [85]. PEEK cranioplasty leads to good aesthetic outcomes and is more comfortable because the material is light and is resistant to thermal modifications [86].

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The main disadvantages of PEEK implants are their expensive prices and the lack of bioactive properties. The last one may lead to further complications like infections, foreign body reactions, and high risk of dislocations [4, 25, 86]. Unfortunately, very few data is available in the literature and the long-term follow-up of PEEK cranioplasty has not been thoroughly documented. The porous polyethylene is a strongly biocompatible material with pore dimensions that allow for cellular migration. Moreover, this material is slightly malleable and does not induce resorption phenomena. When placed in a biological environment, the porous ethylene forms a connective tissue layer which contributes to the stabilization of the implant, along with the titanium screws which are usually used with this biomaterial. However, the porous ethylene has some limitations regarding the malleability and impact resistance, which make it useful just for the small-sized defects. Furthermore, some adverse reactions were observed at the implantation near mucosa. The material is not recognized by X-ray analysis, is poorly detected by CT, and does not induce major artifacts in MRT [3].

Biocomposites Composite materials are manufactured form at least two components which possess different chemical and physical properties. The components will remain macroscopically separate within the new material structure, so composites are heterogeneous materials where in each of the phase conserves its characteristics and main properties and their interaction is intermediated by an interface. The manufacture of this type of material allows the procurement of new, enhanced characteristics which cannot be defined separately for each of the components. Similar with the other types of materials (metals, ceramics, or polymers), a composite material which is biocompatible is termed “biocomposite” [1, 17, 37, 48, 89, 90]. The components or phases of a composite material are defined in two groups: the first component is the matrix (continuous phase), which integrates the material’s volume and constantly supports and maintains the position of the second phase type – the reinforcing material (the dispersed phase). The reinforcing material enhances the matrix’s properties (especially the mechanical properties but some other features like density, biocompatibility, or X-ray transparency as well). Additionally, the composite materials may be designed to fulfill a wide range of preset conditions, by severely controlling the volume fractions and the local or global distribution of the phases. A higher volume of dispersed phase will considerably enhance the mechanical properties, while the long and aligned fibers will prevent crack propagation and will ensure the anisotropic behavior of the composite [1, 90]. The development of a composite material should consider a range of factors like components selection, the choice of proper preparation and processing methods, and aspects of internal and external appearance of the final medical device. During manufacture, the composites are mounted in a preset shape, while the sequence of

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Table 1 Classification of composites based on phases’ dispersion in the material [1, 80, 81] Composite type Simple Complex Graded Hierarchical

Number of dispersed phases Single Multiple Single/multiple Single/multiple

Distribution of dispersed phases in the matrix Homogenous Homogenous Heterogeneous Homogenous/ heterogeneous

Observations – – – A primary composite is dispersed in a second matrix

phase’s addition in the shape depends on the required material and on the previously selected preparation method. Composites preparation requires, of course, the combination of at least two different materials. Therefore, phases’ miscibility, adhesion, and polarity are important factors which interfere in this process. The lack of adhesion between the phases will lead to a failure at the interface, followed by a decrease of mechanical properties. From the chemical point of view, the interaction between the components of a composite material may be strong (ionic, covalent, or coordinative bonds) or weak (van der Waals forces, hydrogen bonds), while a third type of interaction does not imply any chemical bonding between the phases. Composite materials may be categorized based on multiple criteria. Hence, based on phases’ dispersion in the material, one could distinguish between the materials presented in the following Table 1 [1, 80, 81]. Finally, a different classification system for the composite materials depends on the matrix type (polymer, ceramic of metal) or on each component’s type (polymerceramic, metal-ceramic, etc.). Among these different types of biocomposites, the ceramic-polymer ones release the highest concentration of toxic products and are limited by their organization in a preset shape. The metals bring multiple complications due to the corrosion processes while the ceramics coated on metallic implants degrade as the implantation time increases. The ceramic-ceramic composites possess some biological superiority, due to their similarity with the mineral component of bones and calcified tissues, and have an additional advantage as they may be shaped in different forms. Conclusively, the biological response of any material type is strongly influenced by the mechanical load from the body, so an insight into the local biomechanics is imperative for a proper selection of an optimum biomaterial [41]. Polymer-ceramic composite materials have been intensively tested on cell cultures and on animal models: the use of a composite material based on coralline hydroxyapatite and polyglycolic acid has been tested in the rabbit’s skull, and the results confirm the lack of side effects and the ease of fixation on the adjacent bone tissue [91]. The gelatin was also used in another composite material, which has been evaluated in vivo after the introduction of particles of tricalcium phosphate to establish its potential as a substitution material for the cranial bone. The material has proven to be easy to shape, biocompatible, osteoconductive, and biodegradable, with a progressive replacement of the material with the regenerated tissue

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[92]. Another biocomposite material based on gelatin and octacalcium phosphate (OCP/Gel) was tested on rats for different filler concentrations. The results have shown that the material has excellent biodegradable properties and stimulates cell migration [93]. The polymer–β-TCP composites, manufactured using modern tridimensional printing techniques, have also been analyzed from a mechanical perspective. The tests have shown that the resistance to bending of the new material is 5–22 times higher, depending on the used polymer (PCL, respectively, PLA) [94]. A multiple composite material with notable results both in vitro and in animal testing is manufactured from layers of polylactide polymers, amorphous calcium phosphate, and calcium carbonate. The inner layers are porous and biodegradable and are involved in bone regeneration, while the outer layers are more stable for ensuring protection [3]. Currently, one of the most popular biomaterials for cranial reconstruction is based on a polymer matrix (epoxy resin) filled with carbon fibers. This material may be easily shaped during surgical intervention for improving the aesthetic outcome, is resistant to sterilization techniques, does not induce artifacts in imagistic results, is lighter than metals but with superior mechanical properties, and is biocompatible (is able to form a connective tissue layer at the implant–host interface) [3, 17].

