Handbook of ionic substituted hydroxyapatites 9780081028346, 0081028342

Table of contents :
Content: 1. Structure of Biological Apatite: Bone and Tooth2. Analytical Tools for Hydroxyapatite3. Bioceramics: Types and Clinical Applications4. Basics of Hydroxyapatite: Structure, Synthesis, Properties and Clinical Applications5. Role of Substitutions in Bioceramics6. Carbonated Substituted Hydroxyapatite7. Fluoride Substituted Hydroxyapatite8. Magnesium Substituted Hydroxyapatite9. Zinc Substituted Hydroxyapatite10. Silver Substituted Hydroxyapatite11. Iron Substituted Hydroxyapatite12. Silicon Substituted Hydroxyapatite13. Strontium Substituted Hydroxyapatite14. Coating of Hydroxyapatite and Substituted Apatite on Dental & Orthopedic Implants15. 3-D Printing of Hydroxyapatite16. Hydroxyapatite and Tissue Engineering

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Handbook of Ionic Substituted Hydroxyapatites

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Woodhead Publishing Series in Biomaterials

Handbook of Ionic Substituted Hydroxyapatites Edited by

Abdul Samad Khan Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Eastern Province, Dammam, Saudi Arabia

Aqif Anwar Chaudhry Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

Woodhead Publishing is an imprint of Elsevier The Officers’ Mess Business Centre, Royston Road, Duxford, CB22 4QH, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, OX5 1GB, United Kingdom Copyright © 2020 Elsevier Ltd. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-08-102834-6 For information on all Woodhead Publishing publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisition Editor: Sabrina Webber Editorial Project Manager: Ana Claudia Garcia Production Project Manager: Debasish Ghosh Cover Designer: Ana Claudia Garcia Typeset by TNQ Technologies

Contents

Contributors About the editors

1

2

Structure of biological apatite: bone and tooth Ahmed Talal, Shorouq Khalid Hamid, Maria Khan and Abdul Samad Khan 1.1 Introduction 1.2 Structure of bone 1.2.1 Anatomy of bone 1.2.2 Composition of bones 1.2.3 Properties of bone structure 1.3 Structure of tooth 1.3.1 Enamel 1.3.2 Dentin 1.3.3 Pulp 1.3.4 Cementum 1.3.5 Properties of tooth structure 1.4 Conclusion References Analytical tools for substituted hydroxyapatite Mariam Raza, Saba Zahid and Anila Asif Introduction 2.1 Structure of hydroxyapatite 2.2 Fourier transform infrared spectroscopy 2.2.1 Silicon substitution 2.2.2 Strontium substitution 2.2.3 Magnesium substitution 2.2.4 Zinc substitution 2.2.5 Fluoride substitution 2.2.6 Iron substitution 2.2.7 Silver substitution 2.2.8 Carbonate substitution

xv xix

1

1 3 4 4 5

8 8 9 11 12 13

16 16 21 21 21 22 22 23 24 24 25 26 27 27

vi

Contents

2.3

2.4

2.5

2.6

2.7 2.8 2.9 2.10

3

Scanning electron microscopy 2.3.1 Silicon substitution 2.3.2 Magnesium substitution 2.3.3 Zinc substitution 2.3.4 Iron substitution X-ray diffraction analyses 2.4.1 Silicon substitution 2.4.2 Strontium substitution 2.4.3 Magnesium substitution 2.4.4 Carbonate substitution 2.4.5 Zinc substitution 2.4.6 Iron substitution 2.4.7 Silver substitution 2.4.8 Multielemental incorporation analysis Differential thermal analysis/thermogravimetric analysis 2.5.1 Magnesium substitution 2.5.2 Carbonate substitution 2.5.3 Multielement substitution 2.5.4 Zinc substitution 2.5.5 Zr-Ce cosubstitution 2.5.6 Silver substitution Raman spectroscopy 2.6.1 Zinc substitution 2.6.2 Silver substitution 2.6.3 Magnesium substitution 2.6.4 Strontium substitution 2.6.5 Silicon substitution 2.6.6 Fluoride substitution 2.6.7 Iron substitution Nuclear magnetic resonance In vivo/in vitro analysis X-ray fluorescence Conclusion References

Bioceramics: types and clinical applications Hashmat Gul, Maria Khan and Abdul Samad Khan 3.1 Introduction to bioceramics 3.2 Classification of bioceramics 3.2.1 Classification on basis of origin 3.2.2 Classification on basis of type of tissue response 3.2.3 Classification on basis of composition 3.2.4 Classification on basis of crystallinity

27 27 28 28 28

31 31 32 32 32 33 33 33 34

35 35 35 36 36 36 37

37 37 37 38 38 38 39 39

39 42 44 44 44 53 53 53 53 55 55 61

Contents

3.3

4

5

vii

Biomedical applications of bioceramics 3.3.1 Orthopedic applications 3.3.2 Coatings for chemical bonding 3.3.3 Bone tissue engineering 3.3.4 Dental applications 3.3.5 Ocular prosthesis 3.3.6 Otolaryngologic applications References

Basics of hydroxyapatitedstructure, synthesis, properties, and clinical applications Hamad Khalid and Aqif Anwar Chaudhry 4.1 Biological apatite and synthetic hydroxyapatite: differences and similarities 4.1.1 Wet-chemical methods 4.1.2 Solid-state methods 4.1.3 Hydroxyapatite coatings 4.2 Properties of hydroxyapatite 4.2.1 Structural insights 4.2.2 Physical and thermal properties 4.2.3 Mechanical properties 4.2.4 Biological performance 4.2.5 Applications of hydroxyapatite References Role of substitution in bioceramics Sobia Tabassum 5.1 Introduction 5.2 Bioapatites 5.3 Synthetic hydroxyapatite 5.4 Effect of substitution on charge and size of hydroxyapatite crystals 5.5 Types of metallic substituents 5.5.1 Monovalent cationic substituents 5.5.2 Bivalent cationic substituents 5.5.3 Trivalent cationic substituents 5.6 Nonmetallic substitutions 5.6.1 Fluoride substitution 5.6.2 Chloride substitution 5.6.3 Carbonate substitution 5.6.4 Silicon substitution 5.6.5 Boron substitution 5.6.6 Sulfate substitution 5.6.7 Selenium substitution 5.6.8 Tellurium substitution 5.7 Significance of multiple substitutions

61 62 64 66 67 71 72

73

85

85 87 94 94

97 97 98 98 98 101

107 117 117 117 118 119 121 121 123 127

129 130 130 130 131 131 132 132 133

133

viii

Contents

5.8 5.9

6

7

8

Grafting of organic compounds/polymers on the surface of hydroxyapatite Conclusion and outlook References

133 134 135

Carbonate substituted hydroxyapatite Saadat Anwar Siddiqi and Usaid Azhar 6.1 Introduction 6.1.1 Hydroxyapatite 6.1.2 Substitution of hydroxyapatite 6.2 Carbonate substituted hydroxyapatite 6.2.1 A-type carbonated hydroxyapatite 6.2.2 B-type carbonated hydroxyapatite 6.2.3 Type-AB carbonated hydroxyapatite 6.3 Methods of synthesis 6.3.1 Precipitation technique 6.3.2 Hydrothermal technique 6.4 Characterization techniques 6.4.1 X-ray diffraction 6.4.2 Fourier-transform infrared spectroscopy 6.5 Carbonated hydroxyapatite as a coating material 6.6 Carbonated hydroxyapatiteebased composite materials 6.7 Biological studies 6.8 Conclusion References

149

Fluoride-substituted hydroxyapatite Sandleen Feroz and Abdul Samad Khan 7.1 Structure of hydroxyapatite 7.2 Presence of ions in biological apatite 7.3 Ionic substitution of hydroxyapatite 7.4 Fluoride substitution in hydroxyapatite 7.4.1 Method of preparation of fluoride-substituted hydroxyapatite 7.4.2 Structure of fluoride-substituted hydroxyapatite 7.4.3 Biomedical applications of fluoride-substituted hydroxyapatite 7.5 Concluding remarks References Further reading

175

Magnesium-substituted hydroxyapatite Ume Omema, Hamad Khalid and Aqif Anwar Chaudhry 8.1 Biological importance of the magnesium ion and its relevance to calcium phosphates 8.2 Synthesis of magnesium-substituted hydroxyapatite 8.2.1 Coprecipitation method

197

149 150 152

152 153 154 156

157 157 158

159 159 160

162 164 165 166 167

175 176 177 178 179 182 187

189 190 196

197 198 198

Contents

ix

8.3 8.4

8.5

9

10

8.2.2 Solegel method 8.2.3 Batch and flow hydrothermal synthesis 8.2.4 Solid-state methods Magnesium-substituted hydroxyapatite coatings Characterization of magnesium-substituted hydroxyapatite 8.4.1 Electron microscopy 8.4.2 Surface area 8.4.3 X-ray diffraction 8.4.4 Fourier-transform infrared spectroscopy 8.4.5 Raman spectroscopy 8.4.6 Thermogravimetric analysis Assessing biological response to magnesium-substituted hydroxyapatite 8.5.1 In vitro analysis 8.5.2 In vivo analysis (animal model) References

199 199 200

201 203 203 205 206 206 208 209

210 211 212

213

Zinc-substituted hydroxyapatite Kashif Ijaz, Hamad Khalid and Aqif Anwar Chaudhry 9.1 The biological importance of zinc ion and its relevance to calcium phosphates 9.2 Synthesis of zinc-substituted hydroxyapatite 9.2.1 Solegel methodology 9.2.2 Coprecipitation method 9.2.3 Hydrothermal synthesis 9.2.4 Solid-state methods 9.3 Zinc-substituted hydroxyapatite coatings 9.4 Characterization of zinc-substituted hydroxyapatite 9.4.1 Electron microscopy 9.4.2 Surface area analysis 9.4.3 X-ray diffraction 9.4.4 Fourier-transform infrared spectroscopy 9.4.5 Raman spectroscopy 9.4.6 Thermogravimetric analysis 9.5 Biological performance of zinc-substituted hydroxyapatite 9.5.1 Cell response 9.5.2 Antibacterial response 9.5.3 Animal model References

217

Silver-substituted hydroxyapatite Zohaib Khurshid, Muhammad Sohail Zafar, Shehriar Hussain, Amber Fareed, Safiyya Yousaf and Farshid Sefat 10.1 Introduction 10.2 The rationale of silver in apatite

237

217 218 218 219 220 220

220 221 221 224 225 226 227 228

228 228 229 230

233

237 237

x

Contents

10.3 10.4

10.5 10.6 10.7 10.8 10.9 10.10 10.11 10.12 10.13

11

Substitution in hydroxyapatite Methods of Preparations 10.4.1 Wet precipitation 10.4.2 Coprecipitation 10.4.3 Hydrothermal 10.4.4 Solegel 10.4.5 Microwave 10.4.6 Other methods Use of surfactants during preparation Structure of silver-substituted hydroxyapatite Effect on bioactivity of hydroxyapatite Use of micro- and/or nanosilver particles in hydroxyapatite for implant coatings Antibacterial effect of silver-substituted hydroxyapatite Drug-loaded silver-substituted hydroxyapatite Other biomedical applications Dental applications Conclusion References

Iron-substituted hydroxyapatite Christie Lung Ying Kei 11.1 Introduction 11.2 Biological importance of iron 11.3 Synthesis methods 11.3.1 Wet-state methods 11.3.2 Solid-state methods 11.4 Lattice structure of iron-substituted hydroxyapatite 11.5 Physical properties of iron-substituted hydroxyapatite 11.5.1 Magnetization 11.5.2 Mechanical properties 11.6 Biological properties of iron-substituted hydroxyapatite 11.6.1 In vitro bioactivity 11.6.2 Biocompatibility 11.7 Biomedical applications 11.7.1 Drug delivery 11.7.2 Biosensor 11.7.3 Protein and gene delivery 11.7.4 Magnetic resonance imaging 11.7.5 Scaffolds 11.7.6 Hyperthermia therapy 11.8 Conclusions References

239 241 242 242 242 243 243 244

244 245 245 246 247 248 248 249 249 250 259 259 260 261 261 267

268 268 269 269

270 270 270

273 274 274 275 275 276 277

279 279

Contents

12

13

14

Silicon-substituted hydroxyapatite Aysha Arshad and Ather Farooq Khan 12.1 Introduction 12.2 Biological importance of silicon 12.3 Silicon-substituted hydroxyapatite synthesis methods 12.3.1 Precipitation method 12.3.2 Solegel method 12.3.3 Hydrothermal method 12.3.4 Solid-state reaction 12.4 Characterization of silicon-substituted hydroxyapatite 12.4.1 X-ray diffraction 12.4.2 Fourier-transform infrared spectroscopy 12.4.3 Other characterization studies 12.5 Silicon-substituted hydroxyapatite in coatings 12.6 Silicon-substituted hydroxyapatite in biomedical applications 12.6.1 In vitro studies 12.6.2 In vivo studies References Effects of strontium substitution in synthetic apatites for biomedical applications Nujood Ibrahim Alyousef, Yara Khalid Almaimouni, Mashael Abdullah Benrahed, Abdul Samad Khan and Saroash Shahid 13.1 Introduction 13.2 Method of preparation of strontium-substituted hydroxyapatite 13.2.1 Hydrothermal method 13.2.2 Solegel method 13.2.3 Coprecipitation method 13.2.4 Wet chemical synthesis 13.3 Structure of strontium-substituted hydroxyapatite 13.3.1 Crystallographic analysis of strontium-substituted apatite 13.3.2 Structural analysis of strontium-substituted apatite 13.4 Biomedical applications of strontium-substituted hydroxyapatite 13.4.1 Orthopedic and bone tissue regeneration 13.4.2 Implant coating 13.4.3 Osteoporosis treatment 13.4.4 Enamel repair 13.4.5 Guided bone regeneration 13.4.6 Drug carrier References Coating of hydroxyapatite and substituted apatite on dental and orthopedic implants Farasat Iqbal and Hira Fatima 14.1 Introduction

xi

283 283 284 285 285 288 288 288

289 291 292 292

293 295 296 297

299

307

307 307 308 309 310 311

312 312 313

316 316 318 319 320 320 321

322

327 327

xii

Contents

14.2 14.3 14.4

Hydroxyapatite coatings for dental and orthopedic implants Processing of hydroxyapatite and substituted apatite coatings Techniques for hydroxyapatite-based coating onto metallic implant 14.4.1 Solegel dip-coating technique 14.4.2 Electrochemical deposition 14.4.3 Plasma spraying technique 14.4.4 High-velocity suspension plasma spraying 14.4.5 Biomimetic coatings Host tissue interaction with hydroxyapatite coatings Current challenges and future opportunities References

330 332 336

Three-dimensional printing of hydroxyapatite Asma Tufail, Franziska Schmidt and Muhammad Maqbool 15.1 Overview of additive manufacturing 15.1.1 The historical development and need for additive manufacturing 15.1.2 Biomedical applications of additive manufacturing 15.2 General introduction of different additive manufacturing techniques 15.2.1 Vat photopolymerization 15.2.2 Material jetting 15.2.3 Binder jetting 15.2.4 Material extrusion 15.2.5 Powder bed fusion 15.2.6 Sheet lamination 15.2.7 Directed energy deposition 15.3 Additive manufacturing of ceramics and ceramic composites 15.3.1 Powder-based ceramic additive manufacturing 15.3.2 Slurry-based ceramic additive manufacturing 15.3.3 Solid-based ceramic additive manufacturing 15.4 Additive manufacturing of hydroxyapatite and composites 15.4.1 Introduction and biological applications of three-dimensional

355

14.5 14.6

15

15.4.2 15.5

16

hydroxyapatite for biomedical applications Synthesis of three-dimensional hydroxyapatite by different additive manufacturing techniques

Concluding remarks and the future of additive manufacturing of hydroxyapatite References

Hydroxyapatite and tissue engineering Saeed Ur Rahman 16.1 Origin and history of hydroxyapatite 16.2 Tissue engineering 16.3 Main components of tissue engineering 16.3.1 Cells 16.3.2 Growth factors 16.3.3 Scaffolds acting as extracellular matrix

336 337 338 338 339

340 342 344

355 355 357

357 357 358 359 360 360 361 362

362 363 366 368

369 369 370

375 376 383 383 383 384 385 385 386

Contents

16.4 16.5

16.6

16.7

Index

xiii

Scaffold materials Hydroxyapatite as tissue scaffolding material 16.5.1 Hydroxyapatite in combination with natural materials 16.5.2 Hydroxyapatite in combination with synthetic materials Role of hydroxyapatite in tissue engineering 16.6.1 Bone 16.6.2 Periodontal tissue regeneration 16.6.3 Temporomandibular joint 16.6.4 Cartilage 16.6.5 Dentin 16.6.6 Cementum Future horizons References

386 387 387 389

390 390 393 393 394 394 395

396 396 401

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Contributors

Yara Khalid Almaimouni College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Nujood Ibrahim Alyousef College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Aysha Arshad Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Anila Asif Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Usaid Azhar Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Mashael Abdullah Benrahed College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Aqif Anwar Chaudhry Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Amber Fareed Oman

Department of Preventive Dentistry, Oman Dental College, Muscat,

Hira Fatima Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Sandleen Feroz Department of Dental Materials, Foundation University Islamabad Campus, Islamabad, Pakistan Hashmat Gul Department of Dental Materials, Army Medical College, National University of Medical Sciences, Islamabad, Pakistan Shorouq Khalid Hamid Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Shehriar Hussain Department of Dental Materials, College of Dentistry, Jinnah Sindh Medical University, Karachi, Pakistan Kashif Ijaz Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

xvi

Contributors

Farasat Iqbal Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Hamad Khalid Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Abdul Samad Khan Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Ather Farooq Khan Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Maria Khan Pakistan

Department of Oral Biology, University of Health Sciences, Lahore,

Zohaib Khurshid Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Saudi Arabia Muhammad Maqbool Institute of Biomaterials, Department of Materials Science and Engineering, University of Erlangen-Nuremberg, Erlangen, Germany Ume Omema Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Saeed Ur Rahman Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Mariam Raza Science of Dental Materials, PMC, Dental Institute, Faisalabad Medical University, Faisalabad, Pakistan Franziska Schmidt Department of Ceramic Materials, Faculty III Process Sciences, Institute of Materials Science and Technology, Technical University Berlin, Berlin, Germany Farshid Sefat Interdisciplinary Research Center in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom Saroash Shahid Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom Saadat Anwar Siddiqi Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Sobia Tabassum Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan Ahmed Talal Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia Asma Tufail Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

Contributors

xvii

Christie Lung Ying Kei Dental Materials Science, Applied Oral Science, Faculty of Dentistry, The University of Hong Kong, Hong Kong, 5/F, Prince Philip Dental Hospital, Sai Ying Pun, Hong Kong SAR, P.R.China Safiyya Yousaf Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom Muhammad Sohail Zafar Department of Restorative Dentistry, College of Dentistry, Taibah University, Saudi Arabia Saba Zahid Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

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About the editors

Dr. Abdul Samad Khan is currently working as Associate Professor Dental Biomaterial at Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University (Formerly, University of Dammam), and also affiliated with Interdisciplinary Research Centre in Biomedical Materials, COMSATS Institute of Information Technology. Dr. Khan did BDS in 1999 from Pakistan, did MSc in 2005, Dental Materials from Queen Mary University of London (QMUL), UK, and completed PhD in 2009 from QMUL, UK. During this time, he also worked as Research Assistant with Doxa AB, Sweden. Later, in 2013e14, he did Postdoctoral Associateship from University of Sheffield, UK. Dr. Khan achieved research grants from International and National Organizations. His area of expertise is synthesis and characterization of bioceramics and bioactive composites. He regularly published papers on apatites, substituted apatites, and their characterizations. Dr. Aqif Anwar Chaudhry is currently working as Associate Professor Biomaterials and Head Interdisciplinary Research Centre in Biomedical Materials, COMSATS Institute of Information Technology, Lahore, Pakistan. Dr. Chaudhry did his PhD in 2008 from Queen Mary University of London, UK. His area of expertise is synthesis and characterization of nanohydroxyapatite and substituted hydroxyapatite. He produced multiple papers on these topics which were published in peer-reviewed journals.

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Structure of biological apatite: bone and tooth

1

Ahmed Talal 1 , Shorouq Khalid Hamid 1 , Maria Khan 2 , Abdul Samad Khan 1 1 Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia; 2Department of Oral Biology, University of Health Sciences, Lahore, Pakistan

1.1

Introduction

Human body is composed of soft and hard tissue structure that works in dynamic condition. Soft tissues are composed of collagen, elastin, and ground substance. The function of hard tissues is to connect, support, and surround body organs and structures, and hard tissue includes bone and mineralized tooth structures, such as are enamel, dentin, and cementum. The hard tissue is composed of organic and inorganic constituents. Apatite or apatite calcium phosphates are the principal inorganic constituents of bone and teeth. De Jong in 1926 was the first to identify the structure of the apatite CaeP solid phase in the bone by chemical analysis as a crystalline calcium phosphate resembling geological apatite (De Jong, 1926), however, not similar to stoichiometric hydroxyapatite (HA). The bone crystals are thin plates having approximately 500 Å length, 250 Å width, and 100 Å thickness (Finean and Engstrom, 1953). The studies showed that there is an existence of bone mineral ions in nonapatitic arrays, and these ions (bivalent ions, i.e., CO23 and HPO24 ) are mainly present in a hydrated layer on the crystal surfaces (Rey et al., 1990; Eichert et al., 2007). The composition, crystal size, morphology, and stoichiometry of biological apatite are different from the pure HA. The Ca/P molar ratio is 1.67 for pure HA, whereas for enamel and dentin, it is 1.62 and 1.64, respectively. Generally, biological apatites are calcium deficient or nonstoichiometric. The other minors, e.g., magnesium (Mg), carbonate (CO3), sodium (Na), chloride (Cl), acid phosphate (HPO4), etc., and trace elements such as strontium (Sr) and lead (Pb) are associated with these apatites. The biological apatites can be classified as follow: • Carbonate-apatite (CO3)-AP • Fluorecarbonate apatite (F, CO3)-AP It can be represented by the chemical formula given below: A10(BO4)6X2

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00001-X Copyright © 2020 Elsevier Ltd. All rights reserved.

2

Handbook of Ionic Substituted Hydroxyapatites

where • A10 is Ca, Na, Sr, Pb, Cd, Mg, and K • BO4 is (PO4, CO3, VO4, SiO4, AsO4, and HPO4)6 • X is (OH, Cl, CO3F)2 where A represents trace cations with concentration less than 0.1 wt.%. The minor elements, carbonates, magnesium, and fluoride are responsible for the biological apatites stability or instability. Water has also been found in the apatite structure in several forms. The term “apatite” applies to a broad category of structures comprising different constituents. Hydroxyapatite is one such constituent, another being carbonate hydroxyl apatite, where carbonate ions substitute for some of the hydroxyl ions (A type) or the carbonate ions may be present on the phosphate sites (B type) (Bano et al., 2019). Synthetic HA is a representative material for bone substance because of its chemical similarities with the inorganic phase of bone and capability of undergoing bonding osteogenesis. Moreover, it is chemically stable for long period of time in vivo (Khalid et al., 2018a; Cai et al., 2019). In addition to calcium and phosphate, the mineral phase in enamel and dentin contains considerable amounts of sodium, carbonate, and magnesium ions and smaller amounts of potassium, chlorine, and fluorine, and in addition to these, dentin contains citrate ions. Enamel apatite crystals are much larger than those present in both bone and dentin. Human enamel apatite has larger a-axis dimension than pure HA, i.e., 0.944 nm compared with 0.942 nm. The crystallographic properties of enamel and dentin are shown in Table 1.1, and the composition of enamel and dentin is given in Table 1.2. These ions may be incorporated in the enamel and dentin in the form of a magnesium whitlockite and a sodium and carbonate containing apatite. The formula given for

Table 1.1 Crystallographic properties of enamel and dentin. Crystallographic properties: lattice parameters (±0.0003 nm)

Enamel

Dentin

a-axis c-axis

0.9441 0.6880

0.9421 0.6887

Crystallinity index

70e75

33e37

Crystallite size (nm)

0.13  0.03

0.0200  0.0040

Ca/P molar

1.63

1.61

Structure of biological apatite: bone and tooth

3

Table 1.2 Constituents of enamel and dentin. Components

Enamel %

Dentin %

Ca

36.4

36.8

P

17.1

18.0

CO

3.4

6.55

Na

0.64

0.38

Mg

0.43

1.24

F

0.01

0.03

the sodium and carbonate containing apatite could not account for the total amount of carbonate found in enamel, and it may contain excess carbonate and also chloride and fluoride ions: Ca10 (PO4)6 (OH)1.59 (CO)0.15 (Cl)0.1 (F)0.01 The relative proportion of constituents in both enamel and dentin are shown in Table 1.3. Bone and dentin are similar in their mineral contents and chemical composition. Enamel contains more minerals than bone and dentin and contains a slightly carbonated apatite instead of pure apatite; however, it does not contain citrate or highly carbonated compounds. In the subsequent sections, the structure, anatomy, composition, and properties of bone and tooth will be discussed.

1.2

Structure of bone

Structural and mechanical support of the body is one of the most important functions that is provided by the skeleton bones. Additionally, it permits muscle movement by providing levers, protects internal vital structures and organs, maintains the mineral homeostasis and acidebase balance, acts as a pool for the cytokines and growth factors, Table 1.3 Relative proportions of constituents in enamel and dentin. Constituents

Enamel %

Dentin %

Ca10 (PO4)6 (OH)1.59 (CO)0.15 (Cl)0.1 (F)0.01

70

e

Ca9Mg(HPO4) (PO4)6

10

47

Ca8.5 Na1.5(PO4)4.5(CO3)2.5

20

20

Ca9(PO4)4.5(CO3)1.5(OH)1.5

e

30

Ca8(citrate) (PO4)4.5H2O

e

3

4

Handbook of Ionic Substituted Hydroxyapatites

and provides an environment in the marrow spaces for hematopoiesis (Taichman, 2005). Bone is classified into four general categories: long bones, short bones, flat bones, and irregular bones. Long bones involve the humeri, clavicles, ulnae, radii, metacarpals, tibiae, femurs, fibulae, phalanges, and metatarsals, whereas short bones involve the tarsal and carpal bones, sesamoid, and patellae bones. Flat bones involve the skull, mandible, ribs, sternum, and scapulae. Irregular bones involve the bones of basicranium, sacrum, vertebrae, hyoid bone, and coccyx. Formation of long, short, and irregular bones include a combination of both membranous and endochondral ossification, whereas only membranous ossification takes place in the formation of flat bones.

1.2.1

Anatomy of bone

Long bones consist of a long hollow shaft or diaphysis; flared, distal to the growth plates, there are cone-shaped metaphyses, while above to the growth plates, there are rounded epiphyses. The diaphysis consists primarily of dense cortical bone, while the metaphysis and epiphysis are formed by trabecular meshwork bone, which is surrounded by dense cortical bone shell. Skeleton of adult human consists of 80% cortical bone while 20% trabecular bone. The ratios of cortical to trabecular bone differ according to different bones function and skeletal sites. The ratio of cortical to trabecular bone in vertebra is of 25:75, whereas it is 50:50 for the femoral head and 95:5 for radial diaphysis. Dense and solid cortical bone is located around the marrow space, whereas trabecular bone is interspersed inside the bone marrow compartment, and it consists of rods bearing bone marrow and a honeycomb-like network of trabecular plates. Both trabecular and cortical bones are formed by osteons. Haversian systems are the cortical osteons, having cylindrical shape; approximately, they are 400 mm long and 200 mm wide at their base, forming a branching network inside the cortical bone (Clarke, 2008). The haversian system walls are formed by the concentric lamellae. During bone remodeling, old lamellae and new lamellae merge forming interstitial lamellae. Cortical bone outer surface is periosteal, whereas the inner surface is endosteal. Periosteal surface activity is significant in fracture repair and appositional growth. Trabecular bone consists of plates and rods with an average of 50e400 nm thickness. Plates are the trabecular osteons, which are semilunar in shape, normally 35 nm in thickness, and consist of concentric lamellae. Trabecular and cortical bones are naturally formed in a lamellar arrangement with collagen fibrils laid down by osteoblasts in alternating orientations (Clarke, 2008). Outer cortical surface of bone is tightly attached by a fibrous connective tissue called periosteum. The periosteum is connected to the bone by thick collagenous fibers, termed Sharpeys’ fibers, while the inner surface of bone is covered by endosteum, which is a membranous structure, in addition a blood vessel canals (Volkmann’s canals) present inside the bone containing blood vessels, nerves, and lymphatics.

1.2.2

Composition of bones

Bone is composed of inorganic and organic component, where 50%e70% is the inorganic content and 20%e40% is the organic matrix; water is 5%e10%, and lipids

Structure of biological apatite: bone and tooth

5

are approximately 3%. The mineral/inorganic content of bone is mainly HA (Ca10(PO4)6(OH)2), with minor quantities of carbonate, magnesium, and acid phosphate, whereas organic content is mostly composed of type I collagen and bone protein. Inorganic content provides mechanical rigidity and gives load-bearing strength to bone, while flexibility and elasticity is provided by the organic matrix. Collagenous protein is the main protein of bone with a share of 85%e90% of total protein component. Bone matrix is mainly composed of type I collagen, with small numbers of types III, V, and Fibril-Associated Collagens with Interrupted Triple helices (FACIT). FACIT collagens are collection of nonfibrillar collagens that help in developing molecular bridges that are significant for the organization and consistency of extracellular matrices. This family consists of collagens IX, XII, XIV, XIX, XX, and XXI. Noncollagenous proteins comprise 10%e15% of entire bone protein. Nearly 25% of noncollagenous protein is exogenously derived, comprising serum albumin and a2-HS-glycoprotein, which because of their acidic properties bind to HA. Matrix mineralization regulation may be aided by serum-derived noncollagenous proteins, and a2-HS-glycoprotein may control bone cell proliferation. The remaining exogenously derived noncollagenous proteins are composed of growth factors, molecular mediators, and a variety of other various molecules in small amounts, which may affect the activity of bone cells. Osteoblasts synthesize and secrete noncollagenous protein. The noncollagenous proteins are classified broadly into several categories, involving proteoglycans, g-carboxylated (gla) proteins, glycosylated proteins, and glycosylated proteins. Bone proteins aid in multiple functions as regulating bone mineral deposition, turnover, and regulating the bone cell activity (Clarke, 2008).

1.2.3

Properties of bone structure

Bone mass is responsible for 50%e70% of bone strength. In addition, geometry and composition of bone are also essential; however, larger bones have greater strength than small bones, even with equal bone mineral density; thus, as bone diameter increases radially, the strength of bone increases. Other factors can also influence the bone strength, such as volume and amount of trabecular and cortical bone at a specified skeletal location. Bone microstructure, bone turnover, mutations in proteins, osteomalacia, fluoride therapy, and hypermineralization also can affect bone strength (Pocock et al., 1987). To understand the mechanical properties of bone, it is important to understand the mechanical properties of its component phases, in addition to the structural association between them at different levels of hierarchical organizational structure. The level of structures are described in Table 1.4. The hierarchically ordered structure has an irregular, however, optimized organization and orientation of its components forming heterogeneous and anisotropic bone material (Fig. 1.1) (Rho et al., 1998). Bone mechanical properties are influenced by different structural levels. One of the examples is the Young’s modulus; in the large tensile cortical specimen, it is in the range of 14e20 GPa (Reilly et al., 1974), whereas that of microbending cortical specimen showed 5.4 GPa (Choi et al., 1990). Nevertheless, it is not clear whether this difference is because of the testing method or the effect of microstructure. Properties such as

6

Handbook of Ionic Substituted Hydroxyapatites

Table 1.4 Levels of hierarchical organizational structure. Macrostructure:

Cancellous bone Cortical bone

Microstructure (10e500 mm):

Haversian systems Osteons Single trabeculae

Submicrostructure (1e10 mm):

Lamellae

Nanostructure (few hundred nanometerse1 mm):

Fibrillar collagen and embedded mineral

Subnanostructure (less than a few hundred nanometers):

Molecular structure of constituent elements (mineral, collagen, noncollagenous organic proteins)

Collagen molecule Cancellous bone

Lamella

Collagen fiber

Collagen fibril

Cortical bone Bone Osteon

Crystals

Haversian canal

0.5 mm 1 nm 10–500 mm

3–7 mm

Microstructure Macrostructure

Nanostructure Sub-microstructure

Sub-nanostructure

Figure 1.1 Hierarchical structural organization of the bone (Rho et al., 1998).

hardness, elastic modulus, and composite apparent density have notably higher values in mandibular bone as compared with maxillary bone. Mandibular composite apparent density is 76% greater than maxillary (1.18 vs. 0.67 g/cm3). Elastic modulus and hardness in the posterior region are significantly greater than in the anterior region, whereas composite apparent density showed the opposite values. Compared with trabecular bone, cortical bone has higher elastic modulus, whereas the hardness of the two bones are comparable (Seong et al., 2009). Ultimate tensile strength of the metaphyseal shell is

Structure of biological apatite: bone and tooth

7

101  26 MPa in the longitudinal direction, whereas it is 50  12 MPa in the transverse directions (Lotz et al., 1991). Generally, values of the bone mechanical properties at the macrostructural level differ among the type of bone as well as within same bone at different locations (Rho et al., 1995; Goldstein, 1987). Cortical bone has greater stiffness and stress resistance than trabecular bone; however, it has higher brittleness. Trabecular bone has the ability to withstand strains up to 30%, while cortical bone fails with only 2% strains (Osterhoff et al., 2016). At macrostructure level, the bone is classified into the cortical (compact) and cancellous (trabecular) bone types. In cross-sectional view, the end of a long bone such as femur has a dense cortical shell, which has a porous, cancellous bone inside it, whereas flat bones, for example, the calvaria, have a sandwich-like structure that consists of dense cortical layers at the outer surfaces, whereas in the center, there is reinforcing thin cancellous bone. Generally, cancellous bone has greater metabolic activities with more remodeling than the cortical bone, thus making it younger in average than the cortical bone. Consequently, although cancellous and cortical bone may have similar kind of material, however, the mechanical properties of the cortical bone material at the microstructural level may alter by its maturation. Mechanical properties of cortical bone get influenced significantly by the porosity, the level of the mineralization, and the arrangement of the solid matrix (Rho et al., 1998). At microstructure level, the mineralized collagen fibers form lamellae are arranged in planar arrangements of 3e7 mm in width. In some situations, these lamellae form osteon (haversian system) by wrapping in concentric layers of 3e8 lamellae around the central canal. The osteonal segments with lamellar orientation in the longitudinal direction showed elastic moduli of 12 in tension and strength of 120 MPa, whereas the osteons with adjacent lamella orientations in sharp angles to each other showed elastic moduli of 5.5 GPa in tension and strengths of 102 MPa. On the other hand, the modulus of trabecular bone material ranges from 1.0 to 20 GPa (Marshall et al., 1997). Submicrostructure level consists of the bone lamellae with 3e7 mm thickness (Marotti, 1993), however, the orientation and arrangement of the lamella substance is not fully known. There could be variation in the lamellae encountered in cancellous and cortical bone. Commonly, the collagen fibers are arranged in a lamella of an osteon that lies in parallel direction, with an alteration in the arrangement of fibrils from one lamella to the other (Giraud-Guille, 1988). The most prominent structures seen at the nanostructure level is collagen fiber that ranges from hundreds of nanometers to 1.0 mm; it is infiltrated and surrounded by the minerals. The subnanostructures level consists of three main components: collagens, crystals, and noncollagenous organic proteins (Weiner and traub, 1992). The plate-like apatite crystals arise within the separate spaces inside the collagen fibrils, thus causing a limitation in the possible primary growth of the mineral crystals, additionally forcing the crystals to be discontinuous. Commonly, its length and width is 50  25 nm, while its thickness is 2e3 nm. The nanocrystalline bone apatite has minor amounts of impurities; however, it is still significant, and those impurities include HPO4, Na, Mg, citrate, carbonate, K, and others. Type I collagen present in the matrix is its principal organic component. It is secreted by osteoblasts, which themselves self-assemble into fibrils

8

Handbook of Ionic Substituted Hydroxyapatites

with a certain tertiary structure with a periodicity of 67 nm and gaps of 40 nm among the ends of the molecules. Phosphoproteins that include osteopontin, sialoprotein, osteonectin, and osteocalcin are noncollagenous organic proteins. These proteins could act as a pool for phosphate or calcium ions for mineral formation (Rho et al., 1998).