Recent Developments and Future Approaches Tissue Engineering The simultaneous evolution of the complementary areas which collaborate for achieving a better functionality of medical applications allows, in our days, for the use of innovative strategies which aim to achieve biomimicry for bone substitute materials. The use of biomaterials with a degradation rate adapted to the tissue regeneration rate requires substantial development in various aspects of a relatively new field, called “tissue engineering.” Tissue engineering is a multidisciplinary domain which harmonizes the principles of engineering and life sciences in order to develop biological replacements for restoring, maintaining, or enhancing tissue functions, and its approaches recommend the use of porous scaffolds impregnated with cells and growth factors which will enhance the bone regeneration. The development of composite biomaterials which will mimic the bone tissue architecture requires a proper identification and characterization of its tridimensional architecture, followed by the replication of an environment which will promote the cellular communication through cells and extracellular matrix’s components and also through highly specialized proteins. All these components enhance the cellular proliferation, differentiation, and migration. Latest approaches regarding biomaterials used for substitution of osseous tissue are related to: • The use of biocomposite materials with nanoparticles: tissue engineering can use specialized techniques for simulating tissue properties at a macroscopic and microscopic scale, but they are still unable to replicate the microscopic aspects

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which are key factors of bone regeneration. Additionally, the inflammatory reaction generated by the cellular activity and by the formation of the fibrous tissue will affect the regeneration process. “Nanobiomaterials” and especially the composites with nanoparticles are considered “promising platforms” which will provide the structural support required by the cells. Furthermore, the biomaterial design aims to introduce components which will guide the cellular behavior [17, 95]. • Involving cellular therapy in biomaterials development: the cellular regeneration may be simulated by incorporating bone growth factors in a polymeric matrix. This may enhance the protection while improving the implant’s aesthetic outcome [19]. Some bone morphogenetic proteins or polypeptide growth factors are already being prepared for further distribution in porous matrixes that will the osteoinduction and bone growth [16] • Ensuring a proper vascularization from the implanted material: the vascularization may be achieved by developing materials with controlled porosity which will provide permeability and proper conditions for oxygen and nutrients diffusion through their structure. Also, the development of some microsurgical techniques may allow the shaping of “axially vascularized tissues” and reduce the contemporary limitations regarding tissue vascularization [96]. • Using modern manufacture techniques (“solid free-form fabrication”) in correlation with high-resolution imagistic techniques for developing personalized biomedical applications. All these tissue engineering aspects require the preparation of scaffolds that resemble the natural bone tissue, but the methods used by classical engineering involve the use of solid, simple, and large materials which will be shaped in smaller and complex forms without the possibility of controlling the geometric parameters (like the ones regarding porosity). This is one of the reasons why the structures required by tissue engineering are incompatible with the current manufacture techniques. In contrast, a modern series of manufacturing techniques named “solid free-form techniques” starts with small-sized units (powders) and organizes them in a preset shape. The solid free-form techniques are described in more detail in the following pages.

Modern Manufacturing Techniques for Personalized Implants Solid free-form fabrication techniques rapidly gained interest for potential uses in tissue engineering but are not extensively used yet due to some disadvantages of their final products like high costs, lack of geometry’s accuracy, low mechanical properties, or lack of adequate raw materials. Within the solid free-form fabrication techniques, tridimensional (3D) printing creates a solid material after the interaction between a layer of powder and a liquid which is atomized above it. The liquid may act as a binder or may generate a chemical reaction which will bond the solid particles. After solidification a new layer will form over the previous one which will also act as a support for the manufacturing of the entire material. The common advantage of all solid free-form fabrication techniques is the possibility of acquiring a diversity of materials. Natural polymers, like polysaccharides,

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may be used with water-based binders without any solvents. From this point of view, a minor disadvantage related to the solid free-form preparation of polymers is the need for organic solvents which may compromise the biocompatibility of the final product. On the other hand, ceramic materials may be used with polymeric binders, which will lead to the manufacture of composite materials, or with specialized cements which dissolve the particles and form new crystals in a ceramic structure [97]. In the medical devices industry, all these techniques have already been implemented in integrated systems, along with digital imagistic techniques (computerized tomography or magnetic resonance) for the manufacture of large models which mimic anatomical characteristics. In neurosurgery, as we mentioned at the beginning of this chapter, the fabrication of personalized titanium and PMMA implants is already available. These implants have a high anatomical accuracy and are prepared through tridimensional printing, followed by surface abrasion and sterilization [26, 30, 98, 99]. Another technique, called “laser cladding,” allowed for the construction of functional products for low-bearing applications, prepared from bioactive glasses. Implant technique did not induce any significant variation in terms of chemical composition while the structural modifications were minimal. Moreover, the bioactive glasses processed this way were tested in simulated body fluid for evaluating their biological behavior, and the results were similar to those of their precursors [100]. Achieving the best results in terms of accuracy and precision of solid free-form fabrication techniques requires advanced imagistic methods and a proper statistical evaluation of precision and reproducibility. High-resolution computerized tomography allows for the scaffolds’ quantitative analysis simultaneous with the investigation of the mineralization process. Beyond any doubt, the understanding of mechanical properties remains an area with large research opportunities, which will be able to offer significant information regarding devices’ behavior after implantation [30]. Finally, the improvement of cranial defects treatment by applying the principles of tissue engineering and modern manufacturing techniques is a complex concept, whose success depends on various interdisciplinary collaborations. Szpalski et al. identified a number of stages of research and development needed to improve neurosurgical techniques: the in vivo animal model testing for evaluating reparation of defects, optimal biomaterials choice, and developing strategies of vascularization and cell proliferation [20]. So it is possible in the future that biodegradable biomaterials manufactured through tissue engineering techniques will cover bone defects for an appropriate amount of time while releasing bioactive molecules which transform the personalized implant into a functional tissue [16].

Case Report: Cranioplasty with PEEK Personalized Implant Introduction Bone lesions caused by traumatic injuries often require the tissue reconstruction through cranioplasty, and the choice of a proper biomaterial is a challenge for neurosurgeons. PEEK implants have already showed good results in cranioplasties,

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although their applications are still limited. The following pages describe a successful cranioplasty performed with a custom-made PEEK implant.