1.3

Structure of tooth

Teeth are the hardest and chemically most stable tissue in the body. In human, the function of teeth is the communication of food by a process referred to as mastication. The anatomical crown of the tooth is covered with enamel. Enamel is thickest over the region of the crown and tapers toward the neck (cervical region) of the tooth. Where enamel finishes, it is usually continuous with a cementum layer, which covers the root area. The cementum is also a hard tissue, which helps in the attachment of the tooth to the surrounding alveolar bone by means of collagen fibers. The dentin is situated between the pulp and enamel and is interlinked with dentinoenamel junction (DEJ). The pulp is a highly vascular connective tissue, which provides nutrients as well as sensory nerves to the cell forming the dentin (Goldberg et al., 2011b). The schematic structure of tooth is given in Fig. 1.2.

1.3.1

Enamel

Enamel is the hardest calcified structure of the body and covers the crown of the teeth. It is translucent, and its color varies from light yellow to gray white due to variation in its thickness from maximum of 2.5 mm to knife-edge at the cervical margin. It consists of 97% by weight (92% by vol.) of inorganic material, which is crystalline calcium phosphate (HA), 1% organic material, and 2% water. The organic material is a calcium-deficient apatite that contains 2%e5% (wt%) carbonate and numerous other

En

el

Dentin

el

am

am

En

Den

tin

Root

Figure 1.2 Transmission light microscopic image of human molar tooth (Tesch et al., 2001).

Structure of biological apatite: bone and tooth

9

trace elements. The mineral crystals are long, ribbon-like carbonate-apatite crystals, which are closely packed and measure 160e1000 nm long, 40e120 nm wide, and 25e30 nm thick. These crystals grouped together to form enamel rods or interrod enamel. Mature enamel is almost completely composed of minerals, which are in the form of bundles consisting of highly elongated crystals, rendering the enamel hard and brittle. In the enamel, the HA crystals form rod- or prism-like structures that are well defined and are about 4 mm in diameter (Abou neel et al., 2016). Enamel rods (prisms) and interrod enamel (interprismatic substance) form the basic organizational structural units of enamel. Structure of crystals in enamel rods and interrod enamel is same; however, they have differed orientation (Fig. 1.3). Around the long axis of tooth, enamel rods are arranged circumferentially in groups. The orientation of enamel rods is perpendicular to the dentin surface in general. The orientation of the enamel rods in the cervical area is mainly horizontal, whereas they are vertically orientated at the cusp tips.

1.3.2

Dentin

Dentin can be considered as a natural composite, similar to bone, consisting of microscopic filler phase made up of apatite crystallites and an organic matrix made up predominantly of collagen. Dentin is a mineralized connective tissue, which makes up the bulk of teeth; it is produced by odontoblasts, which are highly specialized cells. It consists of 45 vol.% mineral/inorganic component, 33 vol.% organic matrix, and 22 vol.% water content, which is approximately 70 wt.% mineral/inorganic component, 20 wt.% organic material, and 10 wt.% water. The inorganic component consists mainly of HA. About 56% of the mineral phase is inside the collagen, which makes

Figure 1.3 A scanning electron microscope image shows enamel rods and interrod enamel having similar structure but different orientation.

10

Handbook of Ionic Substituted Hydroxyapatites

dentin slightly harder than the bone, yet softer than enamel. Similar to bone and enamel, dentin contains a Ca-deficient HA with more than trace amounts of carbonate. Apatite crystals of dentin are much smaller (approximately 5  30  100 nm) than the enamel apatite crystals and contain 4%e5% carbonate compared with HA. The major components are distributed into particular morphological features to form a complex and vital hydrated composite whose morphology changes with location and undergoes modification with stimuli (age and diseases) (Marshall et al., 1997; Saxena et al., 2019). The minerals in dentin are in the form of crystalline and amorphous calcium phosphates. The crystals are 3 nm in diameter and 64 nm long. Morphologically, dentin is composed of evenly spaced dentinal tubules (approximately 1e2 mm in diameter), which extend from the dentinepulp interface out to enameledentin interface. The tubules have an ‘S’-shaped curvature, which is more pronounced in the crown of the tooth, and a secondary, spiral curvature with a periodicity and amplitude of a few microns. Each tubule contains the cytoplasmic process of a cell (odontoblasts). Around most tubules, there is a heavily mineralized cylinder of peritubular dentin, which is devoid of or has very little matrix. Throughout life, a layer of nonmineralized dentin, known as predentin, separates the odontoblasts from the mineralized regions. The predentinedentin interface has an irregularscalloped appearance. This is due to the occurrence of mineralized dentin with a spheroidal morphology. All of the dental tissues are deposited incrementally, and unlike bone, dental tissues are rarely remodeled (Arola et al., 2009). Close to the pulp, the diameter of the tubules is largest, approximately 2.5 mm, which decreases to less than 1 mm at the DEJ. Dentinal tubules, besides the main tubule, also have many branches and outgrowths. In areas of low tubular density, branches are higher in number, thus forming a canalicular, anastomosing system similar to osteocytes in bone. Three different tubular branches types were recognized by Major and Nordahl (1996) based on size, direction, and location. These tubular branches were classified as major, fine, and microbranches having diameter of 0.5e1 mm, 300e700 nm, and 25e200 nm, respectively. Major branches are mostly present on the peripheries, fine branches are present in areas where tubules have low density, whereas microbranches are present in all parts of the dentin (Mjor and Nordahl, 1996). Collagen constitutes 90% of the dentin organic matrix, and the major component of the dentin collagen is type I; it also contains type III and V traces. Proteoglycans and other noncollagenous proteins constitute approximately 10% of the dentin organic matrix, whereas lipids form the remaining 2%. Odontoblasts produce noncollagenous proteins that play an important role in the dentin mineralization. These are present between the dentinal tubules and are assembled along the dentinal tubule walls (Orsini et al., 2009). According to formation phases, dentin can be divided into five different types: DEJ, mantle dentin, primary dentin, secondary dentin, and tertiary dentin. This classification reflects the changes in the basic components of the structure as defined by changes in their arrangement, interrelationships, or chemistry (Marshall et al., 1997). In contrary to previous belief, DEJ being a simple anatomical interface between dentin and enamel, recent studies have shown that DEJ is distinct from both dentin and enamel, consists of considerable amounts of organic and mineral material, and is 7e15 mm

Structure of biological apatite: bone and tooth

11

wide. In this region, certain enzymes and growth factors are also present, which on being discharged can produce effects in an area away from the DEJ. In human teeth, the DEJ is wavy or scalloped and not smooth. Mechanical connection between dentin and enamel is believed to improve because of this pattern. The scallops size ranges between 25 and 50 mm, and they are bigger and deeper at the dentin cusps and incisal edges, leveling down near the cervical region (Goldberg et al., 2011a). The mantle dentin is the outer dentin layer and is present in the coronal part of the tooth; it is 5e20 mm in thickness and is less mineralized, more elastic, and resilient. Because of these characteristics, it may adapt to dissipate pressures or stress forces. Additionally, in the mantle dentin, dental tubules are either not present or present in reduced numbers and are thin and curved. Hence, the amount of peritubular dentin increases remarkably, whereas intertubular dentin amount decreases toward the pulp from the DEJ. Primary dentin constitutes main portion of the dentin and make up the bulk of the tooth and provides its form and size. It is different from the mantle dentin; its collagen matrix is more compact and odontoblasts form its organic matrix completely. Rapid primary dentin formation is followed by continuous slow dentin formation, which is termed as secondary dentin. Although the exact point when the odontoblast activity changes form primary to secondary dentin formation is not known, it is believed that the primary dentin formation ends with the crown completion. Ending of primary dentin formation also coincides with the tooth becoming functional and its root completion. In response to any external irritation to the tooth structure such as caries, cavity preparation, abrasion, attrition, erosion, or trauma, tertiary dentin is formed as a defensive mechanism. Formation of tertiary dentin aims to provide protection to the pulp by increasing the thickness of dentin between the pulpal tissue and oral microbes, thus isolating the pulp tissue. The external wear of tooth tends to induce more tertiary dentin as compared with caries. The potency and application time of external stimulus effect the regularity and form of tertiary dentin. Tertiary dentin can be further classified into reactionary and reparative dentin, which are produced by original primary odontoblasts and newly differentiated replacement odontoblasts, respectively. Tertiary dentin forms a relatively impermeable barrier between tubular dentin and pulp tissue as it is atubular. Formation of reactionary and reparative dentin under a remaining dentin thickness of as thin as 0.5 mm impart dentinepulp complex renders an exceptional ability to survive even under extensive dentinal damage (Tj€aderhane et al., 2009; Goldberg et al., 2011a).

1.3.3

Pulp

Dental pulp is the unmineralized soft connective tissue, which resides in the pulp cavity of the tooth. Pulp can be divided into two parts anatomically: (i) coronal pulp and (ii) radicular pulp. Coronal pulp present in the pulp chamber in the crown of the tooth, whereas radicular pulp is present in the pulp canals in the root of the tooth. It is considered as a unique organ, which is specialized to perform four functions: (i) formative or

12

Handbook of Ionic Substituted Hydroxyapatites

developmental, (ii) nutritive, (iii) sensory or protective, and (iv) defensive or reparative. The major cells of pulp are odontoblasts, fibroblasts, undifferentiated ectomesenchymal cells, and macrophages in addition to other immunocompetent cells. Odontoblasts, which are pulp-specialized cells and undifferentiated mesenchymal cells, which on stimulation may differentiate into dentin-forming cells, reside in the pulp. They have the ability to produce dentin; this enables the pulp to form a barrier to protect itself from the irritants and partially make up for the loss of dental tissue due to wear or caries. Throughout the tooth life, secondary dentin is laid down at a slow rate at the circumference by the odontoblasts, which withdraw toward the pulp center after secreting the dentinal matrix. In addition, in response to caries processes or restorative procedure, odontoblast may also produce sclerotic, reactionary, and reparative dentine. Fibroblasts have the highest number among all the cells in the pulp, they perform the function of forming and maintaining the pulp matrix, which consists of collagen and ground substance (Yu and abbott, 2007). The matrix of the pulp consists of collagen fibers and ground substance. The principal fibers are type I and type II collagen fibers, whereas the ground substance is composed of glycosaminoglycans, glycoproteins, and water. Ground substance supports the pulp cells and acts as a medium for the nutrients transport and metabolites among the cells and the vasculature. Dental pulp also contains sensory and autonomic nerves to serve its vasomotor, chemoreceptor, and defensive mechanism. Sensory nerves that are branches of maxillary and mandibular divisions of trigeminal nerve take part in pain perception and transduction. These nerves, along with the blood vessels, form the neurovascular bundle and enter the pulp through the apical foramen. As the nerves extend occlusally in the pulp core, they form branches, which contribute to an extensive plexus of nerves in the crown portion of the tooth just below the odontoblasts cell bodies (Linde, 1985; Yu and abbott, 2007; Abd-elmeguid and yu, 2009; Pashley, 1996). Pulp plays a major role in the defensive mechanism against the bacterial invasion into the dental tubules, dentinal fluid, and odontoblastic processes in the dentinal tubules that collectively behave as a positively charged hydrogel (Linden et al., 1995). Dentinal fluid outward flow also plays an important defensive mechanism opposed to the entry of harmful substances as it hinders the rate at which toxic substances diffuse into the dentinal tubules from the mouth (Matthews and Vongsavan, 1994). Additionally, in response to the bacterial infection, dentinal fluid may contain antibodies and antimicrobial agents.

1.3.4

Cementum

Cementum is an avascular hard connective tissue deposited by cementoblasts; it covers the entire anatomic roots of the teeth in thin layer. Unlike the other tissues of the periodontium, it does not normally undergo remodeling. Cementum as one of its main function attaches the principal collagen fibers of the periodontal ligament to the surface of the root; these fibers are embedded to the cementum on one side and into the bone on the other side, thus anchoring the teeth in their sockets.

Structure of biological apatite: bone and tooth

13

It consists of approximately 45e50 wt.% inorganic material and 50e55 wt.% organic matter and water. Cementum inorganic material is HA with minute amounts of amorphous calcium phosphates, whereas cementum organic portion is primarily collagen and protein; approximately 90% of the organic matrix constitutes of type I collagen and approximately 5% accounts for type III collagen. Cementum is light yellow in color and continuously formed throughout life to replace the aged cementum (Nanci and Bosshardt, 2006). Two types of cementum are acellular and cellular cementum, and these cover the coronal half and apical half of the tooth root, respectively; however, no clear boundary exist between these two types (Fig. 1.4) (Matalova et al., 2015). Wear of the tooth occlusal or incisal surface and passive eruption of the tooth are compensated by the increase in the cementum thickness on the root end of the tooth (Bosshardt and selvig, 1997).

1.3.5

Properties of tooth structure

Teeth require unique optimal mechanical properties to perform their function of incision, laceration, and food grinding during mastication. Mechanical properties of

Cementum–dentin junction Alveolar bone

Periodontium

Acellular cementum

Dentin

Predentin

Sharpey’s fibers

Collagen fibers

Fibroblasts Cementoblasts

Cellular cementum

Cementocytes

Odontoblasts

Figure 1.4 Acellular and cellular cementum distribution around the root (Matalova et al., 2015).

14

Handbook of Ionic Substituted Hydroxyapatites

human tooth structure have been measured; however, the reported values vary with each other. This attributes to the technical problems related with preparing and analyzing the small samples. There is variation in enamel and dentin properties from one type of tooth to another. The properties of enamel are according to its location on the tooth, i.e., the cuspal enamel is stronger than enamel on the other surface of the tooth. The properties also vary according to the histological structure. It can be justified as the enamel is stronger under longitudinal compression than the lateral compression. Table 1.5 indicates a variation in the properties of enamel and dentin. Wear properties and reliability of a tooth are affected by the mechanical properties of dentin and enamel. Hard enamel resists the wear, whereas an important role is played by the soft zone in dentin under DEJ in the distribution of strain and resistance of fracture during the process of mastication (Brauer et al., 2011). Enamel and dentin are made up of crystalline calcium phosphate, protein, and water. Ratio and interface between these components dictate the mechanical properties of the tooth structure. Enamel at the microscale is made up of highly mineralized, stiff rods, which are embedded in a soft protein matrix. Individual rods at the nanoscale level present a composite assembly with mineral crystals acting as reinforcements and proteins as matrix (Zimmerman et al., 2010; Habelitz et al., 2001; Eimar et al., 2012). Orientation of HA crystals within a rod strongly effects the mechanical properties of enamel rods (Jeng et al., 2011). Enamel has anisotropic properties; its mechanical properties depend on the direction of the load in relation to enamel rod and interrod orientation. Higher values for mechanical properties are observed when load is applied parallel to the direction of enamel rods, and similarly lower values are observed with the load applied perpendicular to the enamel rods (Shahmoradi et al., 2014). Dentin structure can be considered as a continuous fiber reinforced composite, in which intertubular dentin forms the matrix and the tubule lumens along with their peritubular dentin forms the cylindrical fiber reinforcement. The mechanical properties of dentin depend on the tubular density, orientation, and the mineral phase average density. Dentin structure and properties gradually change from the DEJ toward the pulp. Its hardness and modulus of elasticity increases with increasing distance form DEJ (Kinney et al., 2003; Brauer et al., 2011). Dentin is anisotropic, higher mechanical response is observed when the load is applied parallel to dentinal tubules, and lower

Table 1.5 Properties of human third molar tooth structure (Xu et al., 1998). Microhardness indentation methods

Enamel occlusal section

Enamel axial section

Dentin

Hardness (GPa)

3.23e3.62

3.03

0.58

Toughness (MPa.m )

0.77

0.52e1.30

e

Elastic modulus (GPa)

94

80

20

Indentation energy (mJ)

2.6

2.7

7.5

2

Structure of biological apatite: bone and tooth

15

mechanical response is observed when applied load is perpendicular to the dental tubules. Additionally, higher hardness and elastic modulus are observed for peritubular dentin than intertubular dentin, which can be due to its higher mineralization content (Poolthong et al., 1998; Kinney et al., 1996). Other studies reported that the hardness of dentin and enamel is between 0.2e2.5 GPa and 16.3e29.8 GPa, whereas the elastic modulus is 1.3e4.9 GPa and 70e120 GPa, respectively. The elastic modulus of enamel is critical for its resistance against wear and in facilitating the cutting and chewing of food (Angker and swain, 2006; An et al., 2015). Under compression, compressive strength and the Young’s modulus of human enamel vary from 100 to 400 MPa and from 10 to 80 GPa, respectively (Zaytsev, 2016). The mechanical and thermal properties of human dentin are shown in Tables 1.6 and 1.7, respectively. As the success of a restoration depends largely on the integrity of the interface between the restorative material and tooth structure, understanding the mechanical properties of enamel and dentin is important. The knowledge of these mechanical and thermal properties is crucial in foreseeing dentin/restoration interface behavior and understanding how dentin’s performance and strength is altered by its aging and disease.

Table 1.6 Mechanical properties of human dentin. Properties

Values

Compressive strength (MPa)

230e370

Young’s modulus (GPa)

10.1e19.3

Shear strength (MPa)

36e138

Tensile strength (MPa)

31e104

Microhardness (GPa)

0.25e0.8

Elastic modulus (GPa)

15e21

Table 1.7 Thermal properties of human dentin. Properties Specific heat (cal g

Values 1

1

K )

0.30

Thermal conductivity (W m Thermal diffusivity ( 10

3

1

1

K ) 2

0.57

1

cm s )

Coefficient of thermal expansion (10

1.8e2.6 6

1

C )

9.0

16

1.4

Handbook of Ionic Substituted Hydroxyapatites

Conclusion

Because of high similarity to biominerals such as tooth and bone, calcium phosphates are osteoconductive and exhibit strong affinity to host hard tissue and biocompatible materials. Ability of synthetic apatites to form a chemical bond with hard tissue offers a greater advantage in clinical applications. Because of this property, these biomaterials have been used as implants in clinical bone repair and regeneration materials, coating materials in tissue engineering, drug delivery devices, and tumor treatment, and as dental restorative materials (Khalid et al., 2018a,b; Khan and Syed, 2019). Still, studies are being conducted to evaluate the performances of these biomaterials in clinical applications and their interaction with biological tissues. Therefore, it is important to know the basic structure of biological structure and should have broad understanding of biological apatite to achieve a better clinical performance.

References Abd-Elmeguid, A., Yu, D.C., 2009. Dental pulp neurophysiology: part 1. Clinical and diagnostic implications. J. Can. Dent. Assoc. 75, 55e59. Abou Neel, E.A., Aljabo, A., Strange, A., Ibrahim, S., Coathup, M., Young, A.M., Bozec, L., Mudera, V., 2016. Demineralization-remineralization dynamics in teeth and bone. Int. J. Nanomed. 11, 4743e4763. An, B., Wang, R., Arola, D., Zhang, D., 2015. Damage mechanisms in uniaxial compression of single enamel rods. J. Mech. Behav. Biomed. Mater. 42, 1e9. Angker, L., Swain, M., 2006. Nanoindentation: application to dental hard tissue investigations. J. Mater. Res. 21, 1893e1905. Arola, D., Ivancik, J., Majd, H., Fouad, A., Bajaj, D., Zhang, X.-Y., Eidelman, N., 2009. Microstructure and mechanical behavior of radicular and coronal dentin. Endod. Top. 20, 30e51. Bano, N., Jikan, S.S., Basri, H., Adzila, S., Zago, D.M., 2019. XRD and FTIR study of A&B type carbonated hydroxyapatite extracted from bovine bone. In: AIP Conference Proceedings. AIP Publishing, p. 020100. Bosshardt, D.D., Selvig, K.A., 1997. Dental cementum: the dynamic tissue covering of the root. Periodontology 2000 13, 41e75. Brauer, D.S., Hilton, J.F., Marshall, G.W., Marshall, S.J., 2011. Nano-and micromechanical properties of dentine: investigation of differences with tooth side. J. Biomech. 44, 1626e1629. Cai, B., Jiang, N., Zhang, L., Huang, J., Wang, D., Li, Y., 2019. Nano-hydroxyapatite/polyamide66 composite scaffold conducting osteogenesis to repair mandible defect. J. Bioact. Compat Polym. 34, 72e82. Choi, K., Kuhn, J.L., Ciarelli, M.J., Goldstein, S.A., 1990. The elastic moduli of human subchondral, trabecular, and cortical bone tissue and the size-dependency of cortical bone modulus. J. Biomech. 23, 1103e1113. Clarke, B., 2008. Normal bone anatomy and physiology. Clin. J. Am. Soc. Nephrol. 3 (Suppl. 3), S131eS139.

Structure of biological apatite: bone and tooth

17

De Jong, W., 1926. La substance minerale dans les os. Recl. Trav. Chim. Pays-Bas 45, 445e448. Eichert, D., Drouet, C., Sfihi, H., Rey, C., Combes, C., 2007. Nanocrystalline apatite-based biomaterials: synthesis, processing and characterization. Biomater. Res. Adv. 93e143. Eimar, H., Ghadimi, E., Marelli, B., Vali, H., Nazhat, S.N., Amin, W.M., Torres, J., Ciobanu, O., Junior, R.F.A., Tamimi, F., 2012. Regulation of enamel hardness by its crystallographic dimensions. Acta Biomater. 8, 3400e3410. Finean, J., Engstrom, A., 1953. Low angle x-ray diffraction in bone. Nature 171, 564. Giraud-Guille, M.M., 1988. Twisted plywood architecture of collagen fibrils in human compact bone osteons. Calcif. Tissue Int. 42, 167e180. Goldberg, M., Kulkarni, A.B., Young, M., Boskey, A., 2011a. Dentin: structure, composition and mineralization. Front. Biosci. Elite Edit. 3, 711e735. Goldberg, M., Kulkarni, A.B., Young, M., Boskey, A., 2011b. Dentin: structure, Composition and Mineralization: the role of dentin ECM in dentin formation and mineralization. Front. Biosci. Elite Edit. 3, 711. Goldstein, S.A., 1987. The mechanical properties of trabecular bone: dependence on anatomic location and function. J. Biomech. 20, 1055e1061. Habelitz, S., Marshall, S.J., Marshall, G.W., Balooch, M., 2001. Mechanical properties of human dental enamel on the nanometre scale. Arch. Oral Biol. 46, 173e183. Jeng, Y.-R., Lin, T.-T., Hsu, H.-M., Chang, H.-J., Shieh, D.-B., 2011. Human enamel rod presents anisotropic nanotribological properties. J. Mech. Behav. Biomed. Mater. 4, 515e522. Khalid, H., Suhaib, F., Zahid, S., Ahmed, S., Jamal, A., Kaleem, M., Khan, A.S., 2018a. Microwave-assisted synthesis and in vitro osteogenic analysis of novel bioactive glass fibers for biomedical and dental applications. Biomed. Mater. 14, 015005. Khalid, H., Syed, M.R., Rahbar, M.I., Iqbal, H., Ahmad, S., Kaleem, M., Matinlinna, J.P., Khan, A.S., 2018b. Effect of nano-bioceramics on monomer leaching and degree of conversion of resin-based composites. Dent. Mater. J. 37, 940e949. Khan, A.S., Syed, M.R., 2019. A review of bioceramics-based dental restorative materials. Dent. Mater. J. 38, 163e176. Kinney, J., Balooch, M., Marshall, S., Marshall, G., Weihs, T., 1996. Atomic force microscope measurements of the hardness and elasticity of peritubular and intertubular human dentin. J. Biomech. Eng. 118, 133e135. Kinney, J., Marshall, S., Marshall, G., 2003. The mechanical properties of human dentin: a critical review and re-evaluation of the dental literature. Crit. Rev. Oral Biol. Med. 14, 13e29. Linde, A., 1985. The extracellular matrix of the dental pulp and dentin. J. Dent. Res. 64, 523e529. Spec No. Linden, L.A., Kallskog, O., Wolgast, M., 1995. Human dentine as a hydrogel. Arch. Oral Biol. 40, 991e1004. Lotz, J.C., Gerhart, T.N., Hayes, W.C., 1991. Mechanical properties of metaphyseal bone in the proximal femur. J. Biomech. 24, 317e329. Marotti, G., 1993. A new theory of bone lamellation. Calcif. Tissue Int. 53 (Suppl. 1), S47eS55. Discussion S56. Marshall Jr., G.W., Marshall, S.J., Kinney, J.H., Balooch, M., 1997. The dentin substrate: structure and properties related to bonding. J. Dent. 25, 441e458.

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Matalova, E., Lungova, V., Sharpe, P., 2015. Chapter 26 e development of tooth and associated structures. In: Vishwakarma, A., Sharpe, P., SHI, S., ramalingam, M. (Eds.), Stem Cell Biology and Tissue Engineering in Dental Sciences. Academic Press, Boston. Matthews, B., Vongsavan, N., 1994. Interactions between neural and hydrodynamic mechanisms in dentine and pulp. Arch. Oral Biol. 39 (Suppl. l), 87se95s. Mjor, I.A., Nordahl, I., 1996. The density and branching of dentinal tubules in human teeth. Arch. Oral Biol. 41, 401e412. Nanci, A., Bosshardt, D.D., 2006. Structure of periodontal tissues in health and disease*. Periodontology 2000 40, 11e28. Orsini, G., Ruggeri Jr., A., Mazzoni, A., Nato, F., Manzoli, L., Putignano, A., DI lenarda, R., Tj€aderhane, L., Breschi, L., 2009. A review of the nature, role, and function of dentin noncollagenous proteins. Part 1: proteoglycans and glycoproteins. Endod. Top. 21, 1e18. Osterhoff, G., Morgan, E.F., Shefelbine, S.J., Karim, L., Mcnamara, L.M., Augat, P., 2016. Bone mechanical properties and changes with osteoporosis. Injury 47, S11eS20. Pashley, D.H., 1996. Dynamics of the pulpo-dentin complex. Crit. Rev. Oral Biol. Med. 7, 104e133. Pocock, N.A., Eisman, J.A., Hopper, J.L., Yeates, M.G., Sambrook, P.N., Eberl, S., 1987. Genetic determinants of bone mass in adults. A twin study. J. Clin. Investig. 80, 706e710. Poolthong, S., Swain, M., Sumii, T., Mori, T., 1998. Effect of Tubule Orientation on Some Mechanical Properties of Dentine. Journal of Dental Research. SAGE PUBLICATIONS INC, THOUSAND OAKS, CA, USA, 847-847. Reilly, D.T., Burstein, A.H., Frankel, V.H., 1974. The elastic modulus for bone. J. Biomech. 7, 271e275. Rey, C., Shimizu, M., Collins, B., Glimcher, M.J., 1990. Resolution-enhanced fourier transform infrared spectroscopy study of the environment of phosphate ions in the early deposits of a solid phase of calcium-phosphate in bone and enamel, and their evolution with age. I: investigations in thev 4 PO 4 domain. Calcif. Tissue Int. 46, 384e394. Rho, J.-Y., Kuhn-Spearing, L., Zioupos, P., 1998. Mechanical properties and the hierarchical structure of bone. Med. Eng. Phys. 20, 92e102. Rho, J.Y., Hobatho, M.C., Ashman, R.B., 1995. Relations of mechanical properties to density and CT numbers in human bone. Med. Eng. Phys. 17, 347e355. Saxena, N., Habelitz, S., Marshall, G.W., Gower, L.B., 2019. Remineralization of demineralized dentin using a dual analog system. Orthod. Craniofac. Res. 22, 76e81. Seong, W.J., Kim, U.K., Swift, J.Q., Heo, Y.C., Hodges, J.S., Ko, C.C., 2009. Elastic properties and apparent density of human edentulous maxilla and mandible. Int. J. Oral Maxillofac. Surg. 38, 1088e1093. Shahmoradi, M., Bertassoni, L., M Elfallah, H., Swain, M., 2014. Fundamental Structure and Properties of Enamel. Dentin and Cementum. Taichman, R.S., 2005. Blood and bone: two tissues whose fates are intertwined to create the hematopoietic stem-cell niche. Blood 105, 2631e2639. Tesch, W., Eidelman, N., Roschger, P., Goldenberg, F., Klaushofer, K., Fratzl, P., 2001. Graded microstructure and mechanical properties of human crown dentin. Calcif. Tissue Int. 69, 147e157. Tj€aderhane, L., Carrilho, M.R., Breschi, L., Tay, F.R., Pashley, D.H., 2009. Dentin basic structure and compositiondan overview. Endod. Top. 20, 3e29. Weiner, S., Traub, W., 1992. Bone structure: from angstroms to microns. FASEB J. 6, 879e885.

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Xu, H.H., Smith, D.T., Jahanmir, S., Romberg, E., Kelly, J.R., Thompson, V.P., Rekow, E.D., 1998. Indentation damage and mechanical properties of human enamel and dentin. J. Dent. Res. 77, 472e480. Yu, C., Abbott, P.V., 2007. An overview of the dental pulp: its functions and responses to injury. Aust. Dent. J. 52, S4eS16. Zaytsev, D., 2016. Mechanical properties of human enamel under compression: on the feature of calculations. Mater. Sci. Eng. C 62, 518e523. Zimmerman, B., Datko, L., Cupelli, M., Alapati, S., Dean, D., Kennedy, M., 2010. Alteration of dentineenamel mechanical properties due to dental whitening treatments. J. Mech. Behav. Biomed. Mater. 3, 339e346.

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Analytical tools for substituted hydroxyapatite

2

Mariam Raza 1 , Saba Zahid 2 , Anila Asif 2 1 Science of Dental Materials, PMC, Dental Institute, Faisalabad Medical University, Faisalabad, Pakistan; 2Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

Introduction The global increase in the number of bone fractures and inborn bony defects in both developed and developing countries has highlighted the excessive demand for bone substitutes. However, the potential possibility of rejection followed by morbidity regarding autografts and allografts has whetted the apatite for xenografts (Dumitrescu, 2011). New developments in the chemistry and biomechanics of the bony grafts and fixation devices have revolutionized the invasive procedures and healing times. Bioceramics have remained an area of intense research to satisfy the hunger pang of synthetic bone substitutes since long (Dorozhkin, 2010). Synthetic hydroxyapatite (HA) (Ca10(PO4)6(OH)2), “the Cinderella of bioceramics” has found its extensive use in medicine and dentistry because of its chemical similarity with the inorganic mineral content of bones and teeth. However, the major limitations of HA such as brittleness, decreased mechanical strength, and slow rate of biological interaction as a bone substitute have provoked the urge for ionic substitutions in HA lattice to render it similar to natural bone tissue. The single or multicationic and anionic substitu3 tions (Ag2þ, Mg2þ, Zn2þ, Co2þ, Sr2þ, Si2þ, F, CO2 3 , PO4 ) have influenced the crystallographic, chemical, biological, morphological, thermal, mechanical, and physical properties of synthetic HA. Many substituted HAs are now commercially available for biomedical and dental applications including bone defects repair, bone augmentations, load-bearing orthopedic implants, and dental restorations. This chapter aims to highlight the analytical tools to characterize the established and recent substituted HA. The characterization of materials is important for understanding their properties and applications. The techniques normally adopted are Fouriertransform infrared spectroscopy (FTIR), X-ray diffraction (XRD), scanning electron microscopy (SEM), differential thermal analysis/thermogravimetric analysis (DTA/ TGA), and Raman, nuclear magnetic resonance (NMR), etc.

2.1

Structure of hydroxyapatite

HA (Ca10(PO4)6(OH)2) is hexagonal with 44 atoms per unit cell. The 10 calcium ions (6 CaII and 4 CaI) are distributed among the two sites. There are three oxygen triangles, Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00002-1 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Handbook of Ionic Substituted Hydroxyapatites

(a)

(b)

H O Ca(2) O(1)

c

P

Ca(1)

O(3) P O(2)

a

b

Ca(2) O

H

Ca

P

Ca(1)

Ca(2)

Figure 2.1 (a) The unit cell of hydroxyapatite (HA) in P63/m space group (Mostafa and Brown, 2007); (b) unit cell of HA from the calculated atomic positions (Zilm et al., 2016). Substitutions can occur in HA for the calcium ions, the phosphate groups, or the hydroxyl groups. It is assumed that cations (zinc, magnesium, potassium, strontium, lanthanum, cadmium, lead, copper, iron) are substituted into the lattice at calcium sites, and anions (F⁻, Cl⁻, SiO4⁻, CO2 3 ) are substituted into the lattice at hydroxyl or phosphate sites. Silicon (Si) substitution has been documented favorable for growth and repair of cartilage and bone because of Si release or change in the crystallographic orientation and surface chemistry of SiHA (Carlisle, 1976; Nielsen and Poellot, 2004). The in vivo chemical and biological interaction is enhanced by the strontium (Sr) substitution in the HA lattice, while magnesium substitution has been reviewed to increase the remineralization phenomenon of the calcified tissues (Wong et al., 2004; Legeros, 1993). Zinc (Zn) substitution in HA is of prime importance because of its cell enzymatic reactions, osteoclasts inhibition, and antibacterial/antifungal nature. It is also found in trace elements in human tissues as it is toxic in abundant amount (Chen et al., 2012). Fluoride (F) substitution enhances biological interaction by improving cell attachment and proliferation (Cheng et al., 2005).

two are above and below CaI and one oxygen triangle is located along the c-axis. The hydroxide ions are positioned parallel to c-axis (Fig. 2.1). Stoichiometric HA has a calcium-to-phosphate ratio of 1.67, and unit cell dimensions of a ¼ b ¼ 9.432Å and c ¼ 6.881Å (Mostafa and Brown, 2007). This chapter focuses on different analytical tools used to investigate different substitutions and their effect on properties of HA.

2.2

Fourier transform infrared spectroscopy

FTIR spectrometer, which gives emission or absorption spectrum of materials, simultaneously collects high-spectral-resolution data over a wide spectral range with better sensitivity and speed. It has been extensively used to chemically characterize the natural mineralized tissues as the frequencies of several vibrational modes of organic and inorganic molecules are active in the infrared region (Lopes et al., 2018).

2.2.1 Silicon substitution Si content in HA lattice (Aminian et al., 2011) makes it a significant material to be used in bone tissue engineering and as bioactive bone substitute. Si is well known for its

Analytical tools for substituted hydroxyapatite

23

promising role in bone and cartilage growth. Precipitation of calcium phosphate due to a stimulus provided by solubilized Si species leads to increase in bone mineral density. Human osteoblast cells function to synthesize type I collagen in the presence of Si (Marchat et al., 2013). Si induces changes in HA properties showing enhanced biological properties along with stoichiometric properties as well as effects on physiological properties of connective tissues and bone (Aminian et al., 2011; Khan et al., 2014). Deficiency of Si results in abnormal bone formation. It has significance because of its role in connective tissue metabolism, and its biological unavailability leads to wear and tear operation and proliferation of bone-forming cells (Khan et al., 2014). This substitution of Si serves as a bioactivity enhancer, when substituted in the crystal 4 lattice of HA, replacing phosphate (PO3 4 ) by silicate ions (SiO4 ) results in increased bone formation and enhanced osteoblast cell activity (Marchat et al., 2013; Zhang et al., 2014). In stoichiometric HA, the sites of phosphate ion are substituted by  Si-forming SiO4 4 with its negative charge stabilized by OH ion vacancy (Khan 4 et al., 2014). The same SiO4 substitution was also confirmed by Tian et al. in 2007 (Tian et al., 2008). The poorly formed crystalline surface on HA is also enhanced because of substitution of Si ions (Marchat et al., 2013). It was found that CO2 present 3 in solution incorporates carbonate (CO2 3 ) groups on PO4 sites and reduces Si substitution (Aminian et al., 2011). Much research has been conducted on the substitution of Si into HA lattice; however, the most cited mechanism by Gibson et al. highlights the creation of anionic vacancies at OH sites due to overall decrease of OH groups per unit cell. The Si substitution affected the PeO vibrational bands also indicating the 3 placement of SiO4 4 at PO4 sites, where infrared peaks at 692, 840, 890, and 1 945 cm were attributed to the Si substitution in HA lattice with the formula Ca10(PO4)6-y(SiO4)y(OH)2-y(VOH)y, where y denotes the molar number of silicate groups (Gibson et al., 1999). Tian et al. reported decrease in relative intensity of  1 PO3 4 and OH bands. The SieO vibrational band appears at 880 cm , which corre4 1 sponds to SiO4 group, while the band at 820 cm is specific for SieOeSi vibrational mode (Tian et al., 2008). Kim et al. observed change in the stretching bands of OH at 630 cm1 and 3570 cm1 (Kim et al., 2003). In one study, the IR spectra of SiHA (with Si 1.25 mol) revealed the decrease in the intensity of vibrational and stretching vibrations of OH groups at 630 and 3570 cm1, respectively (Marchat et al., 2013). 2 4 PO3 4 group also reduces because of CO3 and SiO4 ions, which leads to increase in Ca/P ratios (Bang et al., 2015).

2.2.2

Strontium substitution

Sr, in trace amount in body, is known as resorption and preosteoclast maturation inhibitor, for increasing osteoclast apoptosis, collagen synthesis, cell proliferation, and thus, decreasing bone resorption and maintaining bone formation (Marques et al., 2016). Sr-substituted HA materials show enhanced biocompatibility, bioactivity, osteoconduction, mechanical, and degradation properties (Zhu et al., 2018; Kaygili et al., 2015). Sr-doped HA finds its applications in drug delivery systems, as anticancer drugs, and as bone substitute (Ran et al., 2017). In case of Sr-HA samples, PO3 4

24

Handbook of Ionic Substituted Hydroxyapatites

and OHvibrations show good agreement to characteristic peaks of HA reported in literature (Ran et al., 2017). Bigi et al. reported a linear shift of bands observed in FTIR spectra of Sr-substituted HA. The OH stretching band was shifted from lower (3572 cm1) to higher wavenumber (3590 cm1), while the OH vibration band shifted from 630 to 537 cm1 with the incorporation of Sr (Fowler, 1974). Sr substitutions effect the phosphate bands to lower wavenumbers depending on Sr concentration. Broad v3 phosphate peaks at 1092 cm1 and 1034 cm1 were moved to 1076 cm1 and 1028 cm1. Because of Sr, very weak v2 signal appears at 463 cm1 instead of 472 cm1 present in Ca100 (Frasnelli et al., 2017). Bigi et al. synthesized CaeMg HA with Sr molar ratio from 0 mol (Sr0) to 1 mol (Sr100). It was observed that the phosphate group stretching and bending bands shifted from higher (962, 603, 565 cm1) to lower wavenumbers (947, 592, 561 cm1) (Bigi et al., 2007). Badraoui et al. also reported increase in OH stretching and decrease in out-of-plane band shift due to larger radius of Sr ions (Badraoui et al., 2009). Lowry et al. reported 3   CO2 3 substitutions at PO4 and OH sites in HA (Lowry et al., 2018).OH groups 1 1 related to HA lattice at 3572 cm and 633 cm are gradually diminished depending on Sr concentration (Frasnelli et al., 2017).

2.2.3

Magnesium substitution

Magnesium (Mg), because of its presence in calcified living tissues, might enhance bioactivity and biocompatibility of ceramics. Mg occupies 60%e65% of human body’s Mg in teeth and bones, while the rest of Mg is scattered in nervous, muscle, and other soft tissues and fluids in the body. Mg regulates muscle activity and nervous system and controls heart rate (Montoya-Cisneros et al., 2017). It reduces risk of osteoporosis and ensures recovery of damaged bone. Mg deficiency leads to loss of bone, decreases bone mass as a result of increased osteoclasts, and decreased osteoblasts (Stipniece et al., 2014). Presence of Mg facilitates adsorption of cells and proteins on biological apatites and effect HA formation, crystal size, and its morphology (Matsunaga, 2008). The presence of Mg in Mg-doped HA is confirmed through stretching mode of OH group at 3698 cm1; however, there was no other significant changes in the peaks (Farzadi et al., 2014). Mg caused changes in PO3 4 absorption band in the form of broadening and splitting of this group (Stipniece et al., 2014). Therefore, extra peaks ascribed to vibration modes of phosphate characteristic of Mg-bTCP were detected at 945, 974, 988, 1117, and 1152 cm1.

2.2.4

Zinc substitution

Zn is an important element for bone tissues and is an abundant trace metallic component of bone, promoting bone formation and suppressing its resorption. Zn2þ promotes cell differentiation and proliferation of osteoblast and reduces inflammatory response (Walczyk et al., 2016). The IR spectrum of Zn2þ-substituted HA is shown in Fig. 2.2. Xiao et al. reported that Zn2þ-substituted HA had lower OH stretching vibrational

Analytical tools for substituted hydroxyapatite

25

Zn25

Transmittance (%)

Zn20 Zn15 Zn10 Zn5 CaHap 196

3600

3200

1600

1200 800 Wavenumber (cm–1)

Figure 2.2 Infrared spectra of hydroxyapatite (HA) and Zn(II)-doped HA (Guerra-L
opez et al., 2015).

mode that might be due to weak hydrogen bonding between OH and OH or PO3 4 , respectively (Xiao et al., 2008). The OH stretching band was shifted to lower number (3572e3569 cm1) while the OH vibrational mode shifted from lower to higher wavenumber (600e633 cm1) with increasing Zn2þ concentration (Guerra-L
opez et al., 2015). Kohli et al. reported that Zn2þ concentration effects the relative intensities  of PO3 4 and OH groups. This behavior may be to maintain electroneutrality by inducing loss of OH in HA lattice (Kohli et al., 2014). Bands broadening was observed in the region of 1100e1000 and 565 cm1, while the intensity of the bands at 476 and 963 cm1 decreased with increasing Zn2þ concentration (Guerra-L
opez et al., 2015). Kohli also reported a decrease in the intensity of peaks at 565, 602, and 1039 cm1 as a result of lowering of crystallinity due to increased amount of Zn2þ (Kohli et al., 2014; Thian et al., 2013). The additional bands around 1430 and 941 cm1 are due to precipitation of a Zn impurity phase, probably ZnNH4PO4.H2O (Guerra-L
opez et al., 2015). Esfahani et al. reported the transmission of PeO(H) band intensity at 875 cm1due to Zn content. The increasing Zn2þ fraction also changes PeO band shape that might be due to decrease in crystallinity of HA (Esfahani et al., 2016).

2.2.5

Fluoride substitution

Fluoride (F) ions are required for normal skeletal and dental growth, thus preventing from dental caries risk (Bertoni et al., 1998; Jha et al., 1997). F-substituted HA is used for stimulatory purposes in clinical restorative materials and to maintain material stability during the process (Zhang and Zhu, 2006). Partial F substitution of OH gives F-substituted HA, while complete substitution of OH by F ions results in

26

Handbook of Ionic Substituted Hydroxyapatites

formation of fluoroapatite, a material found in dental enamel, which is generally used in dental applications due to good mechanical properties (Wei et al., 2003). Stretching vibrations of OH were observed at 3570 cm1, which were decreased in intensity as F ions were incorporated at OH sites (Bianco et al., 2010). Wei and Evans also observed absence of OH stretching mode in case of fluoro-substituted HA at 3570 cm1 (Wei et al., 2003). New bands at 3543, 727, and 719 cm1 were observed only in F-doped HA samples, and the bandwidth was directly proportional to the concentration of F. The intensities of the bands at 600, 567, and 474 cm1did not show any increase or decrease, revealing no change by F incorporation (Bianco et al., 2010). Baumer et al. said that the amount of F ions can be estimated in samples on the basis of number, position, and relative integrated absorbance of OH peak in the region 3480e3650 cm1 (Baumer et al., 1985).

2.2.6

Iron substitution

Fig. 2.3 shows the FTIR spectra of iron-substituted hydroxyapatite (Fe-HA) nanoparticles. The basic apatite structure in Fe-HA is confirmed by detection of PO3 4 and OH absorption bands (Li et al., 2009). In case of Fe-HA, additional peaks of adsorbed water appeared in the sample (Kamal and Hezma, 2015).

v3PO4

3–

v4PO4

3–

fFe-HA20% Transmittance (a.u.)

eFe-HA15% dFe-HA10% cFe-HA5% bFe-HA1% aHA(pH5) v3CO3 –

H2O bend

OH stretch

2– 2–

v2CO3

H2O stretch

v2PO4 v1PO4

4000

3500

3000

2500

2000

1500

1000

3–

3–

500

Wavelength (cm–1)

Figure 2.3 Fourier-transform infrared spectroscopy (FTIR) spectra of Fe3þ-doped hydroxyapatite (Li et al., 2009).

Analytical tools for substituted hydroxyapatite

2.2.7

27

Silver substitution

It has been reported in previous studies that incorporation of silver (Ag) in the HA crystal lattice occupied CaI site rather than CaII sites as reported in Fig. 2.1 (Rameshbabu et al., 2007; Badrour et al., 1998). However, Singh reported that CaII sites were more preferably occupied by Agþ (Singh et al., 2011). FTIR spectra of Ag-doped HA lattice showed that before calcination, Ag caused the OH bands intensities (660e540 cm1) to decrease before calcinations. But, after calcination, no significant difference between peaks of HA and Ag-doped HA was found (Santos et al., 2015). Rameshbabu et al. did not observe Ag phosphateerelated vibration band that may be obscured as a result of intense 1034 cm1 phosphate band (Rameshbabu et al., 2007). It has been presented in literature that new peak of adsorbed water at 1628 and slight shifting of OH peak at 3572 cm1 (Iqbal et al., 2012; Lim et al., 2013).

2.2.8

Carbonate substitution

FTIR spectroscopy of CO2 3 -substituted HA shows substitutions at two possible  1 anionic sites: PO3 and OH . CO2 4 3 ions substitute phosphate ions at 875 cm , 1 showing a less intense peak and at 1440 cm with an intense peak. Substitution of 1 OH site by CO2 3 ion appears at 1633 cm (Swamiappan, 2016; Seredin et al., 2017). Antonakos et al. reported induction of vacancy at OH-site due to CO2 3 substitution at 3567 cm1. In CO2 -substituted HA, reports show decrease in intensity or 3 absence of OH peak due to increase in CO2 content (Antonakos et al., 2007). 3 The progressive increase in v2 and v3 bands with increasing CO2 3 amounts has been reported by Fleet and Liu. According to them, increase in CO2 content decreases 3 the intensity of OH vibration and OH stretch bands at 631 and 3570 cm1, respectively (Fleet and Liu, 2007).

2.3

Scanning electron microscopy

It is the most unique instrument for the analysis and examination of microstructural characteristics of solid materials.

2.3.1

Silicon substitution

Marchat et al. prepared SiHAs with the new aqueous precipitation method with concentration of Si up to 1.25 mol and revealed the round shape of SiHA particles and larger acicular pure HA particles at low magnification: however, at higher magnification, no disordered or amorphous layer of SiHA particles was seen (Marchat et al., 2013). Bremen et al. prepared Si-substituted HA using hydrothermal method and observed decrease in particle size (44e69 nm) in Si-HA as compared with pure HA. However, no change was observed in particle shape after incorporation of Si in

28

Handbook of Ionic Substituted Hydroxyapatites

HA (Aminian et al., 2010). Tsalsabila et al. also reported decrease in crystallite size after substitution of Si in HA, which was found to be 103 nm (Tsalsabila et al., 2018). Yu et al. concluded that the particle size was decreased with increase in Si content. However, it was decreased less after a certain Si value was achieved because of secondary-phase formation (Yu et al., 2017). Ibrahim et al. observed presence of agglomerates of nanosized precipitates covering the surface of HA in case of Si-substituted HA (Ibrahim et al., 2011). Yacoubi et al. reported change in crystal size along a-axis of Si-substituted HA from 41 to 32 Å and from 42 to 25 Å, as the concentration of Si changed from 4 to 8 mol (El Yacoubi et al., 2014).

2.3.2

Magnesium substitution

Field emission scanning electron microscopy (FE-SEM) images of pure HA and Mg-HA (magnesium-substituted HA) have been shown in Fig. 2.4. Pure HA particles are nanosized needles. However, Mg-HA has needle-like and plate-like crystallites (Fig. 2.7) (Stipniece et al., 2014). Farzadi et al. observed numerous small spherical particle agglomerations of Mg-HA nanoparticles ranging in size 70e130 nm form agglomerates. Mg inhibits growth of HA grains, thus imparting them small particle size. Crystal size of HA changed from 45 to 38 nm after substitution with Mg (Fig. 2.8) (Farzadi et al., 2014). Incorporation of Mg reduced length of HA nanorods due to inhibition of crystallization in HA solution (Gayathri et al., 2018).

2.3.3

Zinc substitution

Zn incorporation increased the grain size of pure HA as shown in Fig. 2.5 (Uysal et al., 2014). Ren et al. observed that the particles of Zn-HA had shown typical shape without sharp face angles. Increasing Zn concentration makes particles to be agglomerated with size ranging from 10 to 50 nm (Ren et al., 2009). It has been reported that Zn-doped HA possesses more compact structure, where the lower size of crystals hinders their determination (Walczyk et al., 2016). According to Popa et al., increasing Zn quantity influences the morphology of HA powder giving them elongated shape. Nanometric dimension of Zn allows nanoparticles to become agglomerated. As Zn content increases, dimension of nanoparticles decreases (Popa et al., 2016). Findings of Predoi et al. support these results confirming the effect of Zn concentration on size of particles. Analyzed particles showed agglomerated morphology (Predoi et al., 2017).

2.3.4

Iron substitution

The FE-SEM image of Fe-HA showed particles were elongated spheroid shapes of around 70 nm, similar in morphology to HA regardless of concentration of substitution by metal ions (Li et al., 2009). Kramer et al. studied the morphology of both pure and Fe-modified HA and observed no change in morphology and size of both the samples. Nanoparticles

Analytical tools for substituted hydroxyapatite

29

Figure 2.4 Scanning electron microscopy micrographs of samples: (a) hydroxyapatite (HA) (scale bar: 1 mm) and (b) magnesium-substituted HA (scale bar: 100 nm) (Farzadi et al., 2014).

Figure 2.5 Scanning electron microscopy results of (a) pure hydroxyapatite; (b) 2Zn (scale bar is 2 mm, magnification is 30,000 ) (Uysal et al., 2014).

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Handbook of Ionic Substituted Hydroxyapatites

HA (9–432)

Si1 25HA

Intensity (a.u)

Si1 00HA

Si0 75HA Si0 50HA Si0 25HA

HA 20

25

30

35

40

45

50

2q (degrees)

Figure 2.6 X-ray diffraction patterns of raw silicon-substituted hydroxyapatite powders synthesized at pH 10.8 with different Si concentrations (Marchat et al., 2013).

(a)

Sr15HA Sr5HA Sr10HA Sr18HA Sr26HA Sr34HA

PDF strontium apatite

Counts

20000

10000

0 31

32

33

(b)

33.6

33.4

33.2

33.0

32.8

32.5

32.3

32.1

31.9

31.7

31.5

31.3

31.1

30.9

30.7

30.5

PDF hydroxyapatite

Figure 2.7 X-ray diffraction analysis of the synthesized powders by a thermal reaction: (a) strontium-substituted hydroxyapatite (HA) synthesis comprising the characteristic peaks of Sr apatite (Powder Diffraction File (PDF) 00-033-1348), (b) pure HA comprising the characteristic peaks of HA (Powder Diffraction File 00-009-0432). There are no significant differences in the pattern up to 15wt% Sr for calcium. The peaks shift to lower angles from a content of 18wt% Sr up to higher ratios and broaden and the intensities decrease. Furthermore, a new phase has been built (Abert et al., 2014).

Analytical tools for substituted hydroxyapatite

31

1000 Intensity

Zn25 Zn20 Zn15 Zn10

500

Zn5 Zn3 CaHap

0 20

25

30

35

40

2q (degrees)

Figure 2.8 X-ray powder diffraction patterns of HA and zinc-substituted HA samples in the range 20 < 2q < 45 degrees (Guerra-L
opez et al., 2015).

were in the form of agglomerated particles with rod-like morphology having 100e300 nm length (Kramer et al., 2013). In 2013, Gamal et al. reported decrease in grain size with increasing concentration of Fe in HA structure as an effect of substitution. This change in grain size is related to small ionic radius of Fe as compared with calcium (Gamal et al., 2013). Felsen et al. observed prolonged spheroidal particles of 200 nm to 1 mm of calcium HA necked to each other, while Fe-substituted samples are based on spherical particles having size less than 500 nm. Spherical agglomerates formed channel-like assembly of powders. Increase of Fe content forms irregular-shaped particles with lower particle size distribution from w50 to w500 nm (Trinkunaite-Felsen et al., 2015).

2.4

X-ray diffraction analyses

XRD is nondestructive technique for characterization of crystalline materials. It is used to investigate phases, structures, orientation of crystals, crystallinity, grain size, strain, and crystal defects (Bunaciu et al., 2015).

2.4.1

Silicon substitution

Marchat et al. revealed that with the increase in the Si content, the full width at half maximum (FWHM) lines increased with broad diffraction lines. However, no other disordered phase was observed, and the patterns of pure HA and varying concentration of SiHA matched well as shown in Fig. 2.6 (Marchat et al., 2013).

32

2.4.2

Handbook of Ionic Substituted Hydroxyapatites

Strontium substitution

Previous studies reported that incorporation of Sr into the HA crystal lattice replaced calcium sites and caused a decrease in overall crystallinity; however, b-TCP formation was comparatively increased (Kim et al., 2004; Li et al., 2007). XRD analysis has highlighted specific differences between the peaks of pure HA and Sr-substituted HA. It is clear from Fig. 2.7 shown below that the peaks at 32.20 and 32.93 degrees are present only in pure HA pattern, and peak at 32.60 degrees is present only in Sr apatite, which clearly indicates the pure phases of pure HA and Sr apatite, respectively (Abert et al., 2014). Another study confirmed the replacement of calcium sites by Sr and the linear increase of the lattice constants with increasing strontium concentration (Bigi et al., 2007). This is in agreement with another study by Boanini where XRD pattern, observed at two selected ranges (24e27 and 38e41 degrees) of pure HA and SrHA, showed that the reflections became broader and shifted to smaller angles with Sr incorporation (Boanini et al., 2011). As the ionic radius of calcium (0.099 nm) is smaller than the ionic radius of Sr (0.12 nm), the cell parameters of SrHA are enlarged as compared with pure HA as shown in Table 2.1 (Boanini et al., 2011).

2.4.3

Magnesium substitution

According to literature, there was no significant difference between XRD patterns of pure HA and Mg-substituted HA. However, with increasing Mg concentration, peak shifting of peaks at 26 , 32 , 33 , 34 , and 40 at higher angles was observed, which was a clear confirmation of decrease in the a-axis dimension of the crystalline particles (Stipniece et al., 2014; Sader et al., 2013). Some studies reported decrease in a-axis and increase in c-axis dimensions after Mg incorporation in HA lattice (Fig. 2.1) (Okazaki and Legeros, 1992; Landi et al., 2006).

2.4.4

Carbonate substitution

Swamiappan reported change in both a- and c-axis as a result of CO2 3 substitution indicated by values of lattice parameters (Swamiappan, 2016). In case of CO2 3 substitution, diffraction peaks shift to lower angles with the greatest shift of w0.1 A for crystalline plane with (100), (200), (211), (110) Miller indices. This is an indication Table 2.1 Lengths (shkl) of the perfect crystalline domains calculated using the Scherrer method, and cell parameters of SrHA and HA (Boanini et al., 2011). Sample

s002 (Ao)

s310 (Ao)

a-axis (Ao)

c-axis (Ao)

HA

469 (5)

224 (6)

9.4269 (3)

6.8840 (2)

SrHA

319 (7)

150 (3)

9.457 (1)

6.9158 (6)

Analytical tools for substituted hydroxyapatite

33

2 of disturbance of unit cell dimension by CO2 3 parallel to “a”-axis. CO3 ion substitution slightly increases parameter “a” that is indicated by unit cell parameter of CO2 3 -free standard. A slight increase in a-axis with decrease of c-axis is due to Atype substitution (hydroxide ions replacement), while B-type substitution (phosphate replacement) does vice versa (Kwa
sniak-Kominek et al., 2017).

2.4.5

Zinc substitution

2þ

Zn tends to occupy CaII site as compared with CaI (Fig. 2.1) and favors tetrahedral orientation. However, increasing concentration of Zn2þ changes the structure from Zn-substituted apatite to amorphous type of structure. It was reported that, when Zn was incorporated in HA lattice, some bonds of Ca2þ ions were broken and formed CaO phase, which is clearly visible in XRD pattern of HA doped with Zn2þ. However, this CaO phase was disappeared when F was also incorporated, which clearly suggests that HA doped with both Zn2þ and F enhanced the HA lattice stability (Uysal et al., 2013, 2014). In another study, an increase in the line width with a decrease in the intensities was observed especially in the region of 30e35 degrees, which is a partial confirmation of disordered structure as evident in Fig. 2.8 (Guerra-L
opez et al., 2015). It was documented in literature that Zn doping of HA will have a decrease in both lattice parameters “a” and “c” due to its small size (Uysal et al., 2014). However,
osarczyk et al. stated that surprisingly the “c” will have an increase in dimension Sl

osarczyk comparatively. It could be due to Zn occupying both Ca(I) and Ca(II) sites (Sl
et al., 2005). In another study by Y. Tang et al., it was observed that “c” lattice parameter decreased in size while “a” lattice parameter increased in size with increase in Zn concentration in HA, which was limited to only 15% (Tang et al., 2009).

2.4.6

Iron substitution

Fe-HA shows comparable peaks for metal substitution, and no significant peak position shifting observed. However, structure of HA is not highly modified by this ion exchange process. With increase in substitution concentration, peak intensity decreases due to decreasing crystallinity (Li et al., 2009). Kramer et al. also observed broadening of peak and decrease in peak intensity as a result of Fe2þ substitution (Kramer et al., 2013). Results given by Ereiba et al. also support decrease in crystal size with increase of Fe2þ content (Ereiba et al., 2013). Fe-doped HA results in slight shifting in position of peak with smaller peak intensity when compared with pure HA. Metal cation-doped HA results in small particle size and high structural strain (Gamal et al., 2013).

2.4.7

Silver substitution

XRD of Ag-substituted HA (Ag-HA) shows a slight increase in lattice parameters of aand c-axis (Table 2.2) (Santos et al., 2015).

34

Handbook of Ionic Substituted Hydroxyapatites

Table 2.2 Comparison between the lattice parameters calculated by Rietveld refinement for the synthesized samples after calcination at 1000 C/2 h (Santos et al., 2015). Lattice parameters Sample

a

b

c

V(A3)

Hydroxyapatite (HA)

9.4070

9.4070

6.8722

526.64

Silver-substituted HA

9.4072

9.4072

6.8730

526.70

Ramli et al. synthesized Ag substituted HA using microwave assisted method. XRD peaks of HA-Ag seemed to be broader due to narrow and nanosized size distribution. But this might be an effect of irradiation that caused rapid heating. Rapid precipitation of particles from solution leads to small particle size and narrow distribution of particles. XRD pattern showed presence of Ag þ peak at 2q w 38.2 degrees (Ramli et al., 2011).

2.4.8

Multielemental incorporation analysis

XRD patterns for as-synthesized samples of pure HA, SrHA, MgHA, and ZnHA are shown in Fig. 2.9. The broadening of peaks is evident from the figure in the sequence MgHA > SrHA > ZnHA > HA, which is due to the low degree of crystallinity. As the crystallite size reduces, peaks start to become broader. Although Mg-HA exhibited larger crystallite size, yet it exhibited the broadest peaks. The reason is that Mg2þ ions inhibit nucleation of HA and subsequently crystallization (Fig. 2.9).

HA (09–432 JCPDS)

Intensity (arbitrary units)

ZnHA

MgHA

SrHA

HA

20

22

24

26

28

30

32

34

36

2θ (degrees)

Figure 2.9 Typical X-ray diffraction patterns of characteristic HA peaks for precipitated samples (Cox et al., 2014).

Analytical tools for substituted hydroxyapatite

2.5

35

Differential thermal analysis/thermogravimetric analysis

DTA measures change in heat content or thermal properties of sample indicated by a peak or deflection. The activation energy of reaction changes peak position with heating rate (Kissinger, 1957). TGA estimates the mass of polymeric materials when subjected to controlled temperature as a function of time or temperature. Ranges of temperature of TGA may vary from ambient to more than 1000 C depending on the instrument and material nature. An inert atmosphere is created inside the equipment using purge gas flowing through the balance. TGA measures mass loss and gives information about composition, thermal stability, and extent of cure (Prime et al., 2009).

2.5.1

Magnesium substitution

All DTA curves of pure HA and Mg-HA powders showed an endothermic peak around 100e150 C, which was due to evaporation of absorbed water or carbon dioxide. However, with increasing Mg concentration, a distortion was observed around temperature 650e1000 C, which could be due to the Mg-HA decomposition in Mg-b-TCP phase as shown in Fig. 2.10 (Stipniece et al., 2014). TGA of Mg2þ substitution shows that it increases the total weight loss. Sudden weight loss in case of Mg2þ HA might be an indication of its decomposition into whitlockite. Increase of Mg2þ content decreases the thermal stability of Mg2þ substituted HA (Askari Louyeh, 2017).

2.5.2

Carbonate substitution

DTA analysis of CO2 3 -substituted HA shows appearance of endothermic peak due to 2  CO2 3 apatite decomposition at 977 C (Swamiappan, 2016). TGA of CO3 -substituted HA shows decomposition of carbonaceous compounds at a temperature between 600 and 1000 C. Carbonated HA shows less thermal stability resulting in degradation of DTA, μV

HAp 1.0Mg-HAp 2.0Mg-HAp 3.0Mg-HAp 10.0Mg-HAp

5.00 2.50 0.00 –2.50 –5.00 0

200

400

600

800

1000

1200

Temperature, °C

Figure 2.10 Differential thermal analysis (DTA) curves of HAp, 1.0 Mg-HAp, 2.0 Mg-HAp, 3.0 Mg-Hap, and 10.0 Mg-HAp powders (Stipniece et al., 2014).

36

Handbook of Ionic Substituted Hydroxyapatites

carbonated HA at this temperature range. Ca/P ratio in CO2 3 -substituted HA increases 2 as a result of replacement of PO3 4 by CO3 ions, and decomposition of carbonated HA makes it thermally unstable. The less thermal stability of HA due to CO2 3 ions present in the structure causes it to degrade at higher temperature (Swamiappan, 2016). Parthiban and Ohtsuki reported weight loss due to CO2 3 ion decomposition from 500 to 1000 C (Parthiban et al., 2011).

2.5.3

Multielement substitution

DTA analysis of substituted HA shows that DTA profiles of Mg2þ, Zn2þ, Sr2þ, and Si exhibit three peaks at 58.04, 256.87, and 286.71 C for Mg-CS/HA, at 62.43, 261.26, and 288.46 C for Zn-CS/HA, at 74.71, 251.61, and 342.75 C for Sr-CS/HA, and at 68.58, 264.12, and 283.20 C for Si-CS/HA. The first peaks were related to moisture content present in the samples, while the second peaks were related to the substituting element (Ran et al., 2017). Thermal stability can be investigated using TGA analysis. According to a research done on doped apatite samples, it was seen that evaporation of bound and adsorbed water formed an endothermic slope curve in range 50e70 C. Substitution of HA showed exothermic peaks at 339.44 for Zn2þ-substituted HA, 364.01 for Mg-substituted HA, 492.00 for Si-substituted HA, and at 408.00 for Zn2þ-substituted HA. These peaks were the result of decomposition and slow oxidation of the degraded components and char, respectively. Thermal stability of different substituted HA can be summarized as Si-CS/HA > Zn-CS/HA > Mg-CS/ Ha > Sr-CS/HA (Ran et al., 2017). According to Cox et al., TGA of Zn2þ-substituted HA showed no significant difference. However, the difference in high temperature behavior of Mg2þ-substituted HA may be due to decomposition of nonstoichiometric HA (Cox et al., 2014). Gopi et al. observed no significant weight loss or steep peak associated with decomposition of the substituted elements (Gopi et al., 2014).

2.5.4

Zinc substitution

TG-DTA profiles of Zn2þ-substituted HA show an increase in weight loss with increase in Zn2þ content indicating that Zn2þ might be responsible to transform apatite to TCP. Increasing Zn2þ content increases the amount of both adsorbed and lattice water that effects the lattice parameter a (Miyaji et al., 2005). Lopez et al. observed an increase in mass loss with increasing Zn2þ concentration in the range 0e500 C. Amount of adsorbed water increases with incorporation of Zn2þ in the sample (Guerra-L
opez et al., 2015).