Clinical Case Details History A male patient, aged 21, suffered a craniocerebral traumatic injury due to a car accident. The subsequent evaluation showed that the patient suffered a comminuted cranial bone fracture with osseous fragments blockage, along with a sub-adjacent cerebral hemorrhagic contusion. The first surgical procedure, performed in a local hospital, allowed for the removal of temporal osseous fragments with favorable postsurgical outcomes. The patient regained its consciousness after approximately 2 days, and his neurological status was completely restored. Case Examination and Operation After approx. 8 months from the initial surgical intervention, the patient came at “Sanador Clinic” in Bucharest for a general and craniocerebral evaluation. The initial CT examination revealed a cranial bone defect in the right temporoparietal area, of approximately 23 mm length, near an adjacent right porencephalic cavity in the parieto-occipital area (24/66 mm maximum axial diameters and 27 mm longitudinal diameter). No other unusual aspects were identified by this initial CT scan, as one may observe in Fig. 1. Since a consequent MRI investigation was contraindicated due to a foreign body with metallic density identified in the pelvic region during the CT examination, an EEG investigation was performed without any further observation of pathological aspects. Also, an ENT checkup was carried out with normal clinical outcomes. The second CT investigation confirmed the presence of a sequelar cerebral lesion in the right posterior temporal area and bone loss with 3.5/2.7 mm diameters in the right retroauricular field. Based on these results, a tridimensional cranial bone reconstruction was simulated using dedicated software (Fig. 2). The previously acquired data, the imagistic aspect, and the proper amount of time since the first surgical intervention allowed for performing a cranioplasty. The alloplastic implant was designed based on the CT results, in order to achieve an accurate anatomical shape of the implant. The material of choice was polyether ether ketone (PEEK), recommended by its advantages in this field previously described: PEEK is a light material which ensures mechanical support, is compatible with the personalized implants’ manufacturing techniques, and does not induce artifacts in the imagistic analyses The personalized PEEK implant was surgically implanted and subsequently mounted using two “CranioFix 2” devices (11 mm diameter), and a keloid scar located near the reconstruction site was removed. The intraoperative aspect of the implanted material is presented in Fig. 3.

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Fig. 1 The initial CT examination (8 months after the traumatic injury)

Fig. 2 Cranial 3D reconstruction. The bone defect is emphasized

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Fig. 3 Intraoperative aspect of the cranioplasty material after implantation and mounting

Fig. 4 3D CT examination at 24 h after the surgery. The cranioplasty, the fixation devices, and the drainage tube are revealed

Postoperative Care The surgical intervention was performed without any complications, and the postsurgical outcome was favorable for both short and long time after implantation. A postsurgical CT result is displayed in Fig. 4. The suture was removed 10 days after the cranioplasty, and the general wound aspect was evaluated (Fig. 5). No other neurological deficits or complications were further identified.

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Fig. 5 Wound aspect and general evaluation (10 days after the surgery)

Clinical Case Follow-Up This clinical experience comprises the use of a relatively new material used in cranioplasty. PEEK already proved in usage in spine surgery, as a bone replacement which does not interfere with the components of the nervous system components. Moreover, this material is already used as a material for cranial bone repair with good outcomes which were also observed in this case. Long-term follow-up will be performed for evaluating material’s degradation and permanently evaluate the patient’s general condition.

Summary This chapter began with some general information regarding the use of biomaterials in neurosurgery and cranioplasty. Currently, a wide variety of biomaterials, especially biopolymers and bioceramics, are used or tested for the restoration of cranial bone defects or deformations, and the choice between them depends on multiple factors, like the size and location of the defect or if a tumor is involved in the pathology or not. Although the biomaterials involved in cranial bone restoration does not induce major influences upon the overall complications rate, the biomaterials choice represents one of the key factors that will lead to a successful surgical intervention. The current use of biomaterials for cranioplasty is divided based in the size of the bone defect: if a small-sized defect needs to be restored, cement may be used for this application so PMMA, hydroxyapatite, or other bioceramics may be used as a biomaterial. On the other hand, if a larger defect needs to be restored, an implant will be required at the implantation site so its manufacture will begin before the

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surgical intervention. Titanium is still used as an implant, along with PEEK, hydroxyapatite, or PMMA. However, each application has its specific requirements and the selected biomaterial must fulfill all of them, or at least the majority. The current research for an optimum biomaterial aims for developing a new materials generation, which will enable bone regeneration by stimulating specific host tissue response, and the porous bioceramics may be considered their precursors. Furthermore, the use of modern techniques for the biomaterial design and manufacturing will lead to a faster development, especially since their current use significantly improve the functionality and the aesthetic aspects of the implants designed for cranial reconstruction [39]. In the next step, the use of biocomposite materials will also improve a wide set of biomedical applications, due to the large number of properties that may be acquired through proper manufacturing control. Both the structure and the biological behavior may be improved through phases’ variation and filler distribution. However, the current use of composite materials is limited due to some limitations [89] regarding the lack of clinical and experimental results about the long-term materials’ behavior, the complex preparation techniques, and the limited specific standard test methods. If these critical aspects will be thoroughly addressed, a new step to composite commercialization may be made, because the perspectives regarding composite materials are optimistic and depend solely on the good collaboration between engineering medicine, chemistry, biology, and biomaterial science specialists [39].

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Marine Biomaterials as Drug Delivery System for Osteoporosis and Bone Tissue Regeneration Joshua Chou and Jia Hao

Contents Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Coral Exoskeletons . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Synthesizing Calcium Phosphates from Coral Exoskeletons . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biomimetic Calcium Phosphate Drug Delivery Systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Drug Loading, Coatings, and Characterization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biomimetic Drug Delivery Systems for Pharmaceutics . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biochemical Modifications of Biomimetic Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Biomimetic Scaffold Effect on Bone Mesenchymal Stem Cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Stem Cell Coating of Biomimetic Scaffold . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Applications of Marine Biomimetic Scaffolds . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Localized Treatment of Osteoporosis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Long-Term Systemic Treatment of Osteoporosis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Marine Biomimetic Scaffolds for Maxillofacial Bone Repairs . . . . . . . . . . . . . . . . . . . . . . . . . . . . Future Prospect and Developments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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There is currently an urgent need to develop sustainable and therapeutically relevant advanced drug delivery systems to treat the prevalence of ongoing human diseases and ailments. The effectiveness of such system will depend