2.5.5

Zr-Ce cosubstitution

Sanyal and Raja confirmed thermal stability of cosubstituted HA due to formation of HA/t-ZrO2 (tetragonal) from 600 C. At higher temperature, monoclinic ZrO2 transforms to t-ZrO2 confirmed by appearance of diffused and broad exothermic peaks of ZrO2 crystallization and formation of t-ZrO2 nucleus, between 350 and 450 C. The peak at 800 C is due to decomposition of CeO2 and formation of t-ZrO2 (Sanyal and Raja, 2016).

Analytical tools for substituted hydroxyapatite

2.5.6

37

Silver substitution

Stability of HA is determined on the basis of slope change that appeared between 700 and 850 C. Agþ -substituted HA shows sharp peak up to 200 C to show adsorbed water. The decomposition temperature of sample increases with addition of Agþ. Substitution of HA with Agþ slightly increases the residual weight (Askari Louyeh, 2017).

2.6

Raman spectroscopy

Raman is a technique that helps to find detailed information about molecular vibrations allowing functional group investigation, bonding types, molecular conformations (Talari et al., 2015), sample’s crystallization state, and molecular environment. Some of the vibrations are Raman-active but inactive in infrared depending on their symmetry. Raman uses reflection method for nondestructive and direct analysis to get sample concentration profile, structure, and chemical composition. Spatial resolution of Raman is 100 times higher than infrared technique (Penel et al., 1997). Raman works on the change in polarization of molecule and has the ability to analyze symmetrical molecules.

2.6.1

Zinc substitution

2þ

Zn , a micronutrient, is an essential trace element that is present in the bone. In vitro and in vivo growth of bone is stimulated as a result of slow release of Zn2þfrom Zn2þsubstituted HA. Zn2þ, as an important factor in functioning of different enzymes, has a major role in medicines. Zn2þ replaces Ca2þ ions in HA lattice, thus improving the bioactivity. Anwar et al. did Raman of Zn2þ-substituted HA to detect substitutions in lattice of apatite and observed similarity in peaks to that of normal HA (Anwar et al., 2016). 1 Lopez et al. observed small shift for v1(PO3 4 ) from 961 to 963 cm as a result of 2þ incorporation of Zn in the structure. The most intense band for asymmetric stretching v3b at 1045 cm1 was slightly shifted to higher wavenumber. The most intense band for bending modes of v2 and v4 vibrations of PO3 4 modes from 600 to 400 cm1 shifts to higher wavenumber in Zn2þ-doped HA. Zn-based samples show blurred structure as compared with fine structure for Ca-HAs (Fig. 2.11) (Guerra-L
opez et al., 2015). Substitution of Zn2þ at interstitial site in hydroxyl channel results in peak near 3411 cm1. The width of peak at 3410 cm1 was obscured, which might be due to Zn2þ substitution at 2b site. Zn may be substituted at nearby CaII site with peak position at 3464 cm1 (Friederichs et al., 2015).

2.6.2

Silver substitution

In case of Agþ-substituted HA, reduction in the intensity of phosphate groups was observed, which were present at 1074, 1046, 960, 429, and 450 cm1 (Askari Louyeh, 2017). Addition of Agþ in HA lattice increases the intensity of doubly degenerate bending peak at position 444 cm1 and originates a new peak at 410 cm1, which indicated presence of AgeO stretching vibrational bond (Singh et al., 2015).

Handbook of Ionic Substituted Hydroxyapatites

Raman Intensity (a.u)

38

Zn20 Zn15 Zn10 Zn5 CaHap 1100 1050 1000 990 980 970 960 950 940

600 550

450

400

Wavenumber (cm–1)

Figure 2.11 Raman spectra of hydroxyapatite (HA) and Zn(II)-doped HA in the range 1100e400 cm1 (Guerra-L
opez et al., 2015).

2.6.3

Magnesium substitution

Mg-substituted HA shows the presence of symmetric stretching OePeO vibrations in the form of an intense band at 910 cm1. Aina et al. reported change in intensity of FWHM of Mg-HA characteristic bands at 910 cm1. Increase in Mg2þ amount leads to formation of TCP phase with appearance of shoulder peak in 924 cm1 band (Aina et al., 2012).

2.6.4

Strontium substitution

Raman of Sr2þ-substituted HA shows that main phosphate band shifts from 910 cm1 to low wavenumber 906 cm1(Aina et al., 2012). Addition of Sr2þresults in main 1 Raman band shift due to v1PO3 and decreases linearly. Same results 4 at 955 cm 1 were reported by Donnell et al., according to which PO3 4 vibrations at 963 cm 2þ and FWHM decrease linearly due to Sr substitution (O’donnell et al., 2008).

2.6.5

Silicon substitution

Substitution of Si brings slight changes in spectra with slight difference in amplitude of OH-stretching vibration at 3570 cm1 (Zou et al., 2005) instead of 3572 cm1 (Koutsopoulos, 2002). Modes of PO3 4 bands that include v1, v2, v3, and v4 bands are assigned at 960, 400e500, 1000e1100, and 550e650 cm1. Si substitution caused broadening in the peaks, while the v2, v3, and, v4 peak intensity was increased with Si addition. An increase in peak amplitude was also observed at 845 and 890 cm1 due to Si (Zou et al., 2005).

Analytical tools for substituted hydroxyapatite

2.6.6

39

Fluoride substitution

In F--substituted HA, Raman spectra show the domination of PO3 4 unit vibrations. Main difference in Raman spectra of pure HA and substituted HA is due to blue shift of v1 peak that is attributed to unit cell volume reduction and the change in splitting v3 vibrations (Campillo et al., 2010). F content is associated with shifting of PeO stretching in PO3 4 band. Increase of F content shifts the band upfield, resulting in shortening of PeO band. The electrostatic attraction of phosphate tetrahedra’s oxygen atoms is increased due to replacement of OH by F ions resulting in increasing vibrational frequency and PeO band shortening. Fluorine increases the full width half height of symmetric phosphate stretch that occurs at 960 cm1 and seems to be associated with crystallinity measurement. Increasing F content also affects the OeH stretch intensity at position 3580 cm1 and broadens the peak with shoulder formation at 3545 cm1. F is responsible for formation of two peaks at 3545 cm1 and 3580 cm1. The intensity of OeH band at 3580 cm1 decreases and finally disappears leaving a weak band at 3545 cm1 (Fig. 2.12) (Chen et al., 2015).

2.6.7

Iron substitution

Fe2þ substitution is responsible for disappearing many peaks in substituted HA and results in formation of new peaks at 224, 244, 291, 409, and 498 cm1, an indication of hematite. In HA, weaker narrow peak of HA was masked by strong hematite peak at 1 612 cm-1. Fe2þ substitution broadens v3PO3 is character4 region. Band at 1320 cm 2þ istic of Fe substitution (Antonakos et al., 2017). Radha et al. reported the appearance of two peaks in Fe2þ-substituted samples at 357 cm1 and 730 cm1, which shows the presence of Fe-bonded oxygen. Increased Fe2þ concentration enhances the HA decomposition rate resulting in broadening of peak at 960 cm1 (Radha et al., 2016).

2.7

Nuclear magnetic resonance

It is the analytical technique that evaluates the structural and chemical composition of any substance through the electromagnetic spectra generated by the nuclei of the substances under consideration in a uniform and strong magnetic field and uses radiofrequencies in the range of 10e1000 MHz (De Graaf, 2019). Proton Magic angle spinning (MAS) NMR spectra of Si-HA has been shown in Fig. 2.13. The significant peak at 0.2 ppm and small insignificant peak at c. 1.6 ppm have been previously filed for hydroxyl groups (Yu et al., 2017). The two wide peaks appeared at c. 5.4 ppm in P/ Si-HA0.8 and P/Si-HA1.6 (silicon-substituted HA) samples. These peaks were attributed due to presence of water molecules on HA surface or eOH groups near to the eOH vacancies; however, the peaks broaden due to the presence of silicon in HA (Hartmann et al., 2001). For the samples synthesized by the precipitation method, the peaks increased with the increase in Si content. There were two sharp resonances observed at c. 72.8 and 69.4 ppm due to the presence of SiO4 4 inside the HA structure (Yu et al., 2017).

40

Handbook of Ionic Substituted Hydroxyapatites

Intensity

(a)

HA FHA1 FHA2 FHA3 FHA4 FHA5 FHA6

900

930

960 990 Raman shift cm–1

1020

1050

3400

3500 3600 Raman shift cm–1

3700

3800

(b)

HA

Intensity

FHA1 FHA2 FHA3 FHA4 FHA5 FHA6 3300

Figure 2.12 Raman spectra of fluorohydroxyapatites containing different fluorine levels (HA: 0 wt%, FHA1: 0.54 wt%, FHA2: 0.83 wt%, FHA3: 1.59 wt%, FHA4: 1.93 wt%, FHA5: 2.2 wt%, and FHA6: 2.94 wt%) in the range of (a) 900e1050 cm1 and (b) 3100e3800 cm1 (Chen et al., 2015).

Frasnelli et al. reported that addition of strontium in HA broadened and shifted the resolved peak at 2.80 ppm to lower fields by 31P SP-MAS NMR. It was also observed that the shifting and broadening of the internal PO3 4 was happened with strontium concentration up to 50%; however, further increase in strontium caused narrowing and shifting of the peak at high field. This could be due to loss of PeO bonds in account for increase in the lattice expansion (Frasnelli et al., 2017). 43Ca solid-state NMR demonstrated that the energetically most preferred site of magnesium incorporation is Ca(II) site. Magnesium incorporation in the HA lattice deforms the structure because of the large space occupied by the MgeO bonds as compared with CaeO bonds. These structural changes will increase the different sites for the PO3 4 and OH anions, in agreement with the experimental 1H and 31P solid-state NMR data (Laurencin et al., 2011). Trandafir et al. reported that the MAS NMR results of the

(b)

(e)

H: HAp1.6

(d) H: HAp0.8

H: HAp

Analytical tools for substituted hydroxyapatite

(a)

(c) P: HAp1.8

(b)

P: HAp0.8

–69.4

–72.8

(a)

P: HAp 10

8

6

4

2

0

–2

–4

12

10

8

6

4

δP / ppm

δH / ppm 1

2

0

–4

0

–20

–40

–60

–80

–100

–120

–140

δSi / ppm

31

Figure 2.13 (left side) Solid-state H (a) and P (b) MAS NMR spectra of HA and Si-HA; (right side) Direct excitation Si29MAS NMR spectra of P: Si-HA1.6 (Yu et al., 2017).

41

42

Handbook of Ionic Substituted Hydroxyapatites

iron-doped HA showed that only a smaller part of the iron atoms were uniformly incorporated in the HA structure in place of calcium, whereas the significant part was scattered in iron-dominating areas under strong dipolar interactions (Trandafir et al., 2014).

2.8

In vivo/in vitro analysis

The advancement of solid trial models for the clinical utilization of biomaterials and in foreseeing expected results is winding up progressively (Fini and Giardino, 2003). Experimenting on a living organism refers to in vivo analysis covering animal studies and clinical trials. However, performing experimentation in a controlled environment outside the living body refers to in vitro analysis, which remained a prime analysis until the last few years. Yu et al. demonstrated the dissolution of Si-HA at the grain boundaries, and it was noticed that silicate incorporation in HA lattice increased the solubility in SBF (simulated body fluid) in the following order P: Si-HA1.6 >P: Si-HA0.8 > P: HA (Yu et al., 2017). Pietak et al. demonstrated that in vitro, aqueous Si effects are dose-related, and the release of Si complexes enhance tissue regeneration and biomineralization, which was also in line with the in vivo effects of SiHA observed by Patel and coworkers (Ananth et al., 2015). Fig. 2.14 showed the crystalline apatite deposits on Si-HA surface immersed in SBF for 25 days. The apatite crystal growth is increased with increase in Si content in HA. Strontium substitution in HA was tested in vitro, and MTT analysis was done at various concentrations of strontium. A statistically significant difference was observed between the HA and SrHA, and the cell viability was increased up to 30% with strontium incorporation. Cell proliferation was also enhanced with increasing strontium concentration in HA (Frasnelli et al., 2017). In a study performed by Landi and co-workers demonstrated that in vitro, 5.7 mol% Mg-incorporated HA showed increased bone breakdown with the passage of time compared with HA. It was highly biocompatible with no carcinogenicity, cytoxicity, genotoxicity, and skin sensitivity. In vivo testing of Mg-doped HA was also done by incorporating the MgHA granules in the femur bones of white rabbits, and it was observed that boneematerial contact increased progressively over the time compared with HA (Landi et al., 2008). Ereiba observed the effect of iron on the bioactivity of HA in vitro. Different concentrations of iron-doped HA were prepared by wet chemical method. It was concluded after characterizing the samples that iron incorporation in HA increased the dissolution rate of calcium, phosphorus, and iron itself and formation of apatite layer on the surface (Ereiba et al., 2013). In another in vivo study, Fe-HA sample was implanted in the sheep animal model to check the tissue response and degradation process. X-ray radiogram analysis clearly showed the continuous degradation and constructive tissue response even after 60 days of implantation (Ulum et al., 2014). Chen tested in vitro and in vivo analysis of silver-doped HA and reported that there was a significant reduction in the number of Gram-positive bacterial colonies on Ag-HA surface comparatively; however, there was no statistically significant difference in the in vitro cytotoxicity observed between HA and Ag-HA surfaces

Analytical tools for substituted hydroxyapatite

43

(a)

(b)

2.00μm

(c)

2.00μm

(d)

2.00μm

(e)

2.00μm

(f)

2.00μm

2.00μm

Figure 2.14 Scanning electron microscopy images of hydroxyapatite (HA) and Si-HA after 25 days in SBF (a) P: HA; (b) P: Si-HA0.8; (c) P: Si-HA1.6; (d) H: HA; (e) H: Si-HA0.8; and (f) H: Si-HA1.6 (Yu et al., 2017).

(Chen et al., 2006). The increased biocompatibility and reduced bacterial growth by Ag-HA was tested by Jelinek et al., where they followed ISO 10993-5:2005 (Tests for cytotoxicitydIn Vitro Methods) and ISO 7405:2008 (DentistrydEvaluation of biocompatibility of medical devices used in dentistry) standards to check biocompatibility direct contact test (Jelinek et al., 2013). Zinc has been found to enhance the new bone growth. In another in vivo case study, Zn-doped HA was implanted in the tibia of adult rabbit to analyze the tissue response and healing mechanics. Laboratory characterizations and evaluations were performed up to 2 months postoperatively in which it was clearly demonstrated that Zn-substituted HA showed enhanced osseointegration than simple HA (Bhattacharjee et al., 2014). Copper-substituted HA has been proven to exhibit better antifungal properties than Zinc-substituted HA (Stani
c et al., 2010).

44

2.9

Handbook of Ionic Substituted Hydroxyapatites

X-ray fluorescence

XRF (X-ray fluorescence) is a nondamaging characterization technique used to analyze the chemical structure and elemental composition of materials. The amount of zinc in the zinc-doped HA was quantitatively confirmed by XRF. The XRF results showed that the zinc content of the samples is lower than those of corresponding amount of starting material (Venkatasubbu et al., 2011). This implies that some of the zinc ions remain in the mother solution after precipitation and it decreased the Ca/P ratio. In another study, Mg-doped HA with Mg content from 0 to 20 mol.% was prepared. The atomic concentrations of the ions in the Mg-HA samples were assessed by XRF, and the molar ratio in the final product was reduced to 1/3 part to that of initial value in the solution affecting stoichiometry of HA (Ren et al., 2010). HA was doped with strontium, and X-ray fluorescence (XRF) results decided the Ca/P ratio to be 1.66. The ratio of Ca2þ/Sr2þ was also confirmed via X-ray photoelectron spectroscopy (XPS) to be 19.43%e1.99%, respectively, in the strontium-doped HA with total amount of Sr2þ 10%, substituted for the total Ca2þ content. This ratio of 20:2 is explanatory for the near-complete substitution of the Sr2þ ions into the HA lattice replacing the Ca2þ ions (Brook et al., 2012).

2.10

Conclusion

Analytical tools are important for material characterization to understand their properties. In this chapter, we focused on characterization of substituted HA using the techniques FTIR spectroscopy, XRD, SEM, DTA/TGA, Raman spectroscopy, solidstate nuclear magnetic resonance (MAS NMR), XRF, and in vivo/in vitro analysis, which helped to detect changes between pure HA and substituted HA. It is observed from the literature that the single- or multicationic and anionic substitutions (Ag2þ, 3 Mg2þ, Zn2þ, Co2þ, Sr2þ, Si2þ, F, CO2 3 , PO4 ) influence the crystallographic, chemical, biological, morphological, thermal, mechanical, and physical properties of synthetic HA.

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Bioceramics: types and clinical applications

3

Hashmat Gul 1 , Maria Khan 2 , Abdul Samad Khan 3 1 Department of Dental Materials, Army Medical College, National University of Medical Sciences, Islamabad, Pakistan; 2Department of Oral Biology, University of Health Sciences, Lahore, Pakistan; 3Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia

3.1

Introduction to bioceramics

Ceramics utilized in the repair and reformation of damaged or diseased parts of musculoskeletal system of the body are termed bioceramics. Bioceramics comprise of alumina, silica, glass ceramic, titanium dioxide, zirconia, calcium phosphates, and bioactive glass (Hench, 1991). Because of their positive interactions with human tissues, bioceramics can be utilized in numerous biomedical applications, including replacements for teeth, knees, hips, ligaments, and tendons, spinal fusion, repair of diseased periodontium, maxillofacial reconstruction, stabilization and augmentation of jaw bones, as temporary bone space fillers after tumor excision, and as thromboresistant carbon coatings in prosthetic heart valves (Dearnaley and Arps, 2005; Eslami et al., 2018; Dorozhkin, 2018).

3.2

Classification of bioceramics

Bioceramics have been classified based on origin, tissue response, composition, and crystallinity. The summary of this classification is tabulated in Table 3.1, where detailed description is discussed in the following sections.

3.2.1

Classification on basis of origin

Bioceramics can be obtained from natural and synthetic sources. Those bioceramics that occur naturally and originate from various living or dead organisms can be termed “Natural Bioceramics.” Most of natural bioceramics come from animals and include biogenic silica, natural pearls, mollusk shells, bones, and teeth (Brundavanam et al., 2017). However, the bioceramics that are synthesized artificially are termed “Synthetic Bioceramics,” which include calcium phosphateebased materials, hydroxyapatite (HA), zirconia, alumina, bioglass, etc. (Chevalier and Gremillard, 2009). HA is a calcium phosphateebased mineral, which is not only produced synthetically but also present naturally in bone and teeth. The detailed information about

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00003-3 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Table 3.1 Classification of bioceramics. Origin

Type of tissue response

Composition

Crystallinity

Natural

Biogenic silica Mollusk shells Natural pearls Teeth Bones

Synthetic

Calcium phosphateebased materials Hydroxyapatite (HA) Bioglass

Bioactive

Calcium phosphateebased materials HA Bioglass

Bioinert

Alumina Glassy carbon

Aluminum based

Alumina Aluminosilicates

Zirconium based

Zircon Cubic zirconia Tetragonal zirconia

Carbon based

Graphite Vitreous carbon

Calcium phosphate based

Amorphous calcium phosphate b -tricalcium phosphate Biphasic calcium phosphate HA Fluorapatite

Silica based

Tricalcium silicate

Crystalline

Aluminosilicates HA Fluorapatite Zirconia

Amorphous

Amorphous calcium phosphate Bioglass

biological apatite or natural HA is given in Chapter 1. HA is researched extensively for its potential to replace viable bone material in numerous clinical procedures (Brundavanam et al., 2017; Habraken et al., 2007; Hutmacher et al., 2007). Nacre, which is the inner lustrous coat of pearl oyster shell and abalone shells, constitutes crystalline calcium carbonate mineral phase in organic matrix. Good mechanical properties and

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55

biocompatibility of nacre with bone tissues (Lamghari et al., 1999; Mouri
es et al., 2002) mark its potential for bone tissue engineering. Biogenic HA has been obtained from fish bone (Mondal et al., 2019), sheep bone (Triyono et al., 2018), goat bone (Aarthy et al., 2019), chicken beak (Alshemary et al., 2018), porcine bone (Sossa et al., 2018), etc., whereas synthetic bioceramics can be synthesized by various techniques including coprecipitation method (Aruna et al., 2018), solegel (Khan et al., 2008), microwave irradiation technique (Khan et al., 2017; Khalid et al., 2018), evaporation-induced self-assembly (Shah et al., 2018), hydrothermal (Ali et al., 2019), continuous hydrothermal flow method (Chaudhry et al., 2008), etc.

3.2.2

Classification on basis of type of tissue response

In late 1960’s, bioceramics were first used as substitutes to metallic materials to upsurge biocompatibility of implants. On basis of host tissue response, bioceramics can be classified as relatively bioinert ceramics and bioactive or surface reactive ceramics (De Aza et al., 2005; Utneja et al., 2015). Bioceramics that do not provoke any tissue reactions on interaction with physiological system are termed “Bioinert,” e.g., alumina (Rokusek et al., 2005) and zirconia (Miao et al., 2007), where zirconia-based material is now used as dental implants, placed mostly in anterior jaw segments because of better aesthetics, lesser bacterial adhesion, nonallergic nature, and improved osseointegration than conventional titanium implants (Kanchana and Hussain, 2013). Bioceramics having a potential of inducing a particular tissue response on interaction with physiological system are termed “Bioactive,” e.g., in surgical practice, the most popular bioactive ceramics are calcium phosphateebased materials (Pina and Ferreira, 2012; Hench, 1991). Bioactive glass exhibits inherent osteogenic potential and is capable of regenerating diseased or lost bone. Bioactive ceramics may further categorize into “Resorbable” and “Nonresorbable” on basis of their solubility, e.g., HA when used as bone filling material heals bony defects via new bone formation on dissolution over time (Pina and Ferreira, 2012; Hench, 1991). Solubility of HA is dependent on its degree of porosity. Tricalcium phosphates (TCPs) are extremely resorbable bioceramics when used as bone fillers in powdered or compact form. Mother of pearl and coral are calcium salts of animal origin, composed chiefly of aragonite calcium carbonates, and are highly resorbable (Dubruille et al., 1999). Nonresorbable ceramics include metal oxides such as zirconia and alumina, which are not only used as bone fillers but also as dental implants (Dubruille et al., 1999; Siddiqi et al., 2017).

3.2.3 3.2.3.1

Classification on basis of composition Alumina-based bioceramics

Alumina, the oxide of aluminum, comes as fine white powder and contains traces (0.6%) of the oxides of magnesium, silicon, sodium, iron, and calcium. Alumina is a brittle solid and fractures on receiving a direct blow and exhibits hardness slightly

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less than diamond, making it impossible to cut or trim with any other metal except diamond. It exhibits poor flexural strength, however, excellent compressive strength and excellent dimensional stability against heat, pressure, and shock (Boutin, 2014). Exceptional hardness of alumina lead it to its use in cutting and milling tools (Boutin, 2014). To develop all ceramic prosthesis, new aluminum-based ceramic systems with greater strength and translucency have been developed constituting high-purity alumina, glass-infiltrated, sintered alumina, and leucite-reinforced ceramic cores € (Ozcan and Vallittu, 2003).

3.2.3.2

Zirconium-based bioceramics

Zirconia is glass-free polycrystalline ceramics in which all atoms are packed into regular crystalline arrays. Zirconia crystals at different temperatures exist in three crystallographic forms, i.e., monoclinic, cubic, and tetragonal. At firing temperature, zirconia is in tetragonal phase, which on cooling to room temperature changes to monoclinic phase. As space occupied by a unit cell of monoclinic is 4.4%, greater than tetragonal, therefore, zirconia crumbles on cooling (Badami and Ahuja, 2014). In late 1980s, stabilization of tetragonal form of zirconia at room temperature was made possible by addition of small amounts of calcium initially and later yttrium or cerium as phase stabilizers (Kelly and Denry, 2008). Metallic oxides are used as molecular phase stabilizers; therefore, zirconia can be classified as magnesiumstabilized zirconia, calcium-stabilized zirconia, and yttrium-stabilized zirconia (Piconi and Maccauro, 1999, EL-Ghany and Sherief, 2016). Tetragonal zirconia is “metastable,” and a highly localized stress before crack propagation is enough to trigger grains of ceramic to transform from tetragonal to monoclinic phase, bringing 4.4% increase in crystal volume at the site of crack tip, which seals the crack and minimizes further crack propagation (Gupta et al., 1977; Badami and Ahuja, 2014). Zirconia exhibits mechanical properties matching stainless steel with good capacity to tolerate cyclic stresses in addition to 900e1200 MPa tensile strength and 2000 MPa compression strength (Piconi and Maccauro, 1999). Surface treatments of zirconia can affect its physical properties such as surface grinding reduces its toughness and aging in wet environment can have detrimental effect on its properties (Swab, 1991; Luthardt et al., 2002; Kosmac et al., 1999). Enhanced material strength, heightened esthetic, and high biocompatibility increases the possibilities of zirconia ceramic to be utilized in many clinical applications (EL-Ghany and Sherief, 2016). Only after the availability of computer-aided manufacturing (CAM), manufacturing of well-fitting prostheses made from zirconia became practical (Badami and Ahuja, 2014).

3.2.3.3

Carbon-based bioceramics

Carbon exists in several forms, i.e., diamond, graphite, vitreous carbon, amorphous carbon, and pyrolytic carbon. Forms of carbon employed in biomedical devices include graphene, diamond-like carbon, carbon nanotubes, and turbostratic carbons (De Aza et al., 2005). Some forms of carbon display good chemical inertia, biocompatibility, and thromboresistance, in addition to their physical characteristics close

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57

to the bone. Their densities and elastic modulus range between 1.5e2.2 g/cm3 and 4e35 GPa, respectively (De Aza et al., 2005).

Turbostratic carbons Turbostratic carbons includes low temperature isotropic (LTI) pyrolytic carbon, ultra-low temperature isotropic (ULTI) pyrolytic carbon, and vitreous carbon. Turbostratic carbons and vitreous carbons exhibit very complicated disordered lattice structures, closely similar to graphite structure, but with arbitrary oriented layers (De Aza et al., 2005). Silicon-alloyed LTI pyrolytic carbons exist in crystalline sizes of 3e5 nm (De Aza et al., 2005) with density ranging between 1700 and 2200 kg m
3, expansion coefficient of 5e6  10
6 K
1, hardness value of 230e370 DPH, Young’s modulus 27e31 GPa, flexural strength 350e530 MPa, fracture strain 1.5%e2%, and fracture 1 toughness 0.9e1.1 mPa m /2 (De Aza et al., 2005). Pyrolitic carbon is normally used as implant coating made from a hydrocarbon gas in a fluidized bed employed via chemical vapor deposition method (L opez-Honorato et al., 2009). However, ULTI vapor deposited carbon exists in crystalline sizes of 8e15 nm. ULTI carbons exhibit density ranging between 1500 and 2200 kg m
3, hardness value of 150e250 DPH, Young’s modulus 14e21 GPa, flexural strength 345e690 MPa, and fracture strain of 2%e5% (De Aza et al., 2005). Because of better thromboresistance and biocompatibility with soft tissues and blood, pyrolytic carbons are used primarily as coatings in circulatory apparatus, blood vessel, and cardiac-valve prosthetic devices. Modern heart valves are mostly coated with LTI pyrolytic carbon, where doping of LTI pyrolytic carbon with silicon improves its mechanical properties, minimizing degradation of prosthetic valves by cyclic fatigue and possible erosive cavitation during patient’s life (De Aza et al., 2005).

Graphene Among various allotropes of carbon, graphene is an innovative nanomaterial, the possible usage of which is being explored in the field of biotechnology and biomedical sciences and biotechnology. Graphene-based materials have the potential to be implemented as biosensors, drug delivery, and tissue engineering (Saghatforoush et al., 2014; Hasanzadeh et al., 2014), where they exhibit incomparable electron transport, high surface area, mechanical, and physicochemical characteristics (Omidinia et al., 2013; Hasanzadeh and Shadjou, 2013). Moreover, graphene raw materials have minute metallic impurities than other carbon nanomaterials (Fan et al., 2010; Zhang et al., 2014).

Diamond-like carbon Diamond-like carbon is a black, amorphous solid with hardness value between diamond and graphite. Its microstructure permits incorporation of other species such as hydrogen, nitrogen, sulfur, silicon, titanium, tungsten, or silver to modify its properties (Dearnaley and Arps, 2005). Up to 40% atm. hydrogen in regions of low electron density of diamond-like carbon strongly effects the tribological and mechanical behavior of diamond-like carbon surface coatings (Dearnaley and Arps, 2005).

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Several methods developed for deposition of diamond-like carbon coatings from numerous carbonaceous precursors includes pulsed laser ablation, direct ion beam deposition, ion beam conversion of condensed precursor, filtered cathodic arc deposition, plasma source ion implantation and deposition, radiofrequency plasma-activated chemical vapor deposition, and magnetron sputtering (Bewilogua and Hofmann, 2014; Dearnaley and Arps, 2005). Diamond-like carbon is a corrosion resistant and hemocompatible biomaterial. It also exhibits superior resistance to Streptococcus aureus adhesion and colonization than generally used metallic biomaterials such as titanium, tantalum, and chromium in presence of serum (Levon et al., 2010). It is clinically tested as stents and mechanical heart valves (Dearnaley and Arps, 2005), whereby amorphous diamond-like carbon can be modified by doping or alloying with various elements. Silicon- and fluoride-doped diamond-like carbon coatings on medical guidewires enhanced their substrate adhesion, surface smoothness, and lubrication behavior, which reduced thrombogenicity and polishing scars (Maguire et al., 2005; Hasebe et al., 2006).

3.2.3.4

Calcium phosphateebased bioceramics

Calcium phosphates are found naturally abundant in earth crust as white solids but acquire the lattice structure of its constituent elements (Pina and Ferreira, 2012). Calcium phosphates are bioactive, bioresorbable materials and are similar to the inorganic part of teeth, bones, and antlers (Pina and Ferreira, 2012). Calcium phosphates have a tendency to establish chemical bonding with bone and adjacent tissues; however, they also aid in providing good stabilization to mechanical load-bearing materials (Pina and Ferreira, 2012). Calcium phosphates are used extensively for bone tissue engineering both in dentistry and orthopedics (Eliaz and Metoki, 2017) and require up to 70%e80% porosity with full connectivity (Best et al., 2008). HA and b-TCP are most popularly used bioceramics of calcium phosphate origin (Pina and Ferreira, 2012). Commercial calcium phosphate products of natural origin derived from bovine bone include Bio-Oss (Geistlich Biomaterials, Geistlich, Switzerland), Osteograf-N (CeraMed Co Denver, CO), Endobone (Merck Co, Darmstadt, Germany), Trubone, and BonAp (Legeros, 2002). Biphasic calcium phosphates are available commercially as Triosit, MBCPTM, and Osteosynt (Legeros, 2002). Various calcium phosphateecontaining compounds have been considered as possible starting reagents for calcium phosphate cements. These cements are based on tetracalcium phosphate (Ca4(PO4)2O), monocalcium phosphate monohydrate (Ca(H2PO4)2$H2O), b-TCP (b-Ca3$(PO4)2), a-TCP (a-Ca3(PO4)2), and octacalcium phosphate (Ca8H2(PO4)6$5H2O) (Suzuki et al., 2006). Commercial calcium phosphate cements include Bonesource (Orthofix Inc, McKinney, TX), Cementek, and aBSM (Etex Corporation, Boston, MA) (Legeros, 2002).

Hydroxyapatite Octacalcium phosphate has been considered as a precursor of biological apatite crystals of bone and teeth. Its structure stacks apatitic layers alternatively with hydrated layers, and the transition of octacalcium to HA is thermodynamically favored.

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59

There is evidence of presence of octacalcium phosphate in the central part of a dentine crystal and apatite in the outer most layers of the same crystal. The resulting biological apatitic crystals are constituted of poorly crystalline HA with a low Ca/P ratio, i.e., Ca-deficient HA containing ions such as carbonate and fluoride (Suzuki et al., 2006). Calcium phosphates with a calcium and phosphate atomic ratio ranging between 1.5 and 1.67 are termed apatites, i.e., HA or fluorapatite (Eliaz and Metoki, 2017). HA is a stable and biocompatible bioceramic for bone tissue engineering (de Lima et al., 2011). HA being highly crystalline and stable with limited solubility in aqueous environment exhibits osteointegration into the newly formed bone on implantation into the body (Pina and Ferreira, 2012). HA may exhibit variable stoichiometry, and on cationic substitutions may modify its surface chemical activity, effecting its biocompatibility. Physicochemical analyses of HA substituted with divalent cations such as cobalt (Co2þ), lead (Pb2þ), copper (Cu2þ), strontium (Sr2þ), zinc (Zn2þ), magnesium (Mg2þ), or iron (Fe2þ) showed that nature of cationic substitution greatly effects characteristics of HA related to crystallinity and calcium release/uptake rates (de Lima et al., 2011). Except Zn-HA, all other substituted HAs induced some degree of apoptosis. Highest apoptosis was observed for Co-HA and Mg-HA. Only Cu-HA impaired simultaneous membrane integrity, mitochondrial activity, and cell density. Highest relative cell densities were observed for Mg-HA and Zn-HA, while Co-HA exhibited improved cell adhesion onto HA surface (de Lima et al., 2011). Physical and chemical modifications of HA implants via incorporation of biological units such as proteins, cells, and growth factors lead to improve osseointegration. Silicate-substituted HA as bone grafting materials have potential of increased bone opposition and organization on implantation (Mankani et al., 2006; Mastrogiacomo et al., 2005; Porter et al., 2005). Synthetic HA is commercially available as Calcitite (Sulzer Calcitek, Carlsbad, CA), Alveograf, Calcitite, and Durapatite, whereas unsintered calcium deficient apatite is available as OsteoGen (Impladent Co, Long Island, NY) (Legeros, 2002).

b-tricalcium phosphate b-TCP is synthesized by sintering precipitates of calcium-deficient apatite with calcium to phosphorous molar ratio of 1:5 or by high-temperature solid-state reactions (Legeros, 2002). Using emulsions of calcium phosphate, porous blocks of b-TCP can be synthesized, with controlled macropore sizes without effecting the microporosity (Bohner et al., 2005). Custom-made macroporous b-TCP bone substitutes can be fabricated using 3D powder printing technique. b-TCP monoliths are produced by rapid prototyping, lessening the processing time for commercial production due to their ease of handling and rapid hardening (Vorndran et al., 2008). b-TCP is resorbed fully over time promoting osteogenesis. Unfortunately, because of poor mechanical properties, b-TCP fails to allow load-bearing applications. However, combining with bioinert polymers, it forms a composite, which facilitates not only chemical bonding with bone but also accelerated implant fixation. The elastic modulus of such composites

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Handbook of Ionic Substituted Hydroxyapatites

can also be simulated to that of human bone by altering the ceramic composition, which favors avoiding stress shielding and succeeding bone resorption (Pina and Ferreira, 2012). Vitoss (Orthovita, Inc, Philadelphia, PA), Augmen, and Synthograf are among the commercially available b-TCP (Legeros, 2002).

3.2.3.5

Silica-based bioceramics

Calcium silicates Biomaterials based on calcium silicate, such as calcium silicate glass, b-calcium silicate, a-calcium silicate, dicalcium silicates, and tricalcium silicates, have the potential of inducing rapid apatite formation in physiological fluid and chemical integration into living bone structure following implantation (Li et al., 2014). Mineral trioxide aggregate (MTA) is a calcium silicateebased bioceramic, mainly composed of a blend of tricalcium silicate, tricalcium oxide, tricalcium aluminate, and silicate oxide (Borges et al., 2017). MTA powder constitutes 50e75wt.% calcium oxide, 15e25wt % silicon oxide, and aluminum oxide mainly with comparatively small amounts of aluminum oxide, iron oxide, and magnesium oxide in white variant (Dammaschke et al., 2005; Torabinejad and White, 1995; Asgary et al., 2006), whereby bismuth imparts radiopacity to MTA cement (Camilleri et al., 2005; Song et al., 2006). MTA particles are mostly irregular with a fraction of elongated particles. Studies have reported that white MTA particles were comparatively finer than gray MTA (Camilleri et al., 2005; Asgary et al., 2006) with particle size ranging from 5 to 50 mm in gray MTA and 5 to 25 mm in white MTA (Asgary et al., 2006). Calcium silicates in MTA on hydration undergoes hydrolysis and yield calcium hydroxide and calcium silicate hydrate (Darvell and Wu, 2011), thus set MTA presents as calcium hydroxide confined within a silicate matrix (Camilleri et al., 2005).

Aluminosilicates Aluminosilicate glasses comprise rare-earth oxides, which display refractive indices of about 1.65, with reasonable coefficients of thermal expansion but very high glass transition temperatures (Shelby and Kohli, 1990). Glass ionomer cements (GICs) are dental restorative materials containing fluoride- and phosphate-enriched aluminosilicate glasses in their powder component, which on reaction with weak polymeric acids sets to form a final structure with large quantity of unreacted glass fillers that reinforce the resulting dental restoration (Sidhu and Nicholson, 2016).

Silicon nitride Silicon nitride, Si3N4, is a biocompatible bioceramic exhibiting favorable cell interactions, osseointegration, and bacteriostatic potentials. High strength and tough implants with minimum wear can be produced with dense polished form. It exhibits exceptional strength, toughness, and crack propagation restriction potentials because of its unique microstructure, consisting of irregular needle-like interlocking grains enveloped by a

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61

thin (5 mm) and relatively less purity. Early failure due to high fracture rates and unanticipated roughening was reported because of suboptimal fixation of the ace tabular component. Second generation of bioceramics was made by addition of calcium and magnesium oxide to alumina and zirconia to decrease grain size. Third generation of bioceramics underwent hipped processing and hot isostatic pressing in addition to grain size optimization, improved purity, and application of presterilization loading to guarantee satisfactory material properties, e.g., BIOLOX forte femoral heads (CeramTec Group). The latest, fourth generation of bioceramics, called alumina matrix composite material, was introduced for clinical usage in early 2000’s and demonstrated better wear resistance, higher grain uniformity, and smaller grain size than former generations. A fraction of tetragonal zirconia crystals and strontium oxides was added to fourth generation ceramics to improve toughness. A small amount of chromium oxide was added to increase hardness, which was compromised because of zirconium addition (Gamble et al., 2017). Total hip arthroplasty alumina bioceramics are chiefly applied in total hip arthroplasty as cups in alumina-on-alumina amalgamation and as femoral heads articulating against polyethylene. Initially, expanded usage of first- and second-generation bioceramics was limited because of greater possibility of fracture due to their brittle nature and high cost. In third- and fourth-generation of bioceramics, fracture risk has diminished greatly because of advancement in fabrication technique with amplified purity and density, increase in grain size, and dispersal in addition to improved quality control (Prakash, 2018). Since 1984, alumina is a standardized material (ISO 6474) for arthroplasty. Frictional characteristics of alumina against itself are excellent and exhibits 4000 times lesser linear wear rate than that of metal on polyethylene. At the same time, hydrophilic alumina surfaces and fluid film lubrication between them diminish adhesive wear of

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alumina. In alumina-on-alumina combination, clearance between components should be about 50 nm to avoid development of Hertz stresses at alumina surface to avoid grain detachment and third-body wear. Both in vitro studies and analysis of retrieved implants showed that formation of limited amount of wear debris leads to moderate biological response (Prakash, 2018). Yettrium-stabilized zirconia has been widely considered for medical use because of its better mechanical properties, despite its difficult sintering. In 1969, it was proposed to utilize zirconia for hip head replacement, and no adverse mechanical behavior, wear, and bone and muscle integration were observed. In 1990, solid samples of zirconia were marked not cytotoxic (Dion et al., 1994; Torricelli et al., 2001; Lohmann et al., 2002) and nonmutagenic (Silva et al., 2002; Covacci et al., 1999). Yettriumstabilized zirconia has been employed successfully in ball head assembly used for total hip replacement and is still used mainly for this purpose (Manicone et al., 2007). A well-designed Morse taper permitting exact fixation of bioceramic ball to femoral stem evading unwanted strains in the head and enhanced surgical practices augment the life span of implant (Prakash, 2018). Currently, most of the ceramic femoral heads exhibit a 12 mm  14 mm Morse Taper, which makes them well-suited with most of the metal femoral stems. The evaluated fracture rate of bioceramic femoral heads is 0.02% for alumina and 0.03% for zirconia, demonstrating that bioceramic head fracture is no longer an issue. Cost, however, is still a limiting factor as bioceramic heads are 6 to 10 times more costly than metal heads (Prakash, 2018; Gamble et al., 2017).

3.3.1.2

Orthopedic fixation devices

Arthrodesis device Silicon nitride is a tough and high-strength commercial ceramic with potential to be considered as implant material. In 1986, in Australia, a clinical trial involving 30 patients was carried out to assess its fitness as an arthrodesis device in lumbar spine (Mcentire et al., 2015). Since 2008, it has been utilized as a successful fusion cage implantation for arthrodesis of cervical and thoracolumbar spine. Silicon nitride holds potential to be used as articulation member in future (Mcentire et al., 2015).

Spinal fusion In lumber spine fusion, bioceramic-based bone grafts are promising bone graft extender in the presence of an osteoinductive stimulus, i.e., local bone graft. For lumbar spine fusion, bioceramics offer numerous advantages over other bone graft extenders such as shape flexibility, inertness, safety profile, and ease of sterilization. However, a necessity for an osteoinductive adjunct, low tensile strength, and susceptibility to fracture limit its use (Nickoli and Hsu, 2014). It is reported that highest spinal fusion rate with apatiteewollastonite-enriched glass ceramic demonstrated by patients was highest, i.e., 100% followed by dense HA block (96.2%), b-TCP (92.5%) (Vitoss; Stryker, Kalamazoo, MI), and coralline HA (86.9%) (Pro-Osteon 200, Pro-Osteon 500; Biomet, Warsaw, IN) and least by synthetic HA (86.7%) and calcium sulfate (86.7%) (Osteoset; Wright Medical Technology, Memphis, TN) (Nickoli and Hsu, 2014).

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3.3.2

Handbook of Ionic Substituted Hydroxyapatites

Coatings for chemical bonding

Currently, metals such as stainless steel, cobalt chromium alloy, or titanium alloy are employed as implants. Unfortunately, most of these implants lack bioactivity; therefore, various bioactive implant surface coatings are carried out to achieve osseointegration (Zhang et al., 2014). These implant surface coating materials should be biocompatible, osteoconductive, and osteoinductive and should exhibit sufficient mechanical stability under physiological loading without detaching from implant surface and antimicrobial properties to minimize prosthetic infection risk (Zhang et al., 2014; Tobin, 2017). Among bioceramics, implant coating materials include calcium phosphates, HA, diamond-like carbon, zirconium nitride, titanium nitride, and titanium niobium nitride (Mcentire et al., 2015, Mazare et al., 2018; Brunello et al., 2018; Walker et al., 2019). Bioceramic coatings are beneficial in articulating prostheses as they eradicate brittle fracture due to their tough metallic substrates, lower production of wear fragments due to greater hardness and wear resistance of ceramic coat, and minimize allergic reactions due to soluble metal ions such as cobalt, chromium, and nickel by acting as a barrier between metallic implant and human bone (Zhang et al., 2014). Electrosprayed surface coatings influence protein interactions, offering potential for topographical and chemical control of cell behavior and drug delivery, however, at the expense of causing stress shielding to the adjacent bone leading to the deterioration of bone cells making new bone (Siebers et al., 2004, 2006a,b; Layrolle et al., 2001). In vivo pitting, scratching, or even delamination of thin ceramic coatings have been detected, leading to prosthesis failure (Mcentire et al., 2015). Prevalence of diseasefree loosening of joint prosthesis after 10 years of surgery is nearly 2% for hip and knee replacements (Zhang et al., 2014).

3.3.2.1

Hydroxyapatite surface coatings

HA is a potential bioactive ceramic surface coating material for orthopedic implants, which is a calcium phosphateebased osteoconductive bioceramic that has shown to promote adhesion and differentiation of osteoblasts to yield natural apatite deposition on implant surface through ion exchange process between HA coating and the surrounding physiological fluids in in vitro models and clinical trials in humans (Mello et al., 2007; Smith et al., 2006; Voigt and Mosier, 2011; Rajaratnam et al., 2008; Lazarinis et al., 2012). Fluoride-doped HA coatings on metallic implants lead to in vitro stimulation of cell proliferation and differentiation; however, it takes longer time to form new bone coatings and adhesive strength between substrate and coatings (Wang et al., 2009; De Groot et al., 1987). Failure of HA-coated implants revealed delamination and resorption of HA coating due to implant-coating attachment failure (Mohseni et al., 2014, LE Guéhennec et al., 2007). New implant coating techniques with HA include thermal spraying, plasma spraying, pulsed laser deposition, dip coating, sputter coating, electrophoretic deposition, hot isostatic pressing, solegel, and ion-beam assisted deposition (Mohseni et al., 2014).

Bioceramics: types and clinical applications

3.3.2.2

65

Diamond-like carbon coatings

Tribological and mechanical properties related to diamond-like carbon films in orthopedic devices have been extensively explored. Diamond-like carbon coated metal-to-polyethylene gliding hip simulator joints exhibited decrease in wear rate on comparison with metal-to-polyethylene and metal-to-metal joints (Tiainen, 2001). An increase in life span of the implant was detected when both surfaces of gliding pair, i.e., ultra-high molecular weight polyethylene and cobalt-chromium-molybdenum, were coated with diamond-like carbon (Sheeja et al., 2005). Diamond-like carbon surface coatings in aqueous environment are unstable and delaminate because of poor adhesion to substrate and high residual stress. Much higher failure rate of diamond-like carbonecoated Tie6Ale4V femoral head was reported than alumina femoral head (Taeger et al., 2003). Radiofrequency plasmaedeposited diamond-like carbon coating on stainless steel medical guidewires exhibit adherence with a coefficient of friction better than that of polytetrafluoroethylene without altering the stiffness of guidewires (Mclaughlin et al., 1996). A study showed that the diamond-like carbon coating on suturing needle reduced the force needed to penetrate cornea of pigs’ eyes by 30%. In microsurgery, the dark color of diamond-like carbon would decrease reflections from illuminations of operation microscopes (Dearnaley and Arps, 2005). Thin diamond-like carbon coatings of about 20e200 nm to soft contact lenses and contact lens cases not only reduce problems of biofilm development but also showed that contact lenses kept in diamond-like carbon coated cases remained uncontaminated (Sleptsov et al., 1996; Dearnaley and Arps, 2005). In dentistry, diamond-like carbon coatings also have the potential to improve osseointegration of dental implants (Roy and Lee, 2007).

3.3.2.3

Calcium phosphates coatings

In calcium phosphate family, only HA and TCP have been extensively evaluated as orthopedic implants. HA coating on femoral prostheses and on sockets is quiet popular fixation means to prevent complications related to PMMA usage. A study in an American multicentre has reported only 0.3% rate of femoral revision for mean follow-up of 8.1 years in 324 implants. Still, it is unclear that HA leads to improved fixation of prosthesis than bone cement (Prakash, 2018). HA chemical composition, coating thickness, surface roughness, and metal substrate’s nature seem to be key factors in achieving required fixation. Poor tensile strength and brittle nature of HA have limited its clinical application as bone-graft substitute. Scarce information is available on clinical usage of ceramic bone-graft substitutes. TCP application in spinal fusion exhibited results similar to those with autogenous bone (Prakash, 2018).

3.3.2.4

Nitride coatings

In late 1980’s for complete joint arthroplasty, physical vapor depositionecoated titanium implants with titanium nitride coatings were introduced, which was clinically investigated after a decade. The nanocrystalline nitride coatings were applied in 1e15 mm thickness

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and are currently known as hip resurfacing implants, for total hip arthroplasty, or as femoral knee components (Subramanian et al., 2011, Mcentire et al., 2015).

3.3.3

Bone tissue engineering

Bone defects display compete recovery with suitable treatment due to good bone regenerative potentials. Still, bone autografts harvested from ilia or tibiae are required in surgeries such as bone tumors or spinal fixation, which may sometimes manifest into pain at the donor site. Because of limited quantity of autografts, a substantial research has been performed on bone tissue engineering using mesenchymal stem cells and bone morphogenetic proteins to replace autografting (Sotome et al., 2004). Development of scaffolds to carry growth factors and mesenchymal stem cells is a major breakthrough in field of bone regeneration (Sotome et al., 2004). These bone scaffolding materials include TCP s, HA, and bioglass alone or in composite form. Synthetic bone graft materials are presented to surgeons in putty form, particulate form, and porous 3D scaffolds (Gul et al., 2015). In dentistry, these are used as temporary bone space fillers, for alveolar ridge augmentation, for periodontal pocket reduction, and for maxillofacial reconstruction.

3.3.3.1

Bioceramic bone fillers

Kikuchi et al. in 2001 developed hydroxyapatite/collagen (HAp/Col) nanocompositee based artificial bone blocks in which c-axis of HA crystals of up to 50 nm length was aligned along collagen fibers. These HAp/Col-based biodegradable and osteoconductive blocks had a potential to be used as a carrier of recombinant human bone morphogenetic protein 2. Limited tissue invasion at implant site without implant disintegration due to dense implant structure limits its applications (Kikuchi et al., 2001). In 2004, Sotome et al. developed HA/collagenealginate bone filler with rh-BMP2 (100 mg/mL, 15 mL), which showed osteogenesis round the implant after 5 weeks of implantation with no obvious material deformation, whereas osteogenesis was seen only in squashed collagen sponge portion (Sotome et al., 2004). Bioglass products in orthopedics and dentistry are surgically placed into bony defect site, whereon gradual resorption, new bone ingrowth occurs (Gul et al., 2015). Dentists and surgeons prefer bioglass in particulate form instead of monoliths, as they offer ease to fill a bony defect. In 1993, first particulate bioglass was introduced commercially by the name PerioGlas (NovaBone Products LLC, Alachua, Florida) for alveolar ridge augmentation. PerioGlas is also employed in conjunction with polymeric membranes in “guided tissue regeneration.” Other commercial bioglass products, which have been used in dentistry and orthopedic, as bone graft are BonAlive (BonAlive Biomaterials, Turku, Finland) and Biogran (BIOMET 3i, Palm Beach Gardens, Florida) (Abbasi et al., 2015).

Bioceramics: types and clinical applications

3.3.3.2

67

Tissue-engineered bone constructs

Because of scarcity of appropriate allograft and autograft materials, tissue-engineered bone constructs are gaining popularity to meet the increasing demands due to their potential to augment healing of bone. Scaffolds used in bone tissue engineering are typically composed of porous materials that deliver mechanical support throughout the repair and regeneration of diseased or damaged bone accompanied by degradation with time. In the field of bone tissue engineering, researches are constantly bringing on innovations since one decade to develop new materials via new processing methods, followed by evaluation for mechanical support with anticipated osseogenesis and angiogenesis capacities and their possible applications (Shadjou and Hasanzadeh, 2016). Human osteoblasts and mesenchymal stromal cells when cultured on graphene films exhibited better cell adhesion and proliferation (Fan et al., 2010; Zhang et al., 2014). A 3D porous graphene/nano-bioglass58S composite scaffold for bone tissue engineering with only 0.5wt.% graphene exhibited improved fracture toughness and compressive strength with promising bioactivity and biocompatibility, suggesting its potential application for bone tissue engineering (Gao et al., 2014). Bioglasse graphene nanoplatelets formed by spark plasma sintering technique exhibited comparatively higher viscosity, sintering temperature, and electrical conductivity than pure bioglass without compromising the bioactivity, thus making possible the fabrication of electrically conductive and bioactive scaffolds for bone tissue engineering (Porwal et al., 2014).

3.3.4 3.3.4.1

Dental applications Dental prosthesis

Ceramics, such as alumina, spinel, lithium disilicateereinforced ceramics, and zirconia can be utilized in the creation of metal-free fixed dental prosthesis (Raigrodski, 2004). Among these, highest average load-bearing capacities were recorded for restorations manufactured with zirconia followed by alumina and least by lithium disilicate. Reinforced ceramics only prove beneficial for making of an anterior crown restoration or a three-unit bridge that exhibits resistance to chewing stresses only on anterior teeth (L€ uthy et al., 2005). Alumina cores are attractive candidates for cores in anterior porcelaineveneer € crowns (Ozcan and Vallittu, 2003). Some of the commercial alumina-reinforced core porcelain systems that can be used in 0.5 mm thickness includes In-Ceram Spinell core, In-Ceram Alumina core, In-Ceram Zirconia core, and Procera AllCeram core (Heffernan et al., 2002). Zirconia restorations are indicated for teeth or implants supported by fixed dental prosthesis (Potiket et al., 2004; Piwowarczyk et al., 2005) (Larsson et al., 2006).

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Among all ceramic-fixed dental prosthesis, zirconia with alumina oxide displayed the maximum initial and long-term strength (Tinschert et al., 2007). Zirconia also offers better esthetics due to its opacity and color matching to natural teeth (Mclaren and Giordano, 2005); still if translucency is needed, then alumina or lithium disilicate surface coating can be carried out. Computer-aided design/manufacturing (CAD/ CAM) system is utilized to create a fitting framework of zirconia prosthesis (Manicone et al., 2007).

3.3.4.2

Laminate veneers

Since early 1980’s, porcelain veneers are a solution to unaesthetic anterior teeth. Clinical studies have indicated favorable clinical performance of porcelain veneers due to their excellent esthetics, high patient satisfaction, and no drastic effects on gingival health (Ho and Matinlinna, 2011). Continual advancement of dental ceramics offers dentists numerous choices for producing extremely esthetic and functional ceramic veneers. Types of ceramic materials with sufficient translucency and potential to be used in small thickness, indicated for dental veneer fabrication, include pressable ceramic and sintered feldspathic porcelain. Ceramics with noncrystalline or glassier microstructure are more translucent than the ceramics with crystalline microstructure (Pini et al., 2012). Traditional dental ceramics are feldspar porcelain, composed of feldspar, quartz, and kaolin, and fired above 870 C to form dental veneers. Feldspar-based ceramics because of their brittle nature fracture easily; therefore, more crystalline ceramics with improved mechanical properties, such as alumina and zirconia, were developed. Ceramics with high crystalline content are generally employed as core materials, while feldspar-based ceramics are utilized as veneers in prosthetic dentistry (Ho and Matinlinna, 2011).

3.3.4.3

Dental implants

Zirconia implants were introduced into dental implantology as an alternative to titanium implants. They are bioinert and exhibit minimal ion leaching than metallic € implants (Ozkurt and Kazazo glu, 2011). Zirconia seems to be a suitable implant material because of its tooth-like color, high strength, high fracture toughness, biocompatibility, and low plaque affinity (Depprich et al., 2008; Piconi and Maccauro, 1999). Moreover, zirconium oxide creates less inflammation and bone resorption € than titanium (Depprich et al., 2008; Ozkurt and Kazazoglu, 2011). Clinical usage of zirconia as dental implants is limited due to difficulty in carrying out surface modifications in zirconia implants, which leads to insufficient osseointegration due to smooth implant surfaces. Round zirconia particles can be sandblasted on titanium € implants surfaces to augment osseointegration of titanium implants (Ozkurt and Kazazo glu, 2011). At present, nine zirconia dental implant systems are available commercially. The first one was introduced in 1987 by name of “the Sigma implant” (Sandhausen, Intermed, Lausanne, Switzerland). Others include “CeraRoot system” (Oral Iceberg,

Bioceramics: types and clinical applications

69

Barcelona, Spain), “Goei system” (Goei Inc, Akitsu-Hiroshima, Japan), “White Sky system” (Bredent Medical, Senden, Germany), “ReImplant system” (ReImplant, Hagen, Germany), “Konus system” (Konus Dental, Bingen, Germany), and “Ziterion system” (Ziterion, Uffenheim, Germany), and “Z-systems” (Z-systems, Konstanz, € Germany) (Ozkurt and Kazazo glu, 2011).

3.3.4.4

Restorative dentistry

Nanosized apatite crystals incorporation into fluoroaluminosilicate glass powder of conventional GICs not only improve the mechanical properties of GICs but also lead to enhanced fluoride elusion and bioactivity on setting. HA addition to GICs also augmented the chemical stability, insolubility, and bond strength with tooth (Najeeb et al., 2016; Moshaverinia et al., 2008a, 2008b). As retrograde filling materials, bioceramics exhibited significantly higher clinical success rates, i.e., 86.4%e95.6%, in apical surgery (Abusrewil et al., 2018).

3.3.4.5

Preventive dentistry

Tooth enamel is the hardest tissue of human body, which is highly mineralized and constitutes about 97% HA. Being a static, nonregenerating tissue after eruption, modern oral care focuses mainly on protection and preservation of tooth enamel. In the oral cavity, enamel is exposed to acids mainly under two conditions; either enamel can be eroded by acidic food or beverages (acid erosion) or the enamel crystallites can be partially demineralized within the caries process in dental plaque. For enamel remineralization, calcium phosphates have been identified as promising biomimetic alternatives because of their similarity to natural enamel (Meyer et al., 2018). Calcium phosphates products used frequently for oral care includes b-TCP, amorphous calcium phosphate, and HA (Meyer et al., 2018; Khan and Syed, 2019). HA-enriched toothpaste use in patients leads to not only enamel remineralization with improved acid resistance by forming a protecting layer on the enamel surface but also improve their periodontal health (Lelli et al., 2014; Makeeva et al., 2016; Harks et al., 2016). However, HA containing mouthwash reported to reduce not only plaque and gingival index but also caused remineralization of white spot lesions on enamel (Hegazy and Salama, 2016). In situ studies reported that HA particles reduce initial biofilm formation on enamel surfaces and causes remineralization of early carious lesions (Kensche et al., 2017; Hannig et al., 2013). HA-enriched mouthwash use reported to reduce plaque and gingival index (Hegazy and Salama, 2016). Various in situ studies have shown that amorphous calcium phosphateeenriched toothpastes are not as effective in reversal of white spot lesions and enamel erosion as is the continued use of fluoride-based dentifrices (Meyer-Lueckel et al., 2015; Wiegand and Attin, 2014; Kensche et al., 2016). Study by Hegde et al. showed that amorphous calcium phosphate when used in chewing gums increased the buffering capacity of saliva (Hegde and Thakkar, 2017). Regression of incipient caries was observed by use of amorphous calcium phosphateeenriched cream as postorthodontic care (Bailey et al., 2009).

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Calcium silicates hold a potential to be used for treating dentin hypersensitivity and enamel white spot lesions by its remineralization effect (Li et al., 2014). In vitro studies have shown that calcium silicate cement application on dentin surface immediately lower the permeability of dentin (Gandolfi et al., 2008, 2012), by penetration into dentinal tubules and forming a dense thick apatite layer on dentin surface (Dong et al., 2011). It is reported that tricalcium silicate treatment of demineralized enamel exhibited similar remineralization of enamel as produced by 1000 ppm fluoride treatment. However, the combination of both produced much greater reduction in enamel demineralization than other treatments (Wang et al., 2012). Bioglass occludes dentinal tubules by deposition onto dentin surfaces and subsequently inducing formation of carbonated HA. Limited studies are reported on enamel remineralization potential of bioglass. Burwell et al. reported that 10 days enamel treatment with blend of bioglass with 5000 ppm fluoride gave pointedly superior remineralization than 5000 ppm fluoride alone (Li et al., 2014).

3.3.4.6

Endodontics

In dentistry, physical properties of bioceramics including absolute biocompatibility, radiopacity, chemical bonding to tooth structure, osseoconductivity, insolubility in tissue fluid, effective hermetic seal ability, and easy handling properties have led to extensive usage of these materials in endodontics (Utneja et al., 2015). Calcium phosphate is a biocompatible material used for pulp capping, apexification, inducing formation of hard tissue, and apical barrier. Sealers based on calcium phosphate are less cytotoxic than zinc oxide eugenol and epoxy-based AH26 sealers and have osteogenic potentials (Utneja et al., 2015). Clinical trials and animal usage tests have revealed effective applications of HA in endodontic procedures including pulp capping, apical barrier development, repair of furcation perforation, periapical defects, and regenerative endodontics (Utneja et al., 2015). Bioglass is a bioactive material, exhibiting better reparative dentin formation potential without any tissue necrosis than calcium hydroxide (Oguntebi et al., 1993). Bioglass-enriched sealers induce superior root development via apexification than TCP (Zeid et al.). Torabinejad and White developed MTA for use as retrograde filling and endodontic repair (Lee et al., 1993). MTA was patented and permitted for endodontic usage in 1995 (Parirokh and Torabinejad, 2010). The bioactivity, biocompatibility, and effective endodontic seal ability of MTA (Kim and Shin, 2014) has gained its popularity in surgical and nonsurgical endodontic usage (Dammaschke et al., 2005). In deciduous dentition, MTA is indicated for pulpotomy, pulp capping, root canal filling, resorption repair, and furcation perforation repair. In permanent dentition, MTA is indicated for partial pulpotomy, pulp capping, root canal sealer, apical/lateral/furcation perforation repair, internal/external resorption repair, repair of horizontal/vertical fracture, retrograde root end filling, apical barrier for developing tooth with open apex and necrotic pulp, and coronal barrier in regenerative endodontics (Macwan and Deshpande, 2014). Most commonly found commercial calcium silicateebased root-end filling cements include Gray ProRoot MTA (Dentsply Sirona, USA), White ProRoot MTA

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(Dentsply Sirona, USA), calcium-enriched mixture (CEM) cement, Biodentine (Septodont, USA), BioAggregate (Innovative BioCeramix Inc., Canada), EndoSequence Root Repair Material (Brasseller, USA), and iRoot BP Plus Root Repair Material (Innovative BioCeramix Inc., Canada) (Abusrewil et al., 2018).

Calcium-enriched mixture cement In 2006, CEM cement was introduced as a new endodontic filling material (Utneja et al., 2015). Film thickness, flow, and initial setting time of this cement are satisfactory (Asgary et al., 2008c). It also exhibits shorter setting time in aqueous environments than MTA, whereas sealing ability is similar to MTA (Asgary et al., 2008a,c). CEM cement’s antibacterial effect is analogous to calcium hydroxide and superior than MTA (Asgary and Kamrani, 2008). CEM cement promotes formation of HA in saline and may hold the potential of inducing hard tissue development (Asgary et al., 2009, 2010; Nosrat et al., 2011). Clinical applications of CEM cement are the same as MTA. CEM exhibits comparable outcomes to MTA when applied for repair of furcation perforation or as pulp capping agent (Asgary et al., 2008b; Samiee et al., 2010). It also exhibits promising results in pulpotomy of permanent molars with irreversible pulpitis and in treatment of internal root resorption (Asgary and Ehsani, 2009; Asgary and Ahmadyar, 2012).

3.3.5 3.3.5.1

Ocular prosthesis Glass ocular implants

Glass is a noncrystalline, oxide-based ceramic and has long history as a biomedical material (Baino and Potestio, 2016). Mules placed the first nonintegrated orbital glass implant in 1985 after evisceration (Anderson et al., 1990; Jordan et al., 2000). Mules implant, being a brittle hollow glass sphere, was prone to fracture on impact and on temperature fluctuations exhibited risk of implosion. Glass orbital implants are no more in use as new better orbital implant materials replaced them (Catalu et al., 2018).

3.3.5.2

Silicon ocular implants

Silicon is a biologically/chemically inert material, exhibiting good flexibility, easy manipulation, and low cost. Solid or porous silicon episcleral implants are the only clinically approved devices available commercially for scleral buckling in surgery of retinal detachment (Baino, 2010). At the end of 1980’s, nonporous silicon orbital implants were commercialized. These were used as such or wrapped within the center of muscle cone and bound to four rectus muscles. These are still now a good choice in absence of pegging. They allow little prosthetic movements than mounded or pegged devices, and the orbital implant has a tendency to migrate over time (Jordan and Klapper, 2010). Nonporous silicon spheres are preferred orbital implants in newborns/preschoolaged kids and in case of severe gunshot orbital injury. Silicon orbital implants exhibited much less pre/postpegging complications such as hypoophthalmos and

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pyogenic granuloma and are commercially available in the United States, which include “flexiglass system” and “flexiglass eye” (Baino et al., 2014).

3.3.5.3

Porous hydroxyapatite ocular implants

In mid-1980s, Perry introduced porous coralline HA sphere, which was implemented clinically in early 1990s, and it became most popular ocular implant after primary enucleation and was commercially available in the name of “Bio-eye sphere” (Arthur, 1991; Hornblass et al., 1995). However, coralline HA was costly; therefore, synthetic HA gained popularity as orbital implants as they were not only cheap but also allowed easy drilling for pegging than Bio-eye (Baino and Potestio, 2016). As compared with silicon ocular implants, interconnected porous construction of HA ocular implant facilitates fibrovascular ingrowth of host tissues, potentially minimizing risk of extrusion, migration, and infection (Nunery et al., 1993; Gradinaru et al., 2015). HA ocular implants are positioned within a covering material before introducing into orbit (Custer, 2000; Li et al., 2001; Babar et al., 2009). Exposed HA ocular implants allow successful treatment via patch grafts of sclera, dermis, and oral mucosa, without the necessity of implant elimination (Owji et al., 2012; Jeong et al., 1996; Rernulla et al., 1995). A composited eye implant developed by Guthoff et al. namely “quasi-integrated implants” constituted two parts, i.e., frontal component composed of synthetic porous HA and posterior component composed of silicon rubber (Jordan and Bawazeer, 2001). It allowed suturing of eye muscles crosswise anterior to the implant to promise better stability and motility (Choi et al., 2006).

3.3.5.4

Aluminum oxide ocular implant

High-purity aluminum oxide is a promising alternative to coralline and synthetic HA (Baino and Potestio, 2016). Because of inert nature, biocompatibility, and good mechanical properties of aluminum oxide, it has been used for decades in orthopedics (Chee et al., 2013). The porous form of aluminum oxide was accepted in 2000 by US Food and Drug Administration and is commercially available in name of “Bioceramic implant” (Catalu et al., 2018).

3.3.6

Otolaryngologic applications

Bioceramics used in production of passive middle ear implants include glass, HA, bioglass, aluminum oxide, Bioverit, Macor, and Ceravital (Beutner and H€uttenbrink, 2009). In the year 2000, HA was reported to be the most popular bioceramic in the United States for ossicle chain replacement (Goldenberg and Emmet, 2001). However, in Europe, titanium was more popular than ceramics, where titanium was used in 68% of cases and HA in only 24% cases (Preuss et al., 2007). Plastipore are commercially available light weight titanium implants, which offer easier handling to many otologists than relatively large and ungainly ceramic middle ear prostheses. However, clinical trials have shown that acoustic edge for titanium was not significant

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(Truy et al., 2007) as numerous other factors impact postoperative hearing. For implant material selection, individual preferences and the allied experience of discrete surgeons also play a significant role, thus hybrid prostheses such as Flex H/A and HAPEX are still used justifiably (Beutner and H€ uttenbrink, 2009).