J. Chou (*) Advanced Tissue Regeneration and Drug Delivery Group, University of Technology Sydney, Sydney, NSW, Australia e-mail: [email protected] J. Hao Oral Implantalogy and Regenerative Dental Medicine, Tokyo Medical and Dental University, Tokyo, Japan e-mail: [email protected] # Springer International Publishing Switzerland 2016 I.V. Antoniac (ed.), Handbook of Bioceramics and Biocomposites, DOI 10.1007/978-3-319-12460-5_57

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primarily on the properties and characteristics of the carrier material. In this chapter, marine materials are investigated as potential drug delivery carriers for bone tissue engineering and in the treatment of osteoporosis. This chapter will explore the unique structures of marine materials that set it apart from its synthetic counterparts and the conversion to biocompatible calcium phosphates combined with synthetic modifications, and case studies will demonstrate the potential clinical applications. Keywords

Bioceramic • Biomimetic • Calcium phosphates • Bone tissue engineering

Introduction The twenty-first century is regarded as the era of tissue engineering and regenerative medicine. The notion that human soft and hard tissues can both be regenerated or expedited was once a figment of imagination and science fiction. Advances in science and the merging of interdisciplinary fields have catapulted innovation and novel technologies with the goal of increasing human longevity. To achieve this, developments in novel material design with complex structures are required to mimic as closely to the human tissues. Moreover, the synthesis process must adhere to strict regulatory guidelines and economic implementation. The main driving force behind the push in regenerative medicine research is the diverse clinical applications of tissue engineering strategies and the specificity of these treatments. A key area of research in tissue regeneration is the development of both organic and inorganic scaffolds. These provide the basis to the fundamental development of direct and/or indirect stimulation of tissue regeneration. There is a continuous need to explore new avenues in which materials, cells, and biologically active molecules can be combined to deliver therapeutic responses. This is critical, as cells and growth factors are key elements for an effective regeneration process. Currently there exist clinically a vast range of biomaterials that have been extensively studied and shown to be capable of accommodating clinical needs. However, with a global rise in an aging population, more ailments and sickness are going to prevalent. Taking into account the rise in society’s middle class with the option of accessing better healthcare, there is a natural demand calling for improvements to current practices and the shortening of medical treatments which has both patient and socioeconomic implications. With global rise in healthcare costs, governments and institutions are keen on developing and supporting tissue engineering and regenerative medicine research. Current researches have been focused on improving existing scaffold materials by making it more bioactive, meaning increasing the material’s ability to stimulate a specific response. This approach has shown to be quite effective and it is a trend that continues to be followed. Among the diverse range of biomaterials available, calcium phosphates are one of the most attractive and extensively studied and

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clinically used bioceramic material. Calcium phosphates are generally used as a bone substitute material due to the compositional similarity to the original human bone, essentially allowing the material to be biocompatible within the body without inducing any inflammatory or host responses. As such, the majority of research using calcium phosphates center on its use in bone tissue regeneration. It should be noted that this is a general misconception, as biocompatible calcium phosphates can be adapted for treatments of other diseases and not just limited to bone tissue engineering. One of the key disadvantages of bioceramics is the brittle nature of the material which only allows its use for non-load-bearing applications. This adds another dimension to the challenge in synthesizing calcium phosphate material with complex and integrated architectural structures which is necessary to allow cells and key nutrients to infiltrate the material to promote and sustain a strong bond between the scaffold and its surrounding. This is particularly crucial for bone substitutes as irregular structures and pore sizes have shown to cause inflammatory response and scaffold rejection [1]. With the increase in knowledge and understanding between bioceramics and bone, it has become apparently clear that scaffolds need to possess structural characteristics similar to natural bone to initiate the required cellular response to promote bone remodeling. Ideally, scaffold materials should favor cellular attachment, growth, and differentiation by providing a highly porous, open-pored, and fully interconnected geometry for the cells. Microporosity with pores less than 10 μm is needed for capillary ingrowth and cell-matrix interactions. At the same time, macropores allows for nutrient supply and waste removal of cells. In addition, scaffolds should be resorbable so as to allow replacement by newly formed bone in the long term. It is also preferable that the external shape of the scaffold material can easily adapt to the defect size. While there are a number of bioceramic materials that address all these criteria and requirements, often their therapeutic effectiveness has limits. To explore an alternative approach and strategy, this chapter will examine a new class of calcium phosphate material derived from marine biomimetics. This chapter will describe in detail the rationale and interests in the use of natural marine-sourced materials as precursors for development of scaffolds for drug delivery systems (DDS).

Coral Exoskeletons The earth’s surface is covered by &70 % water and contains 80 % of all life found on the planet [2]. It is no wonder then that the ocean has been and still is a source of food, let alone a vast source of therapeutic molecules. The marine habitat houses a tremendous amount of varying living organisms of which many are still yet to be discovered due to technological limitations of exploring the ocean depth. However, with integrated advancement of submersible and the realization of exotic organisms containing therapeutic compounds, more species are being uncovered and studied. Nature in many ways can inform us on how to build structures, architectural designs, and the fabrication of new materials with exemplary performance using the most

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energy-efficient process with maximum optimization. To study natural materials can provide fresh and new approach for generating unique scaffolds for use in regenerative medicine with the potential to outperform conventional man-made advances. Most synthetic materials require some form of modification to meet the essential requirements of an adequate bone scaffold. This is however not the case for natural materials, in particularly corals. The coral life cycle begins with the polyps which absorbs the calcium ions and carbonic acid present in the seawater to produce the calcium carbonate in the form of aragonite crystals representing 97–99 % of the coral exoskeleton [3]. The remaining composition is made up of various elements and is dependent on the environment but mainly consists of trace elements of magnesium (0.05–0.2 %), strontium, fluorine, and phosphorous in the phosphate form (0.02–0.03 %) [4]. Coral polyps edify a centripetal exoskeleton upon which they reside, a process reminiscent of bone formation. During human bone remodeling, bone-building osteoblasts secrete collagen, which forms the framework of bone. Similarly, the outer layer of coral consists of calicoblast cells that secrete aragonite which calcifies to form the coral scaffold with an architecture characteristic to each species.

Synthesizing Calcium Phosphates from Coral Exoskeletons The initial testing of coral material displayed unsatisfactory results due to the structural instability once exposed to physiological environment. To address this, a new synthesis process was developed by Roy and Linnehan [5] in which hydrothermal conversion was capable of replacing the carbonate component with phosphate to produce calcium phosphate bioceramics base on the following chemical reaction (Eq. 1.):  10CaCO3 þ 6ðNH4 Þ2 HPO4 þ 2H2 O ! Ca10 ðPO4 Þ6 ðOHÞ2 þ 6ðNH4 Þ2 CO3 þ 4H2 CO3 (1) The general process generally involves placing the coral precursor material in phosphatic solution under 15000psi for 24–48 h at 220–250  C. The calcium/ phosphate (Ca/P) ratio can be altered to synthesize different forms of calcium phosphates including the more stable hydroxyapatite to the more degradable alpha tricalcium phosphate. The key benefit of the hydrothermal treatment is the simplicity of chemically modifying the compositional component of the material but, more importantly, the process is a replacement process in which the physical structures are not affected. This preservation of the unique porous characteristic of coral species is what makes this process attractive. The sintering temperature at 220  C is a lot lower compared with conventional synthesis of tricalcium phosphates (800–1100  C). This enables biomaterial scientist the option and flexibility to synthesize natural bioceramics depending on the clinical application. Calcium phosphate solubility is dependent on the Ca/P ratio and as such for drug delivery applications; one can predict the degradation rate for the intended purpose.

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For more fast-acting delivery systems, one might not necessarily need to convert the coral scaffold but as a calcium carbonate may be appropriate. However, for bone tissue engineering, beta-tricalcium phosphate (Ca/P = 1.5) has remained one of the most attractive material as a scaffold and as a drug delivery vehicle. Different drug delivery systems using coral exoskeletons as a precursor material have been developed and studied [6–8].

Biomimetic Calcium Phosphate Drug Delivery Systems The development of biomimetic calcium phosphate-based drug delivery systems have gained significant momentum in recent years as scientists are realizing the potentials in harvesting and utilizing natural products especially from the marine environment. Natural polymers such as chitin and chitosan are shown to be versatile functional materials that possess excellent biocompatibility, biodegradability, non-toxicity, and adsorption properties [9, 10]. While marine calcareous exoskeletons have been extensively shown and applied as a bone substitute material, little has been done on using it as a scaffold for drug delivery applications. A successful drug delivery system is characterized by the drug loading efficiency and, more importantly, the release rate of the drug. These are defining parameters that significantly impact on the therapeutic outcome. As previously discussed, the structural properties of foraminifera materials possess interconnected uniform porous chambers consisting of both macro- and micropores, which make them ideal candidate as drug carriers. The consistency in these structural properties allows for predictable and constant drug loading and release compared with nonuniform structural drug carriers. In addition, the hydrothermal treatment allows these materials to be biocompatible, and the degradation rate of different types of calcium phosphates can therefore be controlled and predicted. This section will discuss drug loading strategies and the applications of using foraminifera as a precursor material for drug delivery systems in bone tissue engineering.

Drug Loading, Coatings, and Characterization Biodegradable drug delivery systems have always being an attractive alternative to nondegradable materials or direct injection as they do not require removal after serving its purpose. Therefore calcium phosphate-based drug carriers are ideal for delivery pharmaceutics. The therapeutic efficiency of a drug delivery system will depend on the loading efficiency and that correlates with how the drugs are loaded initially. The most common and straightforward approach is by directly immersing the biomimetic material into a concentrated solution containing the desired pharmaceutic which will over time adsorb onto the scaffold. Depending on the type of pharmaceutic, this may be sufficient to load the drugs. Surface charges, chemical composition, and structural properties are all crucial defining factors. To

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increase the loading efficiency, in a previous study, the converted scaffolds were immersed in the concentrated drug solution in a rotary evaporator which allowed for deeper drug penetration into the material due to the vacuum pressure [7]. For a more long-term and sustainable drug release, coatings are required. Strategies and coating methods are extensively studied for calcium phosphate materials and can be adapted for these biomimetic scaffolds. It was shown in a previous study that apatite coating around the scaffold was able to achieve controlled release of the drug simvastatin, thereby generating improvement in the therapeutic efficacy of the overall drug delivery system [7]. Similarly, liposome coatings have also being demonstrated to be achievable with these biomimetic scaffolds [8]. Upon optimizing the drug loading efficiency, the release rate of the drug from the material is the next essential element. Generally once these biomimetic drug delivery systems are inserted into its targeted environment, drugs will start to elude initially through the surface, then through the pores as physiological fluids infiltrate, and, finally, through the degradation of the material itself. Different types of buffer solution are commonly used to “simulate” the likely physiological environments that the material will be expose to gain an understanding and overview on how the carrier material will interact and behave. These biomimetic scaffolds have been evaluated in simulated body fluids (SBF), in phosphate buffer saline (PBS), and in acetate buffer solutions [6–8, 11, 12]. It was found that in the initial stages between 0 and 24 h, burst release of the drugs from the surface and the pores are observed. This generally accounts for a very small percentage of the total incorporated drugs. To characterize these properties and determine the drug-eluding mechanism, a drug release plot based on the Higuchi equation can be made. The Higuchi equation is described in (Eq. 2): rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi  Di e  2Cd  eCs t M t ¼ A Cs τ

(2)

where Mt is the amount of drug release after time t, A is the matrix surface area, D is the diffusion coefficient of the drug, Cs is the solubility, Cd is the concentration of the drug in the matrix, τ is the tortuosity, and e is the porosity of the matrix. The release mechanism was demonstrated using a model drug incorporated with the biomimetic calcium phosphate material and plotting a subsequent drug release profile based on the Higuchi plot shown in Fig. 1. The Higuchi plot shows two phase of drug release: a surface release slope (a) and drug diffusion release from the matrix material slope (b). The initial burst release was due to conventional release of drug from the surface of the carrier material as illustrated by slope (a). The second-order release profile shows a linear slope (b) in the Higuchi plot suggesting that the release of the drug from the matrix material is based on a drug diffusion process through the macro- and micropores of carrier. This provides an overview on the release mechanism profile from using these carrier materials and can allow further optimization by controlling these release profiles.