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Chevalier, J., Gremillard, L., 2009. Ceramics for medical applications: a picture for the next 20 years. J. Eur. Ceram. Soc. 29, 1245e1255. Choi, H.Y., Lee, J.-E., Park, H.J., Oum, B.S., 2006. Effect of synthetic bone glass particulate on the fibrovascularization of porous polyethylene orbital implants. Ophthalmic Plast. Reconstr. Surg. 22, 121e125. Covacci, V., Bruzzese, N., Maccauro, G., Andreassi, C., Ricci, G., Piconi, C., Marmo, E., Burger, W., Cittadini, A., 1999. In vitro evaluation of the mutagenic and carcinogenic power of high purity zirconia ceramic. Biomaterials 20, 371e376. Custer, P.L., 2000. Enucleation: past, present, and future. Ophthalmic Plast. Reconstr. Surg. 16, 316e321. Gamble, D., Jaiswal, P.K., Lutz, I., Johnston, K.D., 2017. The use of ceramics in total hip arthroplasty. Orthop. Rheumatol. 4, 7. Dammaschke, T., Gerth, H.U., Z€uchner, H., Sch€afer, E., 2005. Chemical and physical surface and bulk material characterization of white ProRoot MTA and two Portland cements. Dent. Mater. 21, 731e738. Darvell, B., Wu, R., 2011. “MTA”dan hydraulic silicate cement: review update and setting reaction. Dent. Mater. 27, 407e422. De Aza, P., De Aza, A., De Aza, S., 2005. Crystalline bioceramic materials. Bol. Soc. Esp. Ceram 44, 135e145. De Groot, K., Geesink, R., Klein, C., Serekian, P., 1987. Plasma sprayed coatings of hydroxylapatite. J. Biomed. Mater. Res. 21, 1375e1381. De Lima, I.R., Alves, G.G., Soriano, C.A., Campaneli, A.P., Gasparoto, T.H., Schnaider  Rossi, A.M., Granjeiro, J.M., 2011. Understanding the impact of Ramos, E., De Sena, L.A., divalent cation substitution on hydroxyapatite: an in vitro multiparametric study on biocompatibility. J. Biomed. Mater. Res. A 98, 351e358. Dearnaley, G., Arps, J.H., 2005. Biomedical applications of diamond-like carbon (DLC) coatings: a review. Surf. Coat. Technol. 200, 2518e2524. Depprich, R., Zipprich, H., Ommerborn, M., Naujoks, C., Wiesmann, H.-P., Kiattavorncharoen, S., Lauer, H.-C., Meyer, U., K€ubler, N.R., Handschel, J., 2008. Osseointegration of zirconia implants compared with titanium: an in vivo study. Head Face Med. 4, 30. Dion, I., Bordenave, L., Lefebvre, F., Bareille, R., Baquey, C., Monties, J.-R., Havlik, P., 1994. Physico-chemistry and cytotoxicity of ceramics. J. Mater. Sci. Mater. Med. 5, 18e24. Dong, Z., Chang, J., Deng, Y., Joiner, A., 2011. Tricalcium silicate induced mineralization for occlusion of dentinal tubules. Aust. Dent. J. 56, 175e180. Dorozhkin, S., 2018. Current state of bioceramics. J. Ceram. Sci. Technol. 9, 353e370. Dubruille, J.-H., Viguier, E., Le Naour, G., Dubruille, M.-T., Auriol, M., Le Charpentier, Y., 1999. Evaluation of combinations of titanium, zirconia, and alumina implants with 2 bone fillers in the dog. Int. J. Oral Maxillofac. Implant. 14, 271e277. EL-Ghany, O.S.A., Sherief, A.H., 2016. Zirconia based ceramics, some clinical and biological aspects. Future Dent. J. 2, 55e64. Eliaz, N., Metoki, N., 2017. Calcium phosphate bioceramics: a review of their history, structure, properties, coating technologies and biomedical applications. Materials 10, 334. Eslami, H., Tahriri, M., Moztarzadeh, F., Bader, R., Tayebi, L., 2018. Nanostructured hydroxyapatite for biomedical applications: from powder to bioceramic. J. Korean Chem. Soc. 55, 597e607. Fan, H., Wang, L., Zhao, K., Li, N., Shi, Z., Ge, Z., Jin, Z., 2010. Fabrication, mechanical properties, and biocompatibility of graphene-reinforced chitosan composites. Biomacromolecules 11, 2345e2351.

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Voigt, J.D., Mosier, M., 2011. Hydroxyapatite (HA) coating appears to be of benefit for implant durability of tibial components in primary total knee arthroplasty: a systematic review of the literature and meta-analysis of 14 trials and 926 evaluable total knee arthroplasties. Acta Orthop. 82, 448e459. Vorndran, E., Klarner, M., Klammert, U., Grover, L.M., Patel, S., Barralet, J.E., Gbureck, U., 2008. 3D powder printing of betricalcium phosphate ceramics using different strategies. Adv. Eng. Mater. 10, B67eB71. Walker, T., Rutkowski, L., Innmann, M., Panzram, B., Herre, J., Gotterbarm, T., Aldinger, P., Merle, C., 2019. Unicondylar knee arthroplasty using cobalt-chromium implants in patients with self-reported cutaneous metal hypersensitivity. Bone Joint J. 101, 227e232. Wang, J., De Groot, K., Van Blitterswijk, C., De Boer, J., 2009. Electrolytic deposition of lithium into calcium phosphate coatings. Dent. Mater. 25, 353e359. Wang, Y., Li, X., Chang, J., Wu, C., Deng, Y., 2012. Effect of tricalcium silicate (Ca3SiO5) bioactive material on reducing enamel demineralization: an in vitro pH-cycling study. J. Dent. 40, 1119e1126. Wiegand, A., Attin, T., 2014. Randomised in situ trial on the effect of milk and CPP-ACP on dental erosion. J. Dent. 42, 1210e1215. Zeid, S.T.A., Ezzat, K.M., Kamal, S.M., Reparative Activity of Bioglass and Tricalcium Phosphate to Induce Apical Closure Compared to Normal Apexogensis (Histological and Histochemical Study). Zhang, B., Myers, D., Wallace, G., Brandt, M., Choong, P., 2014. Bioactive coatings for orthopaedic implantsdrecent trends in development of implant coatings. Int. J. Mol. Sci. 15, 11878e11921. Zheng, Y., Yang, Y., Deng, Y., 2019. Dual therapeutic cobalt-incorporated bioceramics accelerate bone tissue regeneration. Mater. Sci. Eng. C 99, 770e782.

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Basics of hydroxyapatitedstructure, synthesis, properties, and clinical applications

4

Hamad Khalid, Aqif Anwar Chaudhry Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

4.1

Biological apatite and synthetic hydroxyapatite: differences and similarities

Hydroxyapatite (HA) is well thought out as a biomaterial and hence widely employed in many health applications. As a source of calcium, HA is particularly used as a vital compound to repair bones and also in toothpastes. HA exhibits a great bioactivity and is highly compatible with the adjacent bone and teeth in living organisms because of its appropriate chemical properties. HA is a ceramic material composed of calcium phosphate owning a high biocompatibility and nontoxicity, and it is an integral part of bone and teeth tissues (Sossa et al., 2018; Raya et al., 2015). When an implant coated with HA meets the osseous tissue, bone growth is promoted toward the implant (Zhou and Lee, 2011; Milovac et al., 2014). The story of HA started in 1926 by Jong et al. when the chemical formula of HA was found to be Ca10(PO4)6(OH)2 and a relation between HA and osseous mineral was observed. Synthetic HA is similar to biological apatite (or osseous mineral), but there are differences in chemical composition and crystallinity. Since then, HA has been investigated widely (Sossa et al., 2018). It is important to understand the structure of the naturally occurring product if it is to be synthesized in a lab or on commercial scale. So, the research on HA carried on and a great advancement was the identification of crystallographic structure of HA by Ponser et al. (Posner, 1969; Posner et al., 1958; Kay et al., 1964). They used chemically precipitated HA for structure illumination. Nowadays, intensive research focusses on development of high-quality HA with improved biocompatibility and via ecofriendly procedures (Huang et al., 2011; Sadat-Shojai et al., 2013). Different methods are adopted to produce HA, e.g., hydrothermal (Zhang et al., 2011; Joki
c et al., 2011), chemical precipitation (Stipniece et al., 2014; Kim et al., 2010; Dhand et al., 2014), and use of biogenic sources (Huang et al., 2011; Kamalanathan et al., 2014). The chemical precipitation and synthesis from biogenic sources are cost effective, have fewer byproducts, are easily available, and have controllable particle size (Sadat-Shojai et al., 2013; Kamalanathan et al., 2014). Especially, the biogenic source

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00004-5 Copyright © 2020 Elsevier Ltd. All rights reserved.

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can produce HA with high crystallinity, does not have contaminations, and is friendly with the environment. Moreover, the HA produced by biogenic source can maintain the natural architecture of the bone tissue, which could behave as osteoconductive material (Mucalo, 2015). There are many biogenic sources of HA, among them egg shells, animal bones, and fish scales are important (Kamalanathan et al., 2014; Mondalet al., 2012; Haberko et al., 2006; Janus et al., 2008; Sobczak-Kupiec and Wzorek, 2012). HA produced form chemical precursors has many advantages such as controllable Ca/P ratio, high crystallinity, and good purity. But there are some disadvantages of chemical precursors methods such as lack of Fe2þ, Mg2þ, Si2þ, Naþ and F ions, which are naturally present in bone tissues, so the biological activity could be altered (Milovac et al., 2014; Akram et al., 2014). Rahavi et al. synthesized and extracted HA using solegel method and calcination of natural bones at 700 C, respectively (Rahavi et al., 2017). They used bones from human and different animals to extract HA. They concluded that X-ray diffraction (XRD) signatures of both types of HAs were in agreement with stoichiometric ratio. Synthetic HA, however, showed small amounts of beta-tricalcium phosphate (b-TCP) and CaO, which showed that conversion of raw materials into products was not 100 %. Moreover, the natural HA showed higher degree of crystallinity. Scanning electron microscopy (SEM) analysis of these HA powders showed aggregated, rough, and dense particles that were denser in synthetic HA, and the author believed that this difference could be due to the crystallinity of calcium nitrate, which was used as precursor of calcium. Similarly, the TEM analysis showed that the dispersion of HA crystallite was not good enough, one possible reason of this agglomeration could be van der Waals attraction. This phenomenon can be justified if we compare the crystallite particles size; 20e40 nm particle size was recorded for synthetic HA, and 71, 52, 97, and 28 nm HA particle size was recorded for human, bovine, camel, and horse bone HA. Comparatively, synthetic HA has smaller and irregular particle size that resulted into more van der Waals attraction as it was also observed in SEM analysis that the synthetic HA showed denser agglomeration. The elemental analysis of these synthesized and natural HA showed that the elements other than the calcium and phosphorus are in very little quantity in synthetic HA (Rahavi et al., 2017). This low concentration of elements cannot alter the overall biocompatibility of the synthetic material (Santos et al., 2004). On the other hand, these elements have higher weight percentage in HA extracted from natural bones. Infact, this is one of the main differences in natural and synthetic HA. Presence of these trace elements in HA could be beneficial for bone implants (Doostmohammadi et al., 2011). Ca/P ratio can also be calculated from elemental analysis data. The synthetic HA showed 1.75 Ca/P, while the theoretical Ca/P ratio is 1.67. It is difficult to control the exact stoichiometric ratio in HA because it depends on the reaction procedures and conditions (Rajkumar et al., 2011). The 1.75 and above Ca/P ratio with presence of CaO could be the result of heat treatment of the final product at 700 or above (Sanosh et al., 2009). This change can also be confirmed from XRD pattern of synthetic material as described before. The possible reason of this disturbed ratio could be the volatile nature of the phosphorus precursor above 650 C. So, the calcium precursor, which is not converted or incorporated into complex structure converts into CaO. The Ca/P

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ratio in HA derived from natural source is higher than 1.67; the scientists suggest that this higher ratio could be due the presence of other calcium phosphates with higher Ca/ P ratio (Doostmohammadi et al., 2012). In a study, it was observed that TCP and calcium oxide is present in HA derived from natural bones (Joschek et al., 2000). Here, it should be noteworthy that the deviation of actual Ca/P ratio in HA, from the theoretical ratio, is not necessarily due to the existence of TCP but it can also happen because of the presence of hydrogen phosphate and carbonate instead of phosphate and hydroxide (Landi et al., 2000). Because the XRD data of HA can lead us to ambiguous conclusion due to presence of a little number of carbonates in the sample, Fourier-transform infrared spectroscopy (FTIR) was utilized. FTIR study could be helpful in diagnosis of HA and can tell what sort of functional groups are present. FTIR studies showed that natural HA had a peak at 879 cm1, which could be due to the presence of hydrogen phosphate. Hydrogen phosphate could be a reason of deviated stoichiometric ration of Ca/P in natural bone HA as we discussed before (Rahavi et al., 2017). Another way to determine the structural and compositional details of synthetic and natural HA is Raman spectroscopy. Khan et al. summarized the Raman studies of synthetic apatite in comparison with natural bone. According to this review, synthetic HA showed some differences from natural bone, e.g., hydroxyl stretch that was present in synthetic HA was absent in natural bones moreover carbonate contents were greater in natural bones (Khan et al., 2008, 2013). However, Awonusi et al., in a study, compared the Raman spectra of bovine cortical bone and carbonated HA (see Fig. 4.1) (Awonusi et al., 2007).

4.1.1 4.1.1.1

Wet-chemical methods Solegel methodology

Solegel is a technique in which a reactant or mixture of reactants is/are hydrolyzed and passed through a solution state and a gel state before turning into a final product. This technique is widely used to synthesize the metal oxides, glass, and ceramics (see Table 4.1). The solegel process can be explained by a series of discrete steps. In the first step, sol is formed form the stable solution of alkoxide or metal precursor. In second step, a gel is formed because of the oxide or alcohol-bridge network by polycondensation or polyesterification, which increases the viscosity of the solution. The third step brings the solidification of reactants; this step is also called aging. After aging, drying and dehydration is carried out. In the final step, the product is treated with high temperature (>800 C) (Cushing et al., 2004). Chemical homogeneity of starting materials is improved by solegel method by molecular level mixing as the polycondensation or polyesterification occurs (Hench, 1997; Hench and West, 1990; Niederberger, 2007). Solegel is also a convenient method to produce irregular surfaces in the form of coatings on implants to mimic the natural HA. This method can be used for dip coating, a suitable substrate such as titanium alloys and 316L stainless steel can be dip coated repeatedly, followed by sintering (Aksakal and Hanyaloglu, 2008;

Handbook of Ionic Substituted Hydroxyapatites

400

500

600

1003 1030 1047 1071

852 878

a 1047 1070

582 590 610 581 590 609

431 450 430 449

Intensity

960

88

700

800

900

Raman shift (cm

1000

1100

b

1200

–1)

Figure 4.1 Raman spectra of (a) cortical bone of bovine and (b) synthetic carbonated apatite (Awonusi et al., 2007). Copyright 2007. Reproduced with permission from Springer Nature.

Table 4.1 List of methods of synthesis with references. No.

Synthesis method

References

1

Solegel

Hench, 1997; Hench and West, 1990; Niederberger, 2007; Aksakal and Hanyaloglu, 2008; Gan and Pilliar, 2004; Liu et al., 2001; Masuda et al., 1990; Deptula et al., 1992; Hu et al., 1993; Breme et al., 1995; Russell et al., 1996; Kordas and Trapalis, 1997; Jillavenkatesa and Condrate, 1998

2

Coprecipitation

Donadel et al., 2005; L
opez-Macipe et al., 1998;
osarczyk et al., 1997; Tas and Aldinger, 2005; Sl
Afshar et al., 2003; Lazi
c, 1995; Pang and Bao, 2003; Phillips et al., 2003; Saeri et al., 2003; Nazir et al., 2011; Khan et al., 2015

3

Emulsion technology

Phillips et al., 2003; Koumoulidis et al., 2003; Segal, 1997; Li et al., 2008; Lim et al., 1997, 1999; Jarudilokkul et al., 2007; Sun et al., 2007

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Table 4.1 Continued No.

Synthesis method

References

4

Batch hydrothermal process

Segal, 1997; Byrappa and Adschiri, 2007; Kaya et al., 2002; Daoud et al., 2005; Takami et al., 2007; Guo and Xiao, 2006; Arce et al., 2004; Liu et al., 1997, 2003; Riman et al., 2002; Wang et al., 2006; Yan et al., 2001; Zhang et al., 2005

5

Continuous hydrothermal flow synthesis

Chaudhry, 2008; Darr and Poliakoff, 1999; Ziegler et al., 2001; Chaudhry et al., 2008; Lester et al., 2013

6

Solid-state synthesis

Jia et al., 2002; Cao et al., 2005; Chen et al., 2004; Shu et al., 2005; Core~no A et al., 2005; Silva et al., 2003; Rhee, 2002; Suchanek et al., 2004; Fernandes et al., 2005; Li-yun et al., 2005

Gan and Pilliar, 2004; Liu et al., 2001). HA particles can also be produced using this method. During the solegel reaction, hydrolysis, condensation, and aggregation happen simultaneously, which makes its chemistry very complex, therefore reproducibility and control over the particle morphology is difficult. Moreover, the heat treatment of amorphous product alters the particle morphology, changes the crystallinity, and inducea agglomeration (Niederberger, 2007). As we already have discussed the importance of the HA in previous section, in this section, we will discuss the different reaction conditions to synthesize HA using sole gel technique. The morphology and other properties of HA prepared using solegel technique is highly dependent on the temperature of reaction, precursors, time of reaction, drying temperature, pH of the starting materials and solegel, and temperature and time of heat treatment (calcination). Different scientists have explored different aspects of this technique. Following, we will try to make a comparison of these different conditions and summarize the results periodically. Masuda et al. synthesized HA using solegel technique. They used alkoxides as the stating material (see Table 4.1). In this study, the authors revealed that the pH of the starting material was the most important factor that directly affected the HA synthesis. So, they used alkaline, neutral, and acidic reaction conditions. It was observed that only HA crystals without any additional phase were synthesized in neutral conditions. After the successful preparation, the HA was calcined at 600 C, resulting in to platelike structure having 1000Å size, which resembled to HA present in living tissues. This study was published in Japanese language; therefore, detailed information of reaction conditions could not be acquired (Masuda et al., 1990). Deputla et al. prepared spherical HA by solegel technique using calcium acetate and phosphoric acid as the starting material. The synthesized spherical nanoparticles had diameter less than 100 mm. In this study, the authors emulsified the freshly prepared precursor solutions using 2-ethylhexamol, and the resulted solegel was dehydrated at 50 C. The resulting

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powder was calcined at 750 C (Deptula et al., 1992). Hu et al. synthesized phosphaterich glass powder, which was further treated with calcium hydroxide. Different conditions were applied to this solegel synthesis, and it was observed that HA was produced at 60 C and 60 min of reaction time. The authors also used ultrasonication for few minutes to accelerate the synthesis (Hu et al., 1993). Breme et al. used HA as coating agent on titanium implants. First, the implant was coated with commercially available HA with or without bounding agent (phosphoric acid). The implant was dip coated on respective suspension and dried on room temperature followed by drying at 600 C and finally sintered at 1200 C. In the second technique, Breme et al. used solegel technique to produce HA on the surface of titanium implant. The implant was dipped in the solution of starting materials, which were CaO as calcium source and PO(OC2H5)3 or PO(OCH3)3 as phosphate source. The thin film formed on the surface of alloy was dried at 130 C for 1 h, which resulted in the formation of gel. Finally, the implant was annealed at 600e800 C for 5e15 min. The HA was characterized by X-ray studies (Breme et al., 1995). Russel et al. used ion beam technology to modify HA, which was prepared by solegel method. In this study, the HA was produced using calcium nitrate tetrahydrate and n-butyl acid phosphate as starting materials. The reaction mixture was spin-coated on the silicon substrate and immediately dried at 350 C followed by annealing at 500, 600, 700, 800, 900, and 950 C in flowing nitrogen environment (Russell et al., 1996). Kordas et al. synthesized HA by solegel method. This time, the HA was synthesized by modification of already established method. Calcium acetate and PO(OC2H5)3 were used as starting material, while different types of alcohols (methyl, ethyl, and propyl alcohol) were used as solvent. After the gel formation, the drying was carried out at 75 C for few days and dried gel was heat treated at 930 C (Kordas and Trapalis, 1997). Jillavenkatesa et al. synthesized solegel-based HA using calcium acetate and triethyl phosphate, though HA with these precursors and method was already reported, but the main purpose of this study was to produce cost-effective and easy-to-synthesize HA. The gel prepared in this method was colloidal in nature, so the difficulty associated with the slow dissociation of phosphate ion was resolved (Jillavenkatesa and Condrate, 1998). Weng et al. used P2O5 to prepare HA by solegel method. For this purpose, the P2O5 was dissolved in ethanol and refluxed for 24 h. Ca(NO3)2$4H2O was used for calcium source. Both reactant solutions were mixed, and resulting solution was refluxed for another 24 h. Resulting solution was dipcoated on appropriate substrate and dried at 150 C for 15 min. The coated substrate was calcined at 500 C for 15 min. To increase the thickness, the substrate was coated again, and the process was repeated for 10 times. Finally, the product was calcined at 700 C.

4.1.1.2

Coprecipitation

Coprecipitation is a process in which two or more than two compounds are precipitated simultaneously in a solvent. The coprecipitation is usually a result of supersaturation, which involves nucleation, growth, and agglomeration of particles in a solution. Most important part of coprecipitation is nucleation, which is the start of particle formation after which secondary processes such as Ostwald ripening and aggregation start.

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These secondary processes are responsible for determining the size morphology and properties of particles. Particle size distribution, morphology, and particle sizes are dependent on reaction conditions such as rate of reactant addition and stirring. Rate of nucleation is inversely proportional to the particle size. Usually, soluble calcium and phosphorus salts are coprecipitated by the addition of a base to result in poorly crystalline HA, which needs proper aging to attain stoichiometric ratio (Cushing et al., 2004). Generally, the sources of calcium ions are calcium nitrate, calcium hydroxide, and calcium chloride. For phosphate ions, usually diammonium hydrogen phosphate, phosphoric acid, potassium hydrogen phosphate, and diammonium hydrogen phos
osarczyk et al., phate are used (Donadel et al., 2005; L
opez-Macipe et al., 1998; Sl
1997; Tas and Aldinger, 2005) (see Table 4.1). The amorphous precipitates are produced by the immediate reaction of calcium and phosphate ions. Normally, the salts are mixed together for 2 h keeping the pH between 10 and 11. It has been often observed that the pH of the reaction system is more dominant in determining the stoichiometry and thermal stability of the resulting calcium phosphate as compared to the Ca:P molar ratio. The initial precipitate suspension is then aged for up to a day to reach the proper stoichiometric ration of calcium and phosphate. The process could be sped up by increasing the temperature up to 95 C (Afshar et al., 2003; Lazi
c, 1995; Pang and Bao, 2003; Phillips et al., 2003; Saeri et al., 2003). During coprecipitation synthesis, agglomeration could be a result of local concentration due to inhomogeneity, which can be avoided by fast stirring and low addition rate of reactants. Moreover, reaction pH plays an important role in determining the stoichiometry of HA. A strict control of pH is important because stoichiometric HA is only obtained at pH above then 10 (Afshar et al., 2003). Microwave assisted method has also been used to reduce synthesis duration for synthesis of thermally stable HA and magnesium-substituted HA (and calcium phosphates) (Nazir et al., 2011; Khan et al., 2015).

4.1.1.3

Emulsion technology

Emulsion is a type of liquid in which two distinct solvents (usually water and oil) or liquids are mixed together in such a way that fine droplets of one are dispersed in another homogenously. Mixture of water, oil, and a surfactant can result in to an emulsion. This emulsion looks clear because nanosized droplets do not scatter light noticeably. When water is dispersed in oil, it is called water-in-oil dispersion. The small water droplets in oil work as small reactors if used in a chemical reaction. These small droplet reactors have increased interfacial area as compared with conventional reactions (Cushing et al., 2004) (see Table 4.1). Both terminologies, microemulsion or macroemulsion, are used for emulsions. When the volume of water is very low in oil, it is called microemulsion. Microemulsions are usually clear and transparent and stable for a long time. But if the water is comparatively higher in volume, then it is called macroemulsion. Macroemulsions are opaque in nature and only stable on continuous stirring (Phillips et al., 2003).

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In an emulsion, when two water droplets containing reactants are collided with each other, chemical reaction takes place. The resulting product has small and homogeneous size when a reaction takes place in such a small droplet (Koumoulidis et al., 2003; Segal, 1997). It is reported that the powders produced by emulsion techniques have higher surface are as compared with those produced by other methods (Koumoulidis et al., 2003). Keeping in view the good homogenized particles produced by emulsion method, scientists have produces HA using this method. However, to carry out this kind of reaction, a large volume of oil (organic solvents) is required (which makes it rather a nonenvironment friendly method), and moreover, subsequent heat treatment is also required for crystallinity (Phillips et al., 2003; Koumoulidis et al., 2003; Li et al., 2008; Lim et al., 1997, 1999). Other parameters are also important for HA synthesis using this method, e.g., Jarudilokkul et al. reported that by increasing the reaction temperature (from 30 to 80 C), the surface are is decreased (from 227 to 98 m2g1) (Jarudilokkul et al., 2007). Similarly Sun et al. reported that using microemulsion technique in hydrothermal conditions and neutral pH, HA with rodlike morphology was produced, and when pH was increased, the particles morphology was converted to spherical (Sun et al., 2007).

4.1.1.4

Batch hydrothermal synthesis

In normal chemical reactions, the temperature of reactor could be increased maximum up to the boiling point of solvent, at which the reaction could be set on reflux. However, if the autogenous pressure is increased, the temperature of solvent can be elevated above its boiling point in a close container just like a pressure cooker. Materials that are insoluble in a certain solvents can be dissolved and recrystallized using this technique. The abovementioned technique is called batch hydrothermal process when water is used as a solvent. Hydro is derived from hydra which means water and thermal means related to heat (Cushing et al., 2004; Byrappa and Adschiri, 2007). Ceramic sols are resulted by the chemical reaction during a batch thermal process under high pressure and temperature usually in the presence of an acid or alkali. However, an acid or alkali is not always required though they have a catalytic effect on chemical reaction (Cushing et al., 2004; Kaya et al., 2002). Segal published a comparison of hydrothermal process with other processes for the synthesis of ceramic powders (Segal, 1997). The hydrothermal synthesis of ceramic powders has number of benefits over other methods. It is an environment friendly process (lies under green chemistry) and is a direct route to fabricate submicron nanoparticles at comparatively lower temperature without the need of subsequent calcination step. The hydrothermal process has been used for the synthesis of different products such as metal oxides, phosphates, and HA (Daoud et al., 2005; Takami et al., 2007; Guo and Xiao, 2006) (see Table 4.1). Guo and Xiao reported the synthesis of HA using batch hydrothermal process (Guo and Xiao, 2006). Calcium nitrate, calcium hydroxide, or calcium carbonate could be used as calcium source, while phosphate could be acquired from phosphoric acid, ammonium hydrogen phosphate, sodium dihydrogen phosphate, or calcium hydrogen phosphate. Solutions of calcium and phosphate source were mixed together and put in

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to closed container. Well-defined, crystalline and phase pure HA could be obtained by hydrothermal process, but the crystallinity and morphology of HA particles is depended on pH, pressure, temperature, and aging time during this type of synthesis. The processing temperature is reported between 90 and 150 C, while aging time is reported between few to 24 h (Guo and Xiao, 2006; Arce et al., 2004; Liu et al., 1997, 2003). Templating agents (e.g., poly(amidoamine), CTAB) can also be used in the hydrothermal batch processing of HA (Liu et al., 2003; Riman et al., 2002; Wang et al., 2006; Yan et al., 2001; Zhang et al., 2005).

4.1.1.5

Continuous hydrothermal flow synthesis

In early 1990s, researchers from Japanese group Arai at Tohoku University developed continuous hydrothermal flow synthesis (CHFS), and since then, few research groups around the globe have initiated use of this versatile technology. Researches have modified and improved CHFS and expanded its use for different materials systems. It is important to understand the basic science of supercritical water used in this technology (Chaudhry, 2008). According to definition, when the temperature and vapor pressure of a fluid exceed the critical point, it is called a supercritical fluid (SCF) (Darr and Poliakoff, 1999). The SCFs have unique set of properties that are totally different from normal gases and liquids. The viscosity and density of SCF can vary considerably when they are near the critical point. When the critical point is achieved, the difference between density of liquid and gas reduces, resulting in the disappearance of interfaces (Cansell et al., 2003; Noyori, 1999). At 374 C and 22 MPa, water becomes supercritical, and above that, it has variations in dielectric constant, viscosity, and ion product (Cushing et al., 2004; Bellissent-Funel, 2001). Water has the ability to dissolve many inorganic salts. However, when the temperature of water is increased on at a constant pressure, its dielectric constant is reduced. When water reaches its supercritical point, the inorganic salts become insoluble and precipitate out. At this point if the salts are supersaturated, the nucleation is increased, which results in the production of nanoparticles (Kashchiev and van Rosmalen, 2003). At this point, metal ions are converted to crystalline metal oxides because of rapid dehydration at high temperature in supercritical state. This could also be a possible reason of small size particles, as rapid dehydration at high temperature does not give enough time for crystal growth (Hakuta et al., 2003; Ziegler et al., 2001). Chaudhry et al. used CHFS technology to produce pure and magnesium-substituted HA for the first time. A series of magnesium substituted HAs (and calcium phosphates) were also produced. It was observed that HA with low magnesium concentration showed crystalline rodlike nanoparticles. Increasing concentration of magnesium up to certain level resulted in stabilization of the magnesium whitlockite (which normally does not precipitate directly), which appeared as a crystalline phase in addition to HA. Further increase in magnesium content led to phase purity (of this magnesium stabilized whitlockite). Generally, increasing magnesium concentration decreased crystallinity gradually (Chaudhry et al., 2008).

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Lester et al. used CHFS technology to produce nano-HA (nHA) of three different morphologies. These three morphologies were sheet, rod, and/or tube. It is described that the morphology of nHA was controlled by the controlling pH and/or reaction temperature. NH4OH was used to control the pH of system, which played an important role in determining the morphology of nHA (Lester et al., 2013) (see Table 4.1).

4.1.2

Solid-state methods

Jia et al. synthesized HA powder by low temperature solid-state reaction or mechanochemical reaction using calcium acetate and ammonium phosphate as precursors (Jia et al., 2002). This solid-state reaction was carried out on low temperature, though this kind of reactions can also perform on high temperature; however, these reactions can be accelerated significantly if treated with microwaves. Microwaves directly increase the molecular movement, which results in rapid temperature increase. So, Cao et al. synthesized HA from calcium nitrate and trisodium phosphate. First, they grounded the starting material and then heated using microwave technique. When the reactants were heated, using microwave, for half a minute, spherical morphology was obtained, which was converted to rodlike structures when microwave heating was extended to 1 min (see Table 4.1) (Cao et al., 2005). Solid-state reactions include mechanochemical mixing, which results into chemical reaction. When water is involved in a mechanochemical reaction, then it is called mechanochemicalehydrothermal reaction (Chen et al., 2004). Mechanochemicale hydrothermal reactions could be carried out in a ball mill. The reactants in appropriate stoichiometric ratio are mixed together and ball milled on a favorable temperature and pressure. Chemical reaction takes place because of contact of reactants between walls of mill and balls (and only balls also) (Shu et al., 2005). Mechanochemical reactions usually take a long time to complete. However, mechanochemical synthesis of HA takes few to 60 h to complete (Core~ no A et al., 2005; Silva et al., 2003), and to convert it into phase pure HA, a heat treatment is required (Rhee, 2002). This heat treatment can lower down the surface area as Chen et al. reported that heat treatment of HA at 900 C for 1 h reduced the surface area form 172 to 7.2 m2 g1 (Chen et al., 2004). The mechanochemical reactions can also be used to produce ion-substituted HA (Zhang et al., 2005; Shu et al., 2005; Core~ no A et al., 2005; Suchanek et al., 2004). Cai et al. reported that if dicalcium phosphate dihydrate and calcium phosphate are grinded together, carbonated HA nanofiber can be synthesized (Shu et al., 2005). Moreover, large quantities of crystalline HA powder can be obtained by solid-state reactions; however, there are chances of impurities that can be added by wearing and tearing of material used for mixing and grinding, e.g., balls of ball mill (Rhee, 2002; Fernandes et al., 2005).

4.1.3

Hydroxyapatite coatings

The metallic implants of titanium and its alloys show the greatest biocompatibility if we compare it with other materials. Usually, these implants are grouped with bioinert materials together with ceramics such as zirconia, titania, and alumina depending on

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the applications and pattern of osteogenesis. Some applications required good biointeraction for tightening and compact packing, in other words a strong bonding with natural bone. Calcium phosphateebased biomaterials such as TCP and HA show good biocompatibility. For this kind of fixation technique, there are some required properties such as mechanical strength, ductility, ease of fabrication, and biocompatibility (Kasuga and Niinomi, 2010). HA-coated implants have been widely used in dentistry and orthopedics. There are variety of methods that are used for coating of HA such as electrophoretic deposition (EPD) (Ducheyne et al., 1986), ion beam sputtering (Ong and Lucas, 1994), dip coating (Li et al., 1996), plasma spraying (Lacefield), thermal spraying (Mardali et al., 2019), and biomimetic coating (Kokubo, 1998) etc.