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Fig. 1 Higuchi plot showing the initial burst release is through (a) the surface of the material followed by (b) diffusion mechanism through the matrix material

Biomimetic Drug Delivery Systems for Pharmaceutics Calcareous precursor materials are ideal candidates not only as scaffolds but also as carriers for pharmaceutics. These natural materials can be easily translated into clinically applicable drug delivery systems. As discussed in the previous section, biomimetic scaffolds can adsorb drugs and release them base on a two-stage process including an initial burst release followed by diffusion through the material by the actions of macro- and micropores. As a proof of principle, two types of drug delivery systems for bone tissue engineering will be discussed in this section.

Antibacterial and Bone-Stimulating Drug Delivery System Despite increase in surgical precautions and sterilization, today’s orthopedic surgeries still face the looming challenge of bacterial infection during bone repair. With reports of increase in bacterial resistance to antibiotics, clinicians are evermore precautious about the use of antibiotics. The reason for concern is during bone fracture repair procedures; the most common infection is from Staphylococcus aureus (S. aureus). Symptoms from S. aureus infection do not present itself till weeks after the surgery which by than would require a revision surgery to remove the infected implant. This directly impacts the patient’s recovery ability and adds to the already-overburdened healthcare system. This has motivated biomaterial scientists to develop scaffolds with antibiotics incorporated to prevent the initial colonization of S. aureus. Precaution must be taken to ensure that the antibiotic concentration released must be at a therapeutic level where the bacteria would not develop resistance to the antibiotic or reoccur. As such, efforts into the development of controlled release antibiotic drug delivery systems have gained significant traction.

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Biocompatible and biodegradable tricalcium phosphates derived from marine calcareous precursors were incorporated with gentamicin antibiotics (TCP-Gen) and tested against methicillin-resistant Staphylococcus aureus (MRSA) removal and prevention in vitro. TCP-Gen were introduced into MRSA culture during the exponential growth phase, and it was determined that within 20 mins the bacteria were completely eliminated and reoccurrence did not occur during the 48 h experimental observation. This suggests that the amount of gentamicin released during the initial burst release was sufficient to eliminate the actions of MRSA growth and more importantly prevent the bacteria from regrowth. It was calculated that the amount of gentamicin released only accounted for 20 % incorporated, and it is predicted that the remaining antibiotics will continue to be released as the material degrades with calcium phosphate ions. Further tests to examine the attachment of MRSA to the biomimetic scaffold also proved the effectiveness of the incorporated gentamicin. Adhered MRSA were completely removed and dysfunctional within 15 mins and reoccurrence were not observed during the 48 h experimental period. These results shows promising potential for the use of these biomimetic tricalcium phosphate scaffolds as controlled release delivery systems capable of sustaining therapeutic level release of gentamicin to combat MRSA. In addition to the incorporated antibiotics, bisphosphonate, which is commonly prescribed to stimulate bone regeneration, was also incorporated together with the gentamicin as a dual drug delivery system. In vitro studies showed that this type of dual system was able to stimulate osteoblast cell proliferation and do not induce any cell toxicity while preserving the underlying antibacterial properties [13]. The results obtained have encouraged further evaluation in in vivo rodent models to determine the biological efficacy of this dual drug delivery system. It is envisioned that this type of biomimetic delivery system could one day be clinically applied as a bone substitute material that can stimulate bone regeneration while preventing the occurrence of bacterial infections.

Controlled Drug Delivery System for Osteoporosis Osteoporosis treatment is at the center of research focus globally with the rise of the middle class population and the significant increase in aging populace. While currently there does not exist any pharmaceutics to completely reverse the osteoporotic condition, drug treatments are able to manage and improve the conditions of patients with a good degree of recovery. As a systemic-wide disease within the body, a therapeutic approach would require a sustained long-term infusion of pharmaceutics within the body to effectively restore the balance of bone remodeling. Most of the drugs shown therapeutic effectiveness all possesses some form of stimulation to osteoblasts and inhibition of osteoclasts. The family of bisphosphonate and its many derivatives have reached a peak in its therapeutic treatments, and signs of side effects associated with long-term use of bisphosphonates are surfacing with mixed response from clinicians. In the past few years, simvastatin has shown in studies to be a strong contender as an alternative medication to treat osteoporotic patients. Developed initially as a cholesterollowering agent, the drug was later found to affect positively on bone remodeling cells. This has generated various forms of calcium phosphate-based materials

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incorporated with simvastatin to be applied for bone tissue engineering [14]. One of the key challenges with the use of simvastatin clinically is the effect of high dosage concentration inducing severe side effects including muscle inflammation and deterioration. Therefore any formulae involving simvastatin would require therapeutic optimization to ensure the concentration does not exceed the therapeutic range. This requires a high degree of control and release mechanism to be effective. Biomimetic tricalcium phosphates were incorporated with simvastatin at a concentration described in literature to be therapeutically effective. To control the release of simvastatin and to ensure a sustained and long-term release of the drug, an additional apatite layer was coated around the scaffold material to preserve the release of the drug and limit the initial burst release. This system was tested in vitro by implanting the samples intramuscularly near the femur bone of osteoporotic mice. The bone mineral content (BMC) and bone mineral density (BMD) of the femur bones on both sides examined by CT analysis showed a much higher statistical increase in the bone BMC and BMD from the apatite-coated scaffolds compared with no coating and with direct injection of simvastatin. It was also observed by direct injection that the mice suffered from severe muscle deterioration and inflammation as a result of having high localized concentration of the drug [7]. This again signifies the importance of controlled release of pharmaceutics to limit the immunological host response. This demonstrates that controlled release can be achieved with the use of biomimetic scaffolds, and further optimizations can increase the therapeutic efficacy of such systems to become clinically applicable.