4.1.3.1

Plasma-sprayed coatings

In this technique, the HA is sprayed on the surface of implant using plasma. As the temperature of plasma is very high, it melts the HA that is meanwhile sprayed. Coatings with plasma technique have strong bonding (between metallic implant and HA) and good dissolution in body fluids. However, there is a large difference in the thermal expansion of metallic implant and coating, which results into residual stress at the interface. When the crystallinity of HA is low and it is in an amorphous state, the coating can dissolve rapidly, which could induce an inflammatory response. Tusi et al. reported a relatively high bonding strength on titanium alloys (Tsui et al., 1998). They optimized the conditions of coating and reported that the strength was between 20 and 40 MPa. These results were obtained because the crystallinity of HA in these experiments was high (83%e94%). For long-term benefits of an implant, the crystallinity of its coating is highly important. So, Ingaki et al. developed a radiofrequency thermal plasma spraying method to obtain strong interaction between HA and metallic implant substrate (Inagaki et al., 2008). The authors used HA/Ti composites for this purpose. HA and Ti were plasma sprayed on the desired substrate using two microfeeders. The coating was carried out in gradient composition, first Ti-rich plasma was sprayed followed by the HA-rich plasma. It was also observed that the plasma in nitrogen environment gave better strength to the coating. Moreover, a postheat treatment provided highly oriented HA, which showed better biocompatibility.

4.1.3.2

Electrophoretic deposition

EPD is a method of particles coating on a surface using electrochemistry. In this method, the particles, e.g., bioceramics, are suspended in a solvent that is transferred in an electrochemical cell, where one electrode is the surface that is to be coated and other is counter electrode. Because of the potential difference of electrodes, the suspended particles are polarized and attracted by the substrate where they deposited in the form of lose coating. To increase the bonding and crystallinity of ceramic particles, a postheat treatment is usually recommended (Augello et al., 2015; Ducheyne et al., 1990).

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Ducheyne and Radin coated the metallic implants with HA using EPD and heattreated the coatings at 900  C for 1 h (Ducheyne et al., 1990). They obtained uniform thicknesses, which were controlled by changing the electrical field strength and deposition time. The deposited ceramic coatings were hampered by the absorbed water, and a vacuum sintering led to phase transformation that, the author suggested, could provoke the considerable changes in the in vitro dissolution experiments. Mehboob et al. used polymer-assisted EPD of HA on medical-grade stainless steel. Polyethylene glycolemodified HA was used for EPD. The purpose of using polymer was to avoid postdeposition heat treatment (Mehboob et al., 2014). Similarly, Iqbal et al. used polyvinyl alcohol (PVA)ecoated HA to deposit on 361L stainless steel substrate using EPD technique. The addition of PVA improved the adhesion of the powder on substrate, which eliminated the heat treatment process (Iqbal et al., 2012).

4.1.3.3

Thin film coatings

Thin, uniform, and dense coatings of HA on metallic substrates can be obtained by physical vapor deposition, which is one of the most promising methods for coating. Calcium phosphate thin films were coated on titanium substrate using radiofrequency (RF) magnetron sputtering. This technique has been widely used for coatings of thin films with excellent bonding of coating materials with substrate (Yoshinari et al., 1997; Wolke et al., 2003). The RF technique was also used for titanium alloys substrate (van Dijk et al., 1995). This method has advantages over the other high-temperature techniques as it can be executed at relatively low temperature that does not alter the properties of substrate.

4.1.3.4

Biomimetic apatite coatings

Formation of a bioactive coating on the surface of any implant by soaking it in a simulated body fluid (SBF) with ion concentrations equal to those of human blood plasma (Kokubo, 1998) is called biomimetic method. Using this method, HA can be deposited on a surface, which is also called bone-like apatite. This method has advantages over conventional methods as materials can be homogenously coated on an implant without the need of heat treatmentdmaking this technique applicable to polymer surfaces as well. Although heat treatment is not applied in this technique, even then a good crystalline nanosized apatite layer was obtained. Formation of biomimetic HA (BHA) using SBF requires two conditions: (1) the surface must have such functional group that can induce nucleation of apatite, e.g., hydroxyl, carboxylic, and phosphate and (2) increased concentration of apatite (Kokubo, 1998; Kokubo et al., 1990). Metallic implants are usually covered with passive oxide layer, e.g., titanium is covered with titanium oxide. Because the surface of passive layer also has titanium hydroxide, therefore, there are great chances of HA nucleation on metallic implants. A large amount of calcium ions should also be dissolved in the SBF to increase the possibility of BHA formation.

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97

Properties of hydroxyapatite Structural insights

The chemical formula of HA is represented by Ca10(PO4)6(OH)2. A stoichiometric HA has monoclinic geometry with P21/b space group (Morgan et al., 2000), and HA having hexagonal system with P63/m space group has little deviation from stoichiometry (Hench). Most of the synthetic methods produce a slight nonstoichiometry and hence a hexagonal structure. Fig. 4.2 represents the sketch of a unit cell of hexagonal HA. Two different positions of calcium ions are mentioned (these refer to Ca(I) and Ca(II) positions). Phosphorus atoms are not shown in the figure because they are hidden by the space occupied by oxygen atoms. Overall, 18 ions are closely packed to make the hexagonal structure. At each hexagonal corner, a calcium ion is surrounded by 3 hexagons (i.e., 12 calcium ions shared by 3 hexagons ¼ 4 calcium ions per hexagon). There is a hydroxyl ion in the center of each unit cell (which makes two hydroxyls per unit cell). Hydroxyl group in center is surrounded by three calcium ion per hexagon forming a ring (i.e., six calcium ions). A “chord” is formed throughout the structure because of these rings that are responsible for many properties of HA. Void spaces between two hexagons are filled with three phosphate tetrahedra per unit cell. Ions in HA can be substituted (also referred to as doping interchangeably in literature) by biologically beneficial ions due to the inherent versatility of this crystal structure. Substitutions can be both cationic and anionic referring to substitutions of the calcium, phosphate, and/or hydroxyl ions.

Calcium Oxygen Hydrogen

Figure 4.2 Schematic representation of hydroxyapatite unit cell. The doted lines are representing the lower level (Sarig, 2004). Copyright 2004. Reproduced with permission from Elsevier Ltd.

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Physical and thermal properties

The physical properties of HA highly depend on the source of method from which HA is obtained or synthesized, as we have already discussed in first section. The HA synthesized in lab shows low crystallinity and high surface area and can have high porosity depending on the processing method used. On the other hand, natural HA is usually derived by heat treatment of bones at higher temperature (usually 800 C), which results in highly crystalline HA (Sobczak-Kupiec et al., 2012). Sintered HA derived from bovine bone has porosity and pore structure that matches the natural bone. This porosity and wettability of HA makes it useful for drug loading. The most stable calcium phosphate salt has a Ca:P molar ratio of 1.67 and normally precipitates at or above pH 10. However, the biological and synthetic apatite differs from the ideal stoichiometric ratio due to substitutions.

4.2.3

Mechanical properties

HA is usually implanted in the form of granules or porous scaffolds. However, it is not used as a load-bearing implant because of its mechanical properties, which limits its use (Lin et al., 2012). To improve the mechanical properties of HA is a great challenge of scientists so far. Usage of carbon nanotubes and different metal oxides such as alumina, zirconia, and titania as reinforcing agent is very common (Mukherjee et al., 2014; Mobasherpour et al., 2009). However, these reinforcing materials may compromise the biocompatibility of HA because usually they are bioinert or nonbiocompatible. On the other hand, HA whiskers or fibers are considered to have better mechanical properties without compromising the biocompatibility (Bose et al., 2009; Suchanek et al., 1997), but HA whiskers do not have thermal stability as they disappear after sintering (Lin et al., 2012). There is another effective way to enhance the mechanical properties of HA, which is the fabrication of nanosized HA. Nanosized ceramics have apparently higher mechanical properties as compared with microsized ceramics (Wang and Shaw, 2009). Another way to improve the mechanical properties of HA is to make composites with polymer and crosslinking agents. Iqbal et al. fabricated chitosan/HA sponge-like scaffolds using freeze-drying technique, and hydroxy propyl methyl cellulose was used as crosslinking agent (Iqbal et al., 2017). The fabricated sponges showed better compression strength and elasticity including good cell support and favorable degradability. Architecture of porous scaffolds also plays an important role in determining its strength (Bignon et al., 2003; Kalita et al., 2003). As the strength is a very important property especially for those scaffolds that are used for the replacement of load-bearing bone. Therefore, it is important to design scaffolds carefully (Hollister et al., 2002; Lin et al., 2004).

4.2.4

Biological performance

There is a clinical need of synthetic bone graft that has ability to stimulate the bone growth and act as temporary bone scaffold. Physiological tissues interact with implanted materials that determine the fate of implant in living system. An implant could be a failure if the boneeimplant interface is loosened. On contrary, the intimate contact of bone and implant and chemical interaction of implant with surrounding bone

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are considered as promising successful implantation. Therefore, the evaluation of any biomaterial for its performance in living organism or tissue culture is important so that in future any misbehavior of biomaterial could be preempted (Bizios, 1994). HA is considered as a well-tolerated material in both in vitro and in vivo studies.

4.2.4.1

In vitro

The HA is considered as a good biomaterial and harmless to cell environment. However, very fine particles of HA may damage the fibroblast cell, whereas large particle HA does not damage (Evans et al., 1992). This damage could be due to the direct contact of the cells with HA particles. Sautier et al. carried out in vitro study of isolated mouse calvaria cells on synthetic HA (Sautier et al., 1991). Initially an electron dense layer at the periphery of the material was observed, which was granular and collagen free. With the passage of time, this dense layer was converted to an amorphous granular form in between the HA aggregates. An osteoid matrix was observed, which was mineralized on the previously prepared layer. In vitro studies of HA/collagen bone-like composites showed excellent biocompatibility and biointegrative activities, which were equal to autogenous bone and much better than other materials. On the bases of in vitro studies, it was concluded that this composite was a potential material in medical and dental field (Kikuchi et al., 2004). Sonocoated scaffolds with nHA also showed increased proliferation of osteoblast cells (Rogowska-Tylman et al., 2019).

4.2.4.2

In vivo (animal model)

Rogowaska et al. prepared scaffolds that were coated with biological HA using sonocoating technique. HA-coated scaffolds were tested in vitro using rabbit model. The animal studies showed that HA-coated scaffolds showed new bone formation. Up to 68% volume of porous scaffolds was filled with new bone formed during in vitro experiment (Rogowska-Tylman et al., 2019). Cardoso et al. (2019) prepared porous scaffolds of calcium deficient HA with PCL using NaCl as leaching salt for porosity. HA was used in the form of needle-shaped whiskers. In vivo studies showed good biocompatibility. Rat model was used for in vivo studies. The in vivo studies confirmed the potential of osteogenesis. Oversall, up to 31% improvement in osteogenic properties was observed as compared with the control. Fang et al. prepared polymeric hydrogel and incorporated nHA by in situ synthesis (Fang et al., 2019). The rabbit animal model was used for in vivo testing. Regeneration of a highly mineralized bone tissue and direct bonding to the nHA containing hydrogel was observed in femoral condyle defect rabbit model. Ripamonti et al. (1992) used baboon model to study in vivo application of HA in combination with osteogenin. Osteogenin is a protein that plays an important role in osteogenesis. Porous HA replicas were obtained from the heat treatment of calcium phosphate from exoskeleton of coral. Osteogenin was obtained from bovine and delivered to bioresorbable and nonbioresorbable HA rods. For assessment of osteogenic activity, a total of 48 rods were bioassayed into eight baboons. Osteogenin fractions were reconstructed with baboon-insoluble collagenous bone matrix, which was implanted in additional four baboons. The results demonstrated that differentiation of bone was observed in nonbioresorbable rods with or without ostegenin. However, bioresorbable HA did not support any bone

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formation even in the presence of osteogenin. Thus, nonresorbable HA replicas and reconstituted insoluble baboon collagenous bone matrix with bovine osteogenin developed the potential therapeutic possibility of bone induction in different areas such as craniofacial, periodontal, and surgery for orthopedic reconstruction.

4.2.4.3

Clinical trials

To evaluate medical, surgical, or behavioral intervention in human, clinical trials are an essential step. A biomaterial that has passed in vitro and in vivo studies successfully can be considered for clinical or human trials in confluence with regional ethical and regulatory permissions. Clinical or human trials of HA have been studied extensively for use in different areas of the human body. These include (but are not limited to) dentine pulp regeneration (Nakashima et al., 2019), impact of implant length on marginal bone lose in atrophied arches (Amine et al., 2019), bone grafting in orthognathic surgery (Alyahya and Swennen, 2019), alveolar ridge preservation, sinus augmentation, periodontal bone defect (Dewi and Ana, 2018; Zhou et al., 2018a), total knee arthroplasty (Zhou et al., 2018b), and anterior cervical discectomy and fusion (Zadegan et al., 2017). Some key reports are further discussed below for elucidation. Pines et al. (1984) carried out clinical trials on 40 patients at risk of osteoporosis because of long-term treatment with prednisolone to determine the efficacy and tolerance of crystalline micro-HA to avoid the appearance or progression of osteoporosis. A total of 32 patients were administrated with 6e8 g HA for 12 months, and 8 patients were served untreated as control group. The results confirmed the tolerance of HA and the superior acceptability of HA as tablet form. Moreover, it was observed that skeletal pain in patients developing osteoporosis was dramatically reduced. Livesley et al. conducted clinical trials to compare the use of HA-coated and uncemented bipolar hemiarthroplasty for the treatment of a displaced subcapital fracture of the femur (Livesley et al., 1993). A total of 82 patients suffering from subcapital fractures of the femur were included in trials. The results were superior significantly in HA-coated group after 1 year postoperative follow-up. Jones et al. (1997) preformed the clinical trial of cylinder implants in different region of mouth. Comparison of plasma-sprayed Ti and HA-coated plasma spray Ti implants was studied in these trials. A total of 65 subjects were included in this study. A total of 352 implants (equally distributed between with and without HA-coated, plasma-sprayed Ti) were implanted in four different sites. The results suggested that HA-coated, plasma-sprayed Ti implants showed superior integration in the beginning. Similarly, Nilsson et al. preformed clinical trials on 57 subjects (Nilsson et al., 1999). It was a randomized study to compare the fixation of HA coated with cemented tibia components in the Tricon II total knee arthroplasty. Radiostereometric analysis was adopted to evaluate the quality of fixation for 5 years after surgery. There was one early delamination of the coating and clinical lessening and two implants got infection. Eight patients (a total of nine knees) died, one patient sustained a stroke, and one patient refused to be investigated after 1 year. In remaining patients, the magnitude of the micromotion between the HA-coated and cemented groups did not differ. Most of the migration was shown within 3 months, in HAcoated implants, and then stabilized, and on the other hand, cemented implants showed initially lower but over the time continuously increasing migration.

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4.2.5

101

Applications of hydroxyapatite

In previous sections, we have discussed the HA-based tissue engineering applications. Following are some other applications of HA.

4.2.5.1

Hydroxyapatite as a drug carrier

HA has been explored in research as viable delivery platforms for clinical applications. HA is a material of particular interest due to its alluring features: interconnecting porosity of bulk material, good biocompatibility, resistance to mechanical force, and sustained release capacity Researchers have been investigating HA nanoparticles for different clinical applications such as immunoadjuvant therapy, drug delivery, and gene/siRNA transportation. Using HA as therapeutic drug/antigens/protein carrier provides slow rate of release and protection from degradation (Shinto et al., 1992; Netz et al., 2001; Komlev et al., 2002; Bhattarai et al., 2008; Barroug and Glimcher, 2002). Barroug and Glimcher reported the adsorption and desorption kinetics of cisplatin in vitro. They prepared plate-like nHA particles having dimension of 93 nm  23 nm. These particles were loaded with the drug for 40 h in ambient conditions. The release profile of the drug showed that after 102 h, only 33% of total absorbed drug was released. The extended studies showed that after 14 days, a total of 58% drug were released (Barroug and Glimcher, 2002). Similarly, Palazzo et al. reported the viability of HA as a drug delivery vector by investigating the adsorption and desorption kinetics of anticancer drugs cisplatin (CDDP), di(ethylenediamineplatinum) medronate, and bisphosphonate alendronate on nHA. The investigators prepared needle-like and plate-like nanoparticles of HA using modified wet-chemical precipitation method. The drug release mechanism of these nHA particles was controlled by morphology and by tuning the surface charges. Positively charged cisplatin was electrostatically bound to plate-like nHA, and counterrally negatively charged alendronate was attached to needle-like nHA through ligand exchangeeinduced binding mechanism. Approximately, 90% of initial drug concentration of alendronate and CDDP was loaded on needle-like nHA with regard to surface area. As the needle-shaped nHA has more negative surface charge, which resulted in more electrostatic interaction between nHA and drug therefore, it was able to load more CDDP (Barroug and Glimcher, 2002; Palazzo et al., 2007). Plasmid DNA and siRNA are degraded in living tissues, therefore their delivery to specific area could be a challenge. Zhe et al. demonstrated that nHA can be used to deliver plasmid DNA (pDNA) in vitro and in vivo. So, rod-shaped nHA (40e60 nm) was synthesized, which was able to carry 1 mg of plasmid DNA per 36 mg. These nanoparticles were used to carry EGFP-N1 pDNA to the required site of gastric cancer cells (SGC-7901). It was observed that the cancer cells were transfected in comparis on to free pDNA, which was used as control. Fluorescent proteins were also successfully expressed on the required side in vitro (in mice model), which also indicates the flexibility of nHA in gene delivery (Zhu et al., 2004). nHA was also used as a nonviral vector in gene delivery. Bisht et al. investigated that 100e120 nm HA particles were able to make complex with pDNA, and this complex showed that expression of green fluorescent protein was increased over a period of 24 h in HeLa cells, which is an evidence of sustain release. The possible reason

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behind this phenomena is that the HA was able to encapsulate the pDNA and protected it from cellular environment and successfully transferred it to nuclear space (Bisht et al., 2005). Small interfering ribose nucleic acid or simply siRNA is important in blocking the impression of different receptors. However, siRNA could be disintegrated in the cell medium, so Yand et al. investigated the ability of nHA to carry siRNA to block NR2B expression, a receptor important in chronic pain. He loaded one of anti-NR2B-siRNA to 35 mg of rod-shaped nHA (40e50 nm). Mice model was used to evaluate the in vivo studies. It was observed that the nociceptive behavior was present even after 7 days of injection, which confirms the sustained release of NR2BsiRNA (Yang et al., 2008). To induce the immune response in the body, antigens are used, e.g., for vaccination purposes. Hepatitis B is a big concern nowadays especially in developing countries. Transportation of hepatitis B surface antigen (HBsAg) was done using cellobiosecoated nHA ranging from 50 to 150 nm by Goyal et al. It was observed that cellobiose-coated nHA was able to load approximately 20% HBsAg, while up to 50% HBsAg was loaded on noncoated nHA. For in vivo studies, HBsAg (10 mg) loaded on coated and noncoated nHA and free HBsAg was administrated subcutaneously in mice. Cellobiose-coated nHA proved to be a more effective adjuvant, and it induced rapid antibody response, which showed that cellobiose-coated nHA can act as an efficient antigen delivery vehicle (Goyal et al., 2006). It is believed that cellobiose coating protected the conformation of specific protein, therefore bioactivity of HBsAg was maintained (Goyal et al., 2006; Kossovsky et al., 1996).

4.2.5.2

Bioimaging

Bioimaging is a process in which the biological systems are noninvasively monitored by visualization in real time. Florescent dyes and other organic compounds are used for bioimaging. For a better performance, it is important to encapsulate the bioimaging material, e.g., a dye in nanoparticulate system (Morgan et al., 2008). The nanoparticulate system can protect the dye in vitro until it reaches its destination. Besides the ability of HA to carry drugs, it also has been used to carry bioimaging molecules. Because the formulation of HA involves a simple chemical reaction, encasement of bioimaging molecules can be achieved by only mixing that molecule during the formation of nHA. Moreover, formation of amorphous HA under the typical synthesis conditions allows the inclusion of a broad range of molecules such as organic fluorophores or other low molecular weight molecules (Panyam and Labhasetwar, 2003; Tung and Amjad, 1998). Morgan et al. reported that nHA (20e30 nm) can be synthesized and loaded with a number of fluorescent dyes. Specifically, he used Cascade Blue, Cy3 Amidite, Rhodamine WT, fluorescein sodium salt, and 10-(3-sulfopropyl) acridinium betaine (SAB) to encapsulate in nHA. For free and encapsulated dye, nearly a fourfold increase in fluorescence quantum efficiency from 0.045 to 0.202 was observed, respectively (Morgan et al., 2008). Another important way to exploit nHA for bioimaging is conjugation of radioisotopes to nHA. Ethylene diamine tetramethylene phosphate has shown high attraction to 153Sm and bone and/or HA and therefore offers a potential avenue to develop nHA conjugated with radioisotopes with Ethylene-diaminetetramethylene-phosphonate (EDTMP) ligands for both bioimaging and radiotherapy functions (Loo et al., 2010).

Basics of hydroxyapatite

4.2.5.3

103

Commercial products

A great number of bioceramics and calcium phosphateebased biomaterials are either available or under consideration under different trade names (see Table 4.2) for orthopedic and dentistry applications (Dorozhkin, 2013a). For example, for augmentation, tooth replacement, alveolar ridge augmentation, and maxillofacial reconstruction, bulk materials in dense and porous form are available (Dorozhkin, 2013b). Other applications include bone defects repair (Silva et al., 2005; Damron, 2007), buttons for burr hole (Easwer et al., 2007; Kashimura et al., 2010), spine fusion (Thalgott et al., 1999; Mashoof et al., 2002; Minamide et al., 2005), orbital implants (including Bio-Eye) (Jordan et al., 2004; Yoon et al., 2008; Tabatabaee et al., 2011; Kundu et al., 2013), and hearing ossicles increment (Wehrs, 1991; Smith et al., 2002; Doi et al., 2007). The trade names and producer of commercial orthophosphates are given below in Table 4.2. Table 4.2 Trade names and producers of calcium phosphate products (Dorozhkin, 2013a). Calcium phosphate

Trade name and producer (when available)

Hydroxyapatite (HA)

BoneSource (Stryker Orthopedics, NJ, USA) Calcitite (Zimmer, IN, USA) Cerapatite (Ceraver, France) Durapatite (unknown producer) HA BIOCER (CHEMAeELEKTROMET, Poland) HAnano Surface (Promimic, Sweden) nanoXIM (Fluidinova, Portugal) NEOBONE (Covalent Materials, Japan) OssaBase-HA (Lasak, Czech Republic) Ostegraf (Ceramed, CO, USA) Ostim (Heraeus Kulzer, Germany) Periograf (Cooke-Waite Laboratories, USA) PermaOS (Mathys, Switzerland) PurAtite (PremierBiomaterials, Ireland) Synatite (SBM, France) Synthacer (KARL STORZ Recon, Germany) Without trade name (Cam Bioceramics, Netherlands) Without trade name (CaP Biomaterials, WI, USA)

HA embedded in silica gel

NanoBone (Artoss, Germany) Continued

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Table 4.2 Continued Calcium phosphate

Trade name and producer (when available)

HA/collagen

Bioimplant (Connectbiopharm, Russia) Bonject (Koken, Japan) Collagraft (Zimmer and Collagen Corporation, USA) CollapAn (Intermedapatite, Russia) HAPCOL (Polystom, Russia) LitAr (LitAr, Russia)

HA/sodium alginate

Bialgin (Biomed, Russia)

HA/poly-L-Lactic acid

SuperFIXSORB30 (Takiron, Japan)

HA/polyethylene

HAPEX (Gyrus, TN, USA)

HA/CaSO4

Hapset (LifeCore, MIN, USA) PerOssal (aap Implantate, Germany)

Coralline HA

Interpore (Interpore, CA, USA) ProOsteon (Interpore, CA, USA)

Algae-derived HA

Algipore (Dentsply Friadent, Germany)

Bovine bone apatite (unsintered)

Bio-Oss (Geitslich, Switzerland) Laddec (Ost-Developpement, France) Lubboc (Ost-Developpement, France) Oxbone (Bioland biomateriaux, France) Tutoplast (Tutogen Medical, Germany)

Bovine bone apatite (sintered)

BonAP (unknown producer) Cerabone (aap Implantate, Germany) Endobon (Merck, Germany) Navigraft (Zimmer Dental, USA) Osteograf (Ceramed, CO, USA) PepGen P-15 (Dentsply Friadent, Germany) Pyrost (Osteo AG, Germany)

Hyman bone allograft

maxgraft (botiss, Germany) Osnatal (aap Implantate, Germany)

Equine

BioGen (unknown producer)

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Table 4.2 Continued Calcium phosphate

Trade name and producer (when available)

a-tricalcium phosphate (TCP)

BioBase (Zimmer, IN, USA) Without trade name (Cam Bioceramics, Netherlands) Without trade name (PremierBiomaterials, Ireland)

b-TCP

adboneTCP (Medbone Medical Devices, Portugal) Antartik TCP (MedicalBiomat, France) Augment Bone Graft (BioMimetic Therapeutics, TN, USA) BioGraft (IFGL BIO CERAMICS, India) Bioresorb (Sybron Implant Solutions, Germany) Biosorb (SBM S.A., France) Bi-Ostetic (Berkeley Advanced Biomaterials, CA, USA) Calc-i-oss classic (Degradable Solutions, Switzerland) Calciresorb (Ceraver, France) CELLPLEX (Wright Medical Technology, TN, USA) Cerasorb (Curasan, Germany) Ceros (Thommen Medical, Switzerland) ChronOS (Synthes, PA, USA) Conduit (DePuy Spine, USA) GenerOs (Berkeley Advanced Biomaterials, CA, USA) HT BIOCER (CHEMAeELEKTROMET, Poland) JAX (Smith and Nephew Orthopedics, USA) Osferion (Olympus Terumo Biomaterials, Japan) OsSatura TCP (Integra Orthobiologics, CA, USA) PORESORB-TCP (Lasak, Czech Republic) SynthoGraft (Bicon, MA, USA) Synthos (unknown producer) Syntricer (KARL STORZ Recon, Germany) Vitoss (Orthovita, PA, USA) Without trade name (Cam Bioceramics, Netherlands) Without trade name (CaP Biomaterials, WI, USA) Without trade name (Shanghai Bio-lu Biomaterials, China)

b-TCP/collagen

Integra Mozaik (Integra Orthobiologics, CA, USA) Continued

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Table 4.2 Continued Calcium phosphate

Trade name and producer (when available)

Biphasic Calcium Phosphate (BCP) (HA þ b-TCP)

4Bone (MIS, Israel) adboneBCP (Medbone Medical Devices, Portugal) Antartik Genta (MedicalBiomat, France) Artosal (aap Implantate, Germany) BCP BiCalPhos (Medtronic, MN, USA) BioGraft (IFGL BIO CERAMICS, India) Biosel (Depuy Bioland, France) BonaGraft (Biotech One, Taiwan) BoneCeramic (Straumann, Switzerland) BoneSave (Stryker Orthopedics, NJ, USA) Calcicoat (Zimmer, IN, USA) Calciresorb (Ceraver, France) Calc-i-oss crystal (Degradable Solutions, Switzerland) CellCeram (Scaffdex, Finland) Ceraform (Teknimed, France) Ceratite (NGK Spark Plug, Japan) CuriOs (Progentix Orthobiology BV, Netherlands) Eurocer (FH Orthopedics, France) GenPhos HA TCP (Baumer, Brazil) Graftys BCP (Graftys, France) Hatric (Arthrex, Naples, FL, USA) Hydros (Biomatlante, France) Indost (Polystom, Russia) Kainos (Signus, Germany) MasterGraft Granules (Medtronic Sofamor Danek, TN, USA) MBCP (Biomatlante, France) OrthoCer HA TCP (Baumer, Brazil) OpteMX (Exactech, FL, USA) OsSatura BCP (Integra Orthobiologics, CA, USA) ossceram nano (bredent medical, Germany) Osteosynt (Einco, Brazil)

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Table 4.2 Continued Calcium phosphate

Trade name and producer (when available) Ostilit (Stryker Orthopedics, NJ, USA) ReproBone (Ceramisys, UK) SBS (Expanscience, France) Scaffdex (Scaffdex Oy, Finland) TCH (Kasios, France) Triosite (Zimmer, IN, USA) Tribone (Stryker, Europe) Without trade name (Cam Bioceramics, Netherlands) Without trade name (CaP Biomaterials, WI, USA)

BCP (HA þ a-TCP)

Skelite (Millennium Biologix, ON, Canada)

BCP (HA þ b-TCP)/ collagen

Allograft (Zimmer, IN, USA)

BCP/fibrin

TricOS (Baxter BioScience, France)

BCP/silicon

FlexHA (Xomed, FL, USA)

Fluorapatite (FA)

Without trade name (CaP Biomaterials, WI, USA)

FA þ BCP (HA þ b-TCP)

FtAP (Polystom, Russia)

Carbonate apatite

Healos (Orquest, CA, USA)

Collagraft (Zimmer, IN, USA)

SRS (Norian, CA, USA)

References Afshar, A., et al., 2003. Some important factors in the wet precipitation process of hydroxyapatite. Mater. Des. 24 (3), 197e202. Akram, M., et al., 2014. Extracting hydroxyapatite and its precursors from natural resources. J. Mater. Sci. 49 (4), 1461e1475. Aksakal, B., Hanyaloglu, C., 2008. Bioceramic dip-coating on Tie6Ale4V and 316L SS implant materials. J. Mater. Sci. Mater. Med. 19 (5), 2097e2104. Alyahya, A., Swennen, G.R.J., 2019. Bone grafting in orthognathic surgery: a systematic review. Int. J. Oral Maxillofac. Surg. 48 (3), 322e331. Amine, M., et al., 2019. Short implants (5e8 mm) vs. long implants in augmented bone and their impact on peri-implant bone in maxilla and/or mandible: systematic review. J. Stomatol. Oral Maxillofac. Surg. 120 (2), 133e142. Arce, H., et al., 2004. Effect of pH and temperature on the formation of hydroxyapatite at low temperatures by decomposition of a CaeEDTA complex. Polyhedron 23 (11), 1897e1901.

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Role of substitution in bioceramics

5

Sobia Tabassum Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Pakistan

5.1

Introduction

The bone grafts offer fast healing where natural healing process is slow and even not possible. Autografts stand as the most successful materials for the regeneration of defected tissues. Although presence of minerals, living cells, and growth factors in autografts help to induce excellent osteoconduction and osteoinduction, but necessity of additional surgery and limited supply restricts their frequent use. The allografts help to overcome some of these limitations, but possibility of disease and infection transfer from donor prompts the use of alternative bone substitute materials. The mineral composition of natural bone provides a road map for the construction of innovative bone grafts. Hydroxyapatite (HA) constitutes 70% of a typical natural bone in composition. However, this HA exists in substituted form (containing ions such as carbonate, fluoride, magnesium, sodium, zinc etc). It is different from synthetic pure HA in terms of stoichiometry, mineral composition crystal size, and physiochemical properties. Cellular interactions with a bone graft play a key role in regeneration of damaged tissues. Substituted ions in HA lattice play a vital role in biological responses of bone þ cells; CO2
3 affects the formation of apatite crystals; Na influences on bone remodel2þ ing; and Zn stimulates bone formation and prevents bone resorption. Biological and physicochemical properties of synthetic HA can be improved and tuned via doping of different metallic ions, nonmetallic ions and oxides, small functional compounds, peptides, and polymers.

5.2

Bioapatites

Bioapatite constitutes the mineral phase of bone, dentin, and enamel. A hydrated domain is present around the bioapatite that contains relatively mobile ions (Kþ, 2
Naþ, Mgþ2, F
, Cl
, HPO2
4 , and CO3 ). These ions are participating in substitution reactions of bioapatite. Degree and type of substitution control structural strength and function of bioapatite. Bioapatite in the human bone can be best described as 2
(Ca,Z)10(PO4,Y)6(OH,X)2, where Z ¼ Naþ, Mg2þ, Kþ, Sr2þ; Y ¼ CO2
3 , HPO4 ;

and X ¼ Cl , F . For comparative look, mineral concentrations and lattice parameters of bioapatite and stoichiometric HA are listed in Table 5.1 (Dorozhkin and Epple,
2002; Supov a, 2015). Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00005-7 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Table 5.1 Mineral concentrations of bone and stoichiometric hydroxyapatite (HA) and their lattice parameters as reported in literature (Dorozhkin and Epple, 2002). Types of ions

Concentration level of substituted ions in bone

Concentration level of substituted ions in stoichiometric HA

Calcium (Ca)

34.8e36.6 wt%

39.6 wt%

Phosphorus (P)

15.2e17.1 wt%

18.5 wt%

Carbonates (CO3)

4.8e7.4 wt%

e

Sodium (Na)

0.9e1.0 wt%

e

Magnesium (Mg)

0.6e0.7 wt%

e

Chlorine (Cl)

0.1e0.13 wt%

e

Fluorine (F)

0.03e0.10 wt%

e

Potassium (K)

0.03e0.07 wt%

e

Strontium (Sr)

0e0.05 wt%

e

Silicon (Si)

0e500 ppm

e

Zinc (Zn)

0e39 ppm

e

Chromium (Cr)

0e0.33 ppm

e

Cobalt (Co)

0e0.025 ppm

e

Manganese (Mn)

0e0.17 ppm

e

Lattice parameters

Bone

Stoichiometric HA

a-axis (nm)

0.9410

0.9430

c-axis (nm)

0.6890

0.6891

5.3

Synthetic hydroxyapatite

HA, a member of apatite family, has a general formula of M5(ZO4)3X, where M represents metallic ions, such as Ca2þ, Cd2þ, Sr2þ, Ba2þ, Pb2þ, Zn2þ, or Mg2þ; ZO3
4 is 3
4
2



represented by PO3
4 , CO3 , SiO4 , or SO4 ; and X represents OH , F , Cl , or CO3
3 . HA is very stable at pH range of 4.2e12.4. Unit cell of HA has a hexagonal structure with the space group P63/m. The unit cell formula of stoichiometric HA is described as Ca(I)4Ca(II)6(PO4)6(OH)2 as shown in Fig. 5.1. The anion channels of HA unit cell compose of six PO3
4 anions, where each P atom is coordinated to four oxygen atoms. They form equilateral triangles centered at the unit cell corners. The 2þ compact arrangement of PO3
4 groups in HA structure generates two sites for Ca 2þ ions; Ca(I) position and Ca(II) position. Out of 10, 4 Ca cations occupy Ca(I) position, where these positions are in columns parallel to the c-axis and coordinate to 9 2þ oxygen atoms of tetrahedral PO3
ions occupy Ca(II) posi4 . The remaining six Ca tion forming two equilateral triangles along the c-axis at z ¼ 1/4 and 3/4 and

Role of substitution in bioceramics

(a)

c axis

119

(b) Ca(II) O(III) OH H

a axis

Calcium Oxygen Hydrogen

P-tetrahedra

Figure 5.1 (a) A prospective view of a hydroxyapatite (HA) crystal unit cell. (b) The types of calcium ion sites found in the HA lattice (Ratnayake et al., 2017).
coordinate to six oxygen atoms of the PO3
4 and one monovalent anion, i.e., OH (El Feki et al., 1999). The volume of Ca(I) site is smaller than Ca(II) site (El Feki et al., 1999).The radii of substituted cations control their localization in these two positions (Nounah et al., 1992). The atomic structure of HA is able to substitute almost half of periodic table elements (Hughes and Rakovan, 2002). Comparatively, a limited substitution is possible for bioapatite as few types of ions are available in body fluids. There are many reports of HA substitution by a range of metals and nonmetals. Nonmetallic anions can substitute the phosphate group and metallic cation can substitute calcium ions of HA as shown in Fig. 5.2. The main shortcomings of orthopedic application of synthetic HA are its comparatively weak osteogenic potential compared to autografts and lack of angiogenic potential. Incident of infection after implantation of biomaterials is also a major challenge frequently faced by clinicians. HA substitution is a way to tune its physicochemical properties such as degree of crystallinity, grain size, mechanical strength, surface charge, porosity, solubility, and degradation kinetics. All these factors control biological response and healing capacity of HA. There are a growing number of in vitro as well as in vivo results mentioning better bioactivity and rate of healing with substituted HA materials. It is also observed that type and degree of substituted element and physiochemical properties of HA control its delivery applications of drug, gene, and protein (Lee et al., 2014).