Biochemical Modifications of Biomimetic Scaffolds Among many strategies into the development of more bioactive scaffolds, significant interest has been vested on the incorporation of inorganic dopants such as magnesium and strontium. These dopants have shown to possess unique properties in stimulating osteoblast proliferation, inhibiting the actions of osteoclast resorption and increasing the mechanical properties of newly formed bone [15–18]. While many studies have shown the potentials for the use of these inorganic compounds combined with scaffolds for different bone tissue applications, an alternative key biological element has been vastly overlooked in comparison. Zinc, an essential element responsible for the regulation and metabolism of cells, has shown in studies to also possess the ability to stimulate osteoblast bone formation when synthesized with tricalcium phosphates [19–23]. Different types and formulations of zinc tricalcium phosphate (Zn-TCP) have been developed in the form of injectable powders in osteoporotic rats [21, 24] and injectable nanoparticles in jawbone [25] and in rabbit femora with over 50 % newly formed bone [26]. Adding on to this, more recent studies have established a relationship between zinc deficiencies in osteoporotic patients in particularly in the elderlies [27]. As tricalcium phosphate bioceramics are resorbable with a higher solubility rate compared with hydroxyapatite, the degree of degradation of Zn-TCP materials becomes crucial. If the Zn-TCP degrades too rapidly, this may result in high levels

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of zinc concentration that lead to severe cytotoxic effects [20–22]. As such, the concentration levels and the release of zinc ions are the two main factors in developing therapeutically effective Zn-TCPs. To date, studies conducted have shown that the optimal zinc concentration level is around 5 % [28] and most studies have based their formulation around this concentration. Another key challenge with the synthesis of Zn-TCP ceramics is the high temperature sintering required which is generally between 800  C and 1100  C. One of the key attracting factors of using marine exoskeletons as precursor materials for synthesizing bioceramics is that they naturally possess inorganic compounds within their chemistry as part of their natural calcification process. Interestingly, both strontium and magnesium are naturally part of the material, and in vitro studies have shown to stimulate osteoblasts and the inhibition of macrophages [12]. Natural materials are designed over time to become more efficient and optimal and are open to scientific discoveries and modifications to be adapted or further improve for human applications. With this in mind, the synthesized biomimetic TCP material was further modified by incorporating zinc within the crystallographic structure of the material. Just as phosphate replaced the carbonate component of the original precursor material by hydrothermal exchange, the synthesized TCP underwent another hydrothermal process in which the solution contained zinc. These zinc ions would replace the calcium ions within the material but only to a small degree. Initial findings showed that 0.5 % zinc incorporation was achieved [6]. The concentration of zinc incorporated is dependent on the zinc solution concentration and the time of the hydrothermal treatment. Micro-CT cross-sectional images shown in Fig. 2 reveal the intricate internal porous architectural structure of the foraminifera material and the preservation of the integrity after conversion to Zn-TCP. One can start to appreciate the elegance of natural scaffold materials with its elaborate structures combined with macro- and micropores.

Fig. 2 Micro-CT images showing the internal porous network of foraminifera precursor material and the preservation of the integrity of the scaffold after hydrothermal conversion to Zn-TCP

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Zinc release studies from the Zn-TCP showed no release in physiological buffer solutions over 7 days. This is indicative on the stability of the zinc within the material. When exposed to an acidic buffer environment (pH 4.5), zinc release was observed, but over 7 days this only accounted for 0.2 % of the total zinc incorporated. These results demonstrate the ability to synthetically modify biomimetic scaffolds by incorporating inorganic ions to change the overall properties of the material. This strategy to combine natural materials with scientific alterations for specific applications can be potentially applied to a vast range of materials with real-life clinical implications.

Biomimetic Scaffold Effect on Bone Mesenchymal Stem Cells The effect of biomaterial scaffolds on cellular response in the initial stages of characterization provides a fundamental understanding and insight into “how” the material will behave in a more complex biological environment. For stem cell-based bone tissue engineering, the material substrate should guide the initial anchorage of cells and their proliferation and assist in stimulating the cells’ differentiation into osteoblasts. The commitment of BMSCs to differentiate into an osteogenic lineage is highly dependent on the physicochemical properties of the surrounding environment and the scaffold on which they are cultured. Surface composition, roughness, and topography all contribute to the osteogenic process. The general focus is to promote an increase in osteoblast cell to stimulate a higher rate of bone regeneration. Other key cells including osteoclasts, osteoblasts, macrophages, etc. are all readily available and are commonly used in in vitro characterization. In a previous study, the osteogenic differentiation of primary BMSC when grown in the presence of biomimetic β-TCP and Zn-TCP was investigated [11]. The first in a series of tests is the biocompatibility of the material. The results from this study showed that both materials were able to sustain a good cell viability of &95 % during the 14-day period. This is reflective of the biocompatibility nature of calcium phosphate materials that is nontoxic to the cells. Examining the cell number, it was found that by 10 day, Zn-TCP exhibits a statistical number in cell growth compared with its counterpart and control. This was also observed at 14 day. Considering that zinc plays a crucial role in cell metabolism and proliferation [29], these results would suggest that the action of the zinc ions being released is having an impact on the proliferation of the BMSC. This would be potentially advantageous in an in vivo environment where the Zn-TCP could essentially stimulate more BMSC growth, thereby increasing the amount of differentiated osteoblasts. Stein and Lian [30], in their osteoblastic differentiation model, reported that MSCs generally proliferate for up to 7–14 days before secreting extracellular matrix proteins that produce early differentiation markers, such as alkaline phosphatase (ALP), which can be observed as early as 7 day. Osteogenic differentiation of osteoblast-like cells is one of the key steps in determining the success of bone and material integration. Among these, ALP activity is one of the widely recognized

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biomarkers in determining osteogenic differentiation, which is regulated by phosphate metabolism [31]. ALP is associated with osteoblast differentiation, and, as such, the level of ALP gives an indication of the stage of osteoblast differentiation. Interestingly at 10 day, ALP levels in Zn-TCP were statistically higher compared with the cell control and β-TCP. By 14 day, ALP level dropped which is an indication that differentiation has completed or near completion. The rate at which osteogenic differentiation occurs is crucial to increase the integration between the scaffold material and newly formed bone. Thirdly, observation of calcium mineralization, which is reflective of the maturation of osteoblasts after differentiation, distinctly shows significantly higher levels of mineralized nodules compared with β-TCP and control group. This again supports the crucial role and effect of zinc on BMSC differentiation which would impact the bone-material integration in an in vivo environment. While synthetic derivatives of Zn-TCP have also shown similar biological response from human bone marrowderived mesenchymal stem cells [29, 32], it should be emphasis the unique structural characteristic of the biomimetic material that is currently being characterized and, more importantly, the release of zinc ions are at a rate capable of sustaining BMSC differentiation into osteoblasts.

Stem Cell Coating of Biomimetic Scaffold One of the key contributing factors that have excited research in tissue engineering is the versatility of stem cells and their ability to differentiate to different cell types. In the context of developing more bioactive scaffolds/biomaterials, the incorporation of stem cells with these materials have therefore being extensively studied and applied [33, 34]. Many strategies and protocols developed are generally centered on isolating host patient cells and either expanded ex vivo than seeded directly onto the scaffold construct before placing the material back into the patient. The success of such construct is highly dependent on the efficiency of the seeding, the rate of cell adhesion to the scaffold, the viability of the cells once seeded, and the distribution of cells on the material [35]. Cell-material interaction is a key factor in the initial stages of cell attachment, and the morphology and architectural structure of the scaffold play a crucial role in determining the cell attachment rate and viability. Studies have shown, for example, that smaller scaffold pore diameters ( 30 μm) enhance cell-seeding efficiency compared with larger pore sizes ( 70 μm) [35, 36]. Since stem cells are anchorage dependent, the faster the cells are able to attach to the scaffold, the likelihood of the cells to remain viable is much higher. One of the simplest approaches for cell seeding onto scaffold/biomaterials is by directly seeding the cells on the construct and allowing the cells to grow and spread throughout the construct over time in an undisturbed environment [37]. A more dynamic approach would include a setup similar to bioreactors in which the seeded scaffold is exposed to a constant flow of media allowing the oscillatory pressure to spread the cells through the material [38]. While these methods have produced satisfactory results, they are however time consuming and often require a high cell

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Fig. 3 Examination of the stem cells distribution throughout the biomimetic scaffold, after different days post-seeding.

concentration for seeding as only a percentage of the cells completely adhere to the material and some loss in cell viability. Therefore there exists a need for an alternative seeding process and to determine if biomimetic scaffolds can provide the framework for stem cell attachment. In this study, a general laboratory centrifuge was used to “spin coat” the stem cells onto the biomimetic scaffold. This approach would provide a quick and easy method for coating other biomaterials and can be easily accessible in laboratories. During the optimization process, it was found that 1-min centrifugation seeded higher numbers of cells onto the scaffold compared with 2- and 3-min centrifugation, which resulted in significantly lower numbers of adhered cells. It is believed that by increasing the centrifugation time, the cells are being pushed off the scaffold. It was also found that the rotational speed of the centrifuge set at 700 g gave the most optimal amount of cell adherence. It was found that the cells seeded were 95 % viable at 7 days postseeding. Figure 3 shows that at 3 days post-seeding, the stem cells were well distributed throughout the material and spreading of the cells could be observed. One can say that the seeded cells were uniformly distributed on the surface of the scaffold. At 7-day post-seeding, it is apparent that the stem cells have spread and covered the vast

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majority of the scaffold material. From these observations one can easily control the level of confluency of cell spreading on the scaffold material depending on the intended application. It can therefore be said that the biomimetic scaffold material not only can sustain cell adherence but at the same time it is capable of allowing and possibly promote cell spreading on the material. This is encouraging and possesses obvious advantages in bone tissue engineering. Future studies will examine the in vivo response of these stem cell-coated biomimetic scaffolds to determine the optimal cell concentration.

Applications of Marine Biomimetic Scaffolds Localized Treatment of Osteoporosis Osteoporosis is a serious and silent musculoskeletal condition in which the normal bone remodeling process is disrupted with bone-destroying osteoclasts taking dominance over osteoblasts, thereby decreasing significantly the mechanical strength of the bone. Intense international research into the development of osteoporosis treatment has so far remained limited and focuses primarily on the management of the condition rather than having a “cure.” It is generally agreed that any kind of treatment for osteoporosis would require a multi-targeted systemic approach, likely in the form of multifunctional drug delivery systems. As previously discussed, foraminifera exoskeletons can be hydrothermally converted to contain the addition of bioactive zinc ions on top of the alreadypresent strontium and magnesium. The aim of this study was to demonstrate the use of these converted marine scaffold templates with multi-stimulatory ion doped as a drug delivery system [6]. As the material degrades in the local environment, the key dopants would be released into the host’s blood stream to provide a systemicwide distribution to target the osteoporosis condition. The in vivo study was examined in an osteoporotic mice model by ovariectomy. The Zn-TCP samples were implanted intramuscularly just above the femur bone (Fig. 4) which possess abundant systemic blood flow to the local femur bone and throughout the host. It should be noted this is an in vivo experimental model to observe the local and systemic effect of drug delivery systems and does not reflect on the actual clinical application. It is envisioned that Zn-TCP would be clinically applied by inserting the material into the medulla cavity where there is minimal obstruction. In addition, the osteoporotic condition affects the overall health condition and survivability of the mice, and as such this model was only used to observe the short-term (4 weeks) localized effect of the treatment. Micro-CT images (Fig. 5) also displayed a higher amount of cancellous bone in the + Zn group compared with the other experimental group and closely match those in the normal group. These results are an indication that Zn-TCPs have the potential of reversing the osteoporotic condition on the localized bone within a 4-week period compared with –Zn and the other control groups.

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Fig. 4 Radiological image of the Zn-TCP samples implanted intramuscularly just above the femur bone

Fig. 5 Examination of the bone mineral content growth after 4 weeks using Micro-CT show cortical bone to be significantly increased by the presence of Zn-TCP compared with the other experimental group

The results showed that at 4 weeks, the cortical bone of the localized femur bone (Fig. 6) was significantly higher in both the + Zn and –Zn experimental group and the density remained at similar levels. Interestingly, the cancellous bone content in the + Zn group showed tremendous growth (22 %) compared with negative growth in the other experimental groups consistent with bone loss due to the osteoporosis condition.

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Fig. 6 Examination of the bone mineral content growth after 4 weeks show (a) cortical bone to be significantly increased by the presence of Zn-TCP compared with OVX (ctrl), and (b) cancellous bone was also significantly stimulated by Zn-TCP compared with negative growth by OVX (ctrl). Asterisk sign represents p