5.4

Effect of substitution on charge and size of hydroxyapatite crystals

The structure of HA can accommodate a range of ionic substitution. Structural limits of HA unit cell control the type and degree of substitutions. Substitutions of HA can

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Handbook of Ionic Substituted Hydroxyapatites

(a) Ca

Zn

Ca1

(b)

OH ion

H

Halide ion

0 Ca2

Zn

P

Calcium ion

c

a

b

Figure 5.2 Types of substitutions in hydroxyapatite (HA) structure. (a) Cationic substitution, when calcium ion in HA is partially replaced with ions such as Mg2þ, Zn2þ, or Agþ. (b) Anionic substitution type A, a smaller hydroxyl ion is replaced by a large halide ion (Ratnayake et al., 2017).

affect the degree of crystallinity, solubility, grain size, and thermal stability and convert HA into other phases of calcium phosphate (CaP). Replacement of Ca2þ ion with cations having different charge and ionic size can unbalance the charge and lattice parameters. Bivalent cations substitute the Ca2þ ions without disturbing the charge, but substitutions of monovalent/trivalent cations lead to charge imbalance that can be overcome by creating supplementary vacancies or by substitution of other cations and anions. Substitution of HA with cations of larger diameter can expand the crystal cell volume. However, this phenomenon does not happen when larger monovalent cations substitute Caþ2 as it creates supplementary vacancies that reduce the channel diameter and consequently reduce the axis parameter (El Feki et al., 2000a). The substitution of Ca2þ (ionic size of 0.099 nm) by a smaller bivalent cation (Mg2þ, 0.069 nm) or a larger bivalent cation (Sr2þ, 0.113 nm) results in contraction and expansion of HA lattice parameters, respectively (Bigi et al., 2007). Usually, cations larger than Ca2þ prefer to occupy site (II). However, other factors such as charge of anions and bond strength also influence the selection of position, i.e., the Mn2þ ion (0.080 nm) prefers the site Ca(I) in fluorapatite, whereas it substitutes both Ca(I) and
Ca(II) sites simultaneously in cholorapatite (Supov a, 2015). The cationeoxygen interaction also influences the position of substituted cations. Ca(I) site has longer metaleO distance, allowing larger cations to occupy this site as reported in Sr2þ-substituted HA (Bigi et al., 2007). However, large concentration of substituted Sr2þ ions increases repulsion between ions expanding the c-axis. This stress is minimized by substitution of Sr2þ at Ca(II) site. Cations of smaller size than Ca2þ, Mg2þ, Zn2þ, In3þ, and Y3þ decrease HA crystal volume, while cations of larger size than Bi3þ and La3þ increase

Role of substitution in bioceramics

121

crystal volume (Webster et al., 2004). Mgþ2- and Znþ2-doped HAs have bigger grain sizes and broader size distribution. All trivalent cations used in this study (In3þ, Y3þ, Bi3þ, La3þ) decrease average grain sizes of HA (Webster et al., 2004).

5.5

Types of metallic substituents

5.5.1

Monovalent cationic substituents

Many monovalent metallic cations such as silver (Agþ), sodium (Naþ), potassium (Kþ), and lithium (Liþ) can substitute bivalent Ca2þ cation in HA crystal. They unbalance the charge and create defected sites in HA lattice (Matsunaga and Murata, 2009).

5.5.1.1

Sodium substitution

Sodium (Na) is an abundant trace element in hard tissue after calcium and phosphorus. Naþ ion substitution is very frequent in bioapatite; it occurs at ca 0.9e1.00 wt% level. Naþ in biological apatite plays a key role in cell adhesions and bone metabolism (Yokota et al., 2017b). Mostly, the literature reports the synthesis of sodiumsubstituted carbonated hydroxyapatite (C-HA) rather than sodium-substituted HA (Na-HA) (Sang Cho et al., 2014). The reason might be to overcome charge imbalance or improve bioactivity with carbonate group (Yokota et al., 2017b). The smaller size of sodium ion than that of Ca2þ ion does not make any significant modification in the hexagonal system of HA. Naþ preferring the Ca(II) site in HA lattice slightly increases the c-axis and cell volume (El Feki et al., 1999). The charge vacancies generated by Naþ substitutions are compensated by cosubstitution of other cations and anions (El Feki et al., 2000b,c; Wilson et al., 2004). Substitution of Naþ into C-HA decreases its crystallinity (Sang Cho et al., 2014) and improves thermal stability (Kannan et al., 2008). In vitro regeneration of calvarial bone defects with Na-HA shows better osteoconductivity compared with HA (Sang Cho et al., 2014). Biphasic mixtures of Na-HA and b-tricalcium phosphate (b-TCP) show better biological behavior than individual components.

5.5.1.2

Potassium substitution

Potassium (K) in bone matrix influences biomineralization, cell adhesion, and biomechanical processes. Dentin hypersensitivity (DH) is a very common oral problem in adults. Potassium salts such as potassium nitrate containing oral care products such as toothpastes, mouth rinses, and oral gels are known to reduce DH (Frechoso et al., 2003; Hodosh et al., 2007). Potassium chloridee and fluorohydroxyapatite-based dentin varnishes show good results to address hypersensitivity and show good biocompatibility with gingival and pulpal fibroblasts cells (Lochaiwatana et al., 2015). There are few reports about the Kþ substitution in HA and its biocompatibility (Ratnayake et al., 2017). K-HA synthesis is reported by soaking of HA in potassium salts solutions such as potassium carbonate and potassium chloride (Nordstr€om and Karlsson, 1992).

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Handbook of Ionic Substituted Hydroxyapatites

The concentration of potassium in HA controls adsorption of bovine serum albumin (BSA) (Weissmueller et al., 2014). Substitution of Kþ into the HA improves thermal stability (Kannan et al., 2007). Bigger ionic radius (0.133 nm) of Kþ should prefer to occupy Ca(II) site. However, the experimental results have shown the reverse trend, 83% of the Kþ ions are at Ca(I) site and 17% at site (II) (El Feki et al., 2000a). Kþ substitution in HA contracts a-axis, makes irregular changes in c-axis, and forms channel vacancies (Cacciotti, 2014; Kannan et al., 2007). In another report of Kþ substitution of HA, it has occupied Ca(I) site and increased both a- and c-axes parameters of HA lattice (Yokota et al., 2017a).

5.5.1.3

Silver substitution

Metallic prostheses are coated with bone regenerative materials to facilitate their rapid fixations with host bone by formation of uniform bone growth at interface. Infection is the most frequent reason of implant failure. Once implant is infected, the pathogens grow rapidly and form a biofilm. It shows resistance to host defense system as well as antibiotics (Costerton et al., 1999; Campoccia et al., 2006). In most cases, the only solution is implant replacement surgery. This causes economical and physical burdens on patients and hospital resources (Arciola et al., 2012). As the infections are the main cause of implant failure (Belt et al., 2001), antibacterial metal ions should be used to install antibacterial potential in implants materials. Silver (Ag) has been known to treat infections since ancient times. It is also biocompatible, nontoxic, and thermally stable at optimal dose level for human cells. It shows antifungal and antibacterial potential. It acts as nontoxic material at 35 ppb concentration and prevents microbes to develop. Agþ ions can substitute Caþ2 at Ca(I) site, and their bigger ionic size changes the lattice parameters of HA ions (Santos et al., 2015). The substituted ions decrease thermal stability and enhance solubility of HA. Ag-HA, with Agþ content more than 4%, shows thermal decomposition at low temperature. Cytotoxicity and antimicrobial properties of Agþ-containing materials are optimized by its concentration and synthetic parameters. The Ag-doped HA samples and their coating on implant materials display strong inhibition to bacterial growth via preventing bacteria adhesion, which ultimately prevents biofilm formation. Ag substitution in HA shows good in vitro biocompatibility as well as excellent antibacterial activity against Escherichia coli, Staphylococcus aureus, and Candida albicans (Riaz et al., 2018; Mo et al., 2007; Fu et al., 2016; Rameshbabu et al., 2007; Yan et al., 2015). Antimicrobial properties of Agþ-containing HA also depend on surface area and degree of crystallinity. Fine crystals of Ag-HA show fast release of Agþ ions. Ag-HA coatings on Ti-6Al-4 have shown good antibacterial properties against Gram-negative bacteria compared with Gram-positive bacteria (Chung et al., 2006). In another study, Ag-HAecoated titanium surfaces allow less number of bacteria to attach with implant surface as compared with HA-coated surface (Chen et al., 2006). The 0.2 wt% of Agþ-doped HA inhibits growth of Candida krusei and Klebsiella pneumoniae. A higher concentration of Ag (0.5 wt%) in HA samples has inhibited growth of Bacillus subtilis and E. coli (Costescu et al., 2013). Colonization and adhesion of Pseudomonas aeruginosa

Role of substitution in bioceramics

123

decrease in Ag-HA coatings by increasing Agþ concentration (Roy et al., 2012) A comparative study of antimicrobial activity of Agþ-, Cu2þ-, and Zn2þ-substituted HA against E. coli has shown the effectiveness of Ag-HA (Kim et al., 1998). Its optimal amount should be incorporated into HA to manage toxic effect. Its concentration range of 1e3 mg/mL has displayed toxic effects (Peetsch et al., 2013). High concentration of Ag in HA leads to decreased cellular and alkaline phosphatase (ALP) activities. It is, therefore, a matter of primary concern to attain balance between the positive effects and adverse consequences of Ag ions (Fielding et al., 2012).

5.5.1.4

Lithium substitution

Trace amount of lithium (Li) in the human body stimulates cell proliferation and can control mineral density of bone (Zamani et al., 2009). It is also well known for treatment of depression (Post, 2018). Lithium may influence parathyroid hormone level affecting calcium homeostasis and is used to treat osteoporosis (Cohen et al., 1998). It shows osteogenic effects when administered in low bone mass animal models (Wang et al., 2016b). Despite its many benefits, there are also few reports that deal with its toxicity and management (Gitlin, 2016; Ott et al., 2016). Therefore, its optimal concentration should be used in biomaterials to get maximum benefits. Lithiumsubstituted HAs (Li-HA) are the promising biomaterials (Popescu et al., 2018). Li substitution improves CaP thermal stability compared with its substitution with other ions (Si, Sr, and Mg) (Matsumoto et al., 2009). Li-HA significantly improves bone healing as it enhances the osteoblast activity (Wang et al., 2016b). Li-substituted HA prepared by sintering of HA with lithium carbonate shows good osteoblast proliferations (Shainberg et al., 2012). Li substitution increases density of the HA after sintering (Fanovich et al., 1999). Lithium-doped HA coatings onto medical-grade titanium substrates are noncytotoxic on MG-63 osteosarcomaederived cells, dermal fibroblasts, and immortalized keratinocytes. Moreover, morphologies of Li-HA coating on implant have helped to inhibit growth of S. aureus and C. albicans (Duta et al., 2019).

5.5.2

Bivalent cationic substituents

Bivalent metallic cations such as magnesium (Mg2þ), strontium (Sr2þ), zinc (Zn2þ), manganese (Mn2þ), copper (Cu2þ), and cobalt (Co2þ) can substitute bivalent Ca2þ cation without disturbing the charge balance.

5.5.2.1

Magnesium substitution

Magnesium (Mg), an abundant element in human body, is required to maintain many physiological and vascular functions (Cooke and Losordo, 2002; Maier et al., 2004). It is needed for synthesis of nucleic acid, proteins, lipids, and in enzymatic reactions of energy generation (O’Neill et al., 2018). Approximately, 65% of total body’s magnesium is found in hard tissues. Its deficiency adversely effects every stage of skeletal metabolism and bone development by effecting both osteoblastic and osteoclastic cellular activities (Percival, 1999). Its prolonged deficiency can lead to osteoporosis

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Handbook of Ionic Substituted Hydroxyapatites

(Stendig-Lindberg et al., 2004; Castiglioni et al., 2013). Mg doping can install osteogenic potential in biomaterials (Zou et al., 2011; Yu et al., 2017; Rude et al., 2005). Mg-doped CaP materials improve cell attachments and proliferation, ALP production, and bone densification (Xue et al., 2008). Angiogenesis is a key requirement of tissue regenerations process (Chim et al., 2013). Lack of angiogenesis leads to restorative failure (Malhotra and Habibovic, 2016). Particularly, in vivo failure of regenerative biomaterial is largely due to the lack of angiogenesis (Amini et al., 2012). Bone repairing magnesium calcium phosphate cement promotes osteogenesis and angiogenesis as shown by both in vitro and in vivo experimental results (Wang et al., 2016a). Mg2þ substitution in HA crystal can reduce the lattice parameters as ionic radius of Mg2þ is smaller than ionic radius of Ca2þ. However, the charge balance in HA is not disturbed by substituting Ca2þ ions. Its substitution in HA decreases its crystallinity and can destabilize the apatite structure and facilitate thermal conversion of HA to b-TCP (Bertoni et al., 1998; Laurencin et al., 2011). HA substituted with 5.7 mol% Mg2þ ions has more resemblance to biological apatite in terms of morphology and crystallinity. In vivo experiment of Mg-HA reveals that there is no induced cytotoxicity, genotoxicity, and carcinogenicity (Landi et al., 2008a). Mg substitution also decreases HA density, compressive strength, and microhardness (Cacciotti et al., 2009). Fracture toughness of HA containing Mg addition up to 1 wt% is approximately equal to that of bone (Zyman et al., 2006).

5.5.2.2

Strontium substitution

Calcium and strontium (Sr) belong to group 2A elements. Sr content is 0.035% of the Ca in body, and its higher concentration is observed in regions of higher metabolic turnover (Dahl et al., 2001). Sr is a nontoxic key element for bone health. It is used for osteoporosis treatment. Its prolonged use does not show toxic outcomes (Meunier et al., 2009). Strontium ranelate, a drug prescribed to treat osteoporosis, works via dual mode of action to reduce bone resorption, stimulates osteoblast differentiation, and inhibits osteoclast activity. Osteoporotic patients on this drug have lesser incidence of fractures (Bonnelye et al., 2008; Meunier et al., 2004). Sr addition also decreases the toxic effects of other metallic ions, i.e., Ag-HA coatings on metallic implant completely decline ALP activities; however, addition of SrO into it effectively offset toxic effect of Ag (Fielding et al., 2012). Excellent regenerative property of Sr encourages its doping into bone grafts. A commercial product called Stronbone by RepRegen Ltd., UK, is a strontium-containing biomaterial having good bone regenerative capacity (Tian et al., 2009). Addition of Sr into biomaterials enhances bioactivity of the material as shown by several in vitro and in vivo studies (Cardemil et al., 2013; Boanini et al., 2018; Tian et al., 2009; Hulsart-Billstr€ om et al., 2013). Sr-doped calcium polyphosphate scaffolds promote osteogenic and angiogenic potential (Chen et al., 2008). Sr-doped polyphosphate has improved in vivo bone growth in 16 weeks compared with undoped analogue (Tian et al., 2009). The concentration of 3e7 wt% is required for better osteoblast activity and osteogenic differentiation (Capuccini et al., 2009b). Calcium phosphate cement (CPC) containing Sr-HA exhibited twice higher compressive strength (2.92 MPa) compared with pure CPC (Shen et al., 2012). Substitution of

Role of substitution in bioceramics

125

Sr2þ in HA can change its lattice parameters. Both a- and the c-axes are increased progressively by increasing its amount into HA (Bigi et al., 2007; Kavitha et al., 2014; Capuccini et al., 2009a). The structural refinement studies indicate that the Sr2þ ions occupy both Ca(I) and Ca (II) sites of the HA structure (Bigi et al., 2007). Its bigger ionic radius (0.13 nm) prefers Ca(I) site when its lower amount is to be substituted. However, the selectivity is reversed by increasing concentration of Sr2þ substitution (Terra et al., 2009; Li et al., 2007; Renaudin et al., 2009) Higher amount of Sr2þ in HA destabilizes its structure and increases solubility (Pan et al., 2009). Sr2þ substitutions of 0.3 and 1.5 mol% produce a single-phase HA samples (Li et al., 2007). However, higher concentration of Sr (15 mol%) decreases the crystallinity and facili2
2þ tates incorporation of HPO2
4 and CO3 ions (Landi et al., 2008b). Sr -substituted HA has improved proliferation, differentiation, and angiogenic expressions of human osteoblast MG63 (Lin et al., 2013). Revision hip hemiarthroplasty of goat model by using PMMA bone cement containing Sr-HA showed significantly higher bonding strength of material with bone (3.36  1.84 MPa) compared with PMMA bone cement (1.23  0.73 MPa) (Ni et al., 2006).

5.5.2.3

Zinc substitution

All biological tissues require zinc (Zn) as an important vital trace element for normal growth. It takes part in different cellular, immunological, and neurological functions (Prasad, 2013). It also possesses antioxidant and antiinflammatory properties; therefore, it can take part in curing of several chronic diseases (Prasad, 2014). Bone contains major amount of the total Zn present in the body (Barrea et al., 2003). It promotes osteoblasts activities and inhibits osteoclasts activities (Moonga and Dempster, 1995). It is also present in many important metalloenzymes to manage their structure, catalytic, or regulatory actions. It is a part of ALP enzyme that is required for bone regeneration. ALP contains four Zn ions and two Mg ions, so deficiency of Zn decreases bone density (Cho et al., 2007; Coleman, 1992). In dental application, zinc phosphate cements are frequently used since long time. Substitution of Zn2þ into the HA lattice is well investigated. By increasing its content in HA, a-lattice parameter is increased (Xiao et al., 2008; Kumar et al., 2012; Ren et al., 2009) and c-lattice parameter is decreased(Miyaji et al., 2005). In another report, Zn2þ substitution decreases lengths of both a- and c-axes (Thian et al., 2013). In bioapatite, Ca(II)
site facilitates substitution of Zn without disturbing the HA crystal structure (Supov a, 2015). Similarly, Zn is substituted at Ca(II) site in synthetic HA (Ma and Ellis, 2008). Its higher substitution in HA decreases crystallinity and facilitates the formation of other CaP phases (Bigi et al., 1995; Wang et al., 2010). Heat treatment of Zn-HA also generated other phases of CaP, i.e., TCP (Sogo et al., 2004). The reported value about the maximum capacity of HA to substitute Zn2þ has contradictions (Ergun et al., 2002; Xiao et al., 2008). Zn2þ-substituted HA was stable when substitution limit of Zn2þ was around 9.4 wt%, whereas further Zn substitution formed parascholzite phase (Miyaji et al., 2005). Other studies have reported 15 and 25 mol% as substitution limit of Zn2þ into HA (Bigi et al., 1995; Miyaji et al., 2005). In vitro results showed enhanced bioactivity and antimicrobial activity of Zn-HA (Thian et al., 2013; Stanic

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Handbook of Ionic Substituted Hydroxyapatites

et al., 2010). Zn-HA improves bone regeneration via increasing osteogenic cells proliferation and differentiation and decreasing the activity of bone resorption cells (Wang et al., 2010; Chou et al., 2016; Hall et al., 1999; Seo et al., 2010).

5.5.2.4

Manganese substitution

Manganese (Mn) is a trace element, participating in many metabolic pathways. It plays significant role in the formation of cartilage and bone tissues. It is an essential cofactor in different enzymes that activate mitochondria and participate in synthesis of mucopolysaccharides and glycoprotein. Therefore, it helps to protect cell from bacterial and viral infections. Its deficiency retards endochondral osteogenesis (Medvecký et al., 2006). In a study, its deficiency in young chicks deformed the bones and decreased level of important bone markers in serums (Zhaojun et al., 2013). Mn2þ substitution into HA does not disturb biocompatibility of HA (Li et al., 2012b; Gy€orgy et al., 2004). Ti surface coated with thin film of Mn-C-HA enhances osteoblast cells proliferations compared with uncoated surface (Gy€ orgy et al., 2004). In other experiments, Mn-containing CaP coating on Ti substrate improves osteocalcin production compared with Sr-containing CaP coating (Bracci et al., 2009). Mn2þ substitution into HA facilitates its thermal conversion to b-TCP (Mayer et al., 2006) (Medvecký et al., 2006). Its amount in HA crystal lattice affects its morphology; relatively higher amount (1.23 wt%) forms smaller needle-like morphology, and relatively lower content (0.73%) constructs platelet crystals of micron size (Mayer et al., 2008). In another study, Mn2þ-substituted HA contains low degree of crystallinity. Its contents from 0% to 4% have remarkable effect on particle size and morphology, i.e., lower concentration results in longer rod-shaped particles of w110 nm and higher amount produces smaller semispherical-shaped particles of w75 nm (Anwar and Akbar, 2018).

5.5.2.5

Copper substitution

Copper (Cu) is required for the function of several proteins. It can stimulate endothelial cells (ECs) for the formation of new blood vessels (Lakhkar et al., 2013). Scaffold of dicalcium phosphate dihydrate loaded with low concentration of cupric Cu2þ ions (22 ng) improves vascularization compared with scaffolds without Cu2þ ions (Barralet et al., 2009). CPC impregnated with 0.56 mg Cu2þ/cm2 showed enhanced osteoblast proliferation (Ewald et al., 2012). The extent of Cu2þ substitution in HA crystals is reported to a maximum of 15 wt% (Othmani et al., 2018). Composite coatings of Ag-HA, Ni-HA, Zn-HA, and Cu-HA were analyzed against S. aureus, and the best antibacterial results were shown by Cu-HA (Sanpo and Tharajak, 2017). Hydrothermally prepared Cu-HA samples possess micro-/nanostructural morphologies. These structural morphologies influenced EC proliferation (in vitro) and blood vessel formation (in vivo) (Elrayah et al., 2018).

5.5.2.6

Cobalt substitution

Cobalt (Co) is a vital element and fundamental part of vitamin B12, which plays an important role in production of red blood cells. Its presence in tissue repairing

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materials can stimulate angiogenesis. Periosteum of bone marrow stromal cells, when impregnated with 100 mM CoCl2, showed upregulation of vascular endothelial growth factor (VEGF) that controls angiogenesis (Fan et al., 2010). CoCl2-impregnated collagen scaffold exhibited better ossification and higher degree of vascularization (Bose et al., 2013). In HA, cosubstitution also introduced good proangiogenic property without suppressing its osteogenic property (Kulanthaivel et al., 2016). In an animal study, 5e12 wt% Co2þ-substituted HA samples were investigated, and the sample containing its highest concentration showed better bone regeneration in terms of improved osteogenesis and angiogenesis (Ignjatovic et al., 2013). Another study also reported improved osteogenic potential with Co2þ-substituted HA compared with HA (Ignjatovic et al., 2015). Its substitution into HA also installed paramagnetic properties (Ignjatovic et al., 2013) and extended its applications for target drug delivery, magnetic imaging, and cancer hyperthermia treatments. They show very slow release from HA lattice; therefore, Co-HA and HA have shown similar degradation rate (Kramer et al., 2014).

5.5.3

Trivalent cationic substituents

HA can be doped with trivalent elements of p-block, d-block, and with f-block elements, i.e., aluminum (Al), gallium (Ga), indium (In), bismuth (Bi), iron (Fe), yttrium (Y) lanthanum (La), etc. During the substitution of HA with trivalent cations, the charge imbalance can be compensated by generation of vacant cation sites. Two Mþ3 ions can replace three Ca2þ ions to maintain charge balance.

5.5.3.1

Aluminum substitution

Aluminum (Al) is a trace element in human hard tissues. It plays different roles in the process of bone metabolism. Although many reports showed toxic effect of Al on bone health such as induced aberrant bone metabolism (osteomalacia and osteoporosis) (Cannata et al., 1983; Chappard et al., 2004, 2015; Li et al., 2012a), however, its lower dose level induces osteogenesis (G omez-Alonso et al., 1999). Al can inhibit dental caries (Putt and Kleber, 1985). Al-doped HA with its different concentrations (0.5e2.5 mol%) showed in vitro biocompatibility with L929 cell lines (Kolekar et al., 2016). A comparative study of trivalent metal ions such as Al3þ-, La3þ-, and Fe3þ-substituted HA samples shows that hydroxyl ions (AleOH, FeeOH) are present on the surface of Al-HA and Fe-HA. These ions enhance their hydrophilicity and BSA absorption capacity (Kandori et al., 2010).

5.5.3.2

Iron substitution

Iron (Fe) is a vital element for many cellular and enzymatic functions; it also facilitates oxygen and electron transport and participates in processes of DNA synthesis. Iron deficiency, anemia, is a widespread health issue and badly affects the body metabolic functions and bone health (Wright et al., 2013). Fe activates lysyl hydroxylase and 25hydroxycholecalciferol hydroxylase to stimulate bone mineralization and synthesis

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(Zofkova et al., 2017). In vivo study revealed poorly mineralized skeleton due to its deficiency (Parelman et al., 2006). Owing to these important roles of iron, there is an increased attention to incorporate it into biomaterials. Fe-HA samples are biocompatible, and MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) assay shows good proliferation of MG63 cells with them (Sarath Chandra et al., 2012). Smaller ionic size of Fe2þ (0.0835 nm) than Ca2þ changes lattice parameters of HA. Fe3þ substitution into HA decreases its crystallinity, inhibits its growth in direction of c-axis, and promotes its growth in a-axis (Zuo et al., 2012). HA is a diamagnetic material (Sarath Chandra et al., 2015). Fe3þ substitution can introduce magnetic properties in HA materials and (Kramer et al., 2013; Yang et al., 2018) extend its applications to cancer curing by using hyperthermia therapy and target delivery system via external magnetic field (Hou et al., 2009). In vivo experiments showed dramatic reductions of tumor size in 15 days when Fe-HA implanted area was explored to magnetic field (Morrissey et al., 2005). The 10% Fe substitution into HA resulted in phase-pure HA. Fe-HA samples showed improved bioactivity, biocompatibility, and good antibacterial property. Moreover, Fe substitution showed the sustained release of methotrexate antibacterial drug (Sheikh et al., 2018). Comparative biocompatible study of HA, Mn-HA, and Fe-HA showed that all substituted HAs were biocompatible to osteoblast cells, but twice the number of cells were present on Fe-HA sample as compared with HA and Mn-HA samples (Li et al., 2012b).

5.5.3.3

Other trivalent substitutions

Gallium (Ga) ions are biocompatible and are used clinically to inhibit bone resorption. It is a suitable candidate for osteoporosis and hypercalcemia cancer treatment (Melnikov et al., 2008, 2009; Bernstein, 1998). Ga-containing CPC improved bone remodeling process in rabbit critical bone defect model (Mellier et al., 2015). Gallium-doped amorphous calcium phosphate showed antibacterial activity against P. aeruginosa (Yang et al., 2017). Ga does not replace Ca2þ in HA lattice, but intercalation (up to 11.0 wt%) of gallium nitrate and sodium gallate has been observed in HA lattice without creating any distortions (Melnikov et al., 2009). Indium (In), a trace body element (Andersson et al., 2017), is substituted into HA, and it has enhanced adhesion and differentiation rates of osteoblasts (Webster et al., 2004). Bismuth (Bi) is a p-block element, and it is not found in human body. Bi-HA has shown good osteoconductivity, biocompatibility with osteoblasts (Selvakumar et al., 2016; Webster et al., 2004; Ciobanu et al., 2015), and antibacterial effect against S. aureus and E. coli (Selvakumar et al., 2016; Webster et al., 2004). Yttrium (Y) is a radioactive therapeutic metal and belongs to d-block. Y-HA improves electrical conductivity, hydrophilicity (Owada et al., 1989), and biocompatibility when cultured with human osteoblast cells (Sato et al., 2006). It also improves the mechanical properties (Nathanael et al., 2011). Y3þ substitution into HA results in shrinkage of lattice structure by decreasing lattice parameters of aand c-axes (Nathanael et al., 2011).Comparative biocompatibility study of phasepureesubstituted HA samples with various divalent (Mg2þ and Zn2þ) and trivalent (La3þ, Y3þ, In3þ, and Bi3þ) metallic cations was carried out. Among them, Zn2þ-,

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In3þ-, and Bi3þ-substituted HA samples were the most effective substituents in terms of better osteoblast adhesion and differentiation (Webster et al., 2004). In another report, phase-pure HA substitutions with di- and trivalent cations such as magnesium, zinc, cadmium, and yttrium were also synthesized. All substituted HA samples showed smaller grain size and decreased a- and c-axes because of smaller size of substituted cations (Mg2þ, Zn2þ, and Y3þ) than calcium ions. This study showed that yttrium could be substituted up to 7%, while 2% substitution of other cations was noted (Webster et al., 2002). In vitro osteoblast adhesion test showed that the best adhesion was showed by Y-HA compared with HA, Cd-HA, zinc (Zn-HA), and magnesium (Mg-HA). The adhesion of Y-HA was increased when concentration of Y cations (2e7 mol%) was increased. Better cell adhesion of Y-HA supported its higher absorbing ability of calcium and proteins (vitronectin and collagen) (Webster et al., 2002).

5.5.3.4

f-block cation substitutions

Many trivalent metal cations belonging to f-block have the capacity to replace Ca2þ ion of HA. These include lanthanum (La3þ), cerium (Ce3þ), praseodymium (Pr3þ), neodymium (Nd3þ), samarium (Sm3þ), europium (Eu3þ), gadolinium (Gd3þ), terbium (Tb3þ), holmium (Ho3þ), erbium (Er3þ), thulium (Tm3þ), ytterbium (Yb3þ), and uranium (U3þ) (Tite et al., 2018). Lanthanum (La)-substituted HA (La-HA) showed good biocompatibility (Lou et al., 2015; Ahymah Joshy et al., 2011; Jadalannagari et al., 2014). La-HA coating on Ti surface showed good bonding strength (Lou et al., 2015). La-doped HA can be used to prevent dental caries as La doping into HA inhibited the dissolution rate of HA in phosphate-buffered saline (PBS). La-doped HA samples pose higher specific surface area and hardness compared with undoped analogues (Ahymah Joshy et al., 2011). Europium (Eu) and cesium (Ce) are present in the human body and show antibacterial activity (Sundarabharathi et al., 2019; Iconaru et al., 2013) and bone tissue regeneration capacity (Victor et al., 2017; Andronescu et al., 2015; Priyadarshini et al., 2017). Their ionic radii are similar to Ca2þ, so they can easily substitute Ca2þ in HA lattice. f- Block cation doping also makes HA a potential candidate for magnetic resonance imaging. Luminescent and magnetic properties in HA have been installed by codoping of europium (Eu3þ) and dysprosium (Dy3þ). Eu-Dy-doped samples showed good biocompatibility (Tesch et al., 2017).

5.6

Nonmetallic substitutions

Bioapatite performs different functions in bone, dentin, and enamel. Hydroxyl groups present on the surface of apatite tailor its solubility and dynamic equilibrium with collagen. These groups can be substituted by nonmetallic anions (F
, Cl
, CO2
3 ). 2
3
2
2
2
2
On the other hand, oxyanions (SiO4
4 , CO3 , BO3 , HPO4 , SO4 , SeO3 , SeO4 )
3 can substitute PO4 of the HA. This nonmetallic substitution is also called as anionic substitution.

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Handbook of Ionic Substituted Hydroxyapatites

Fluoride substitution

Fluoride anion (F
) can substitute hydroxyl group of bioapatite. Smaller size of F
is best fitted in HA lattice in place of OH
and results in a compact unit cell structure by decreasing length of a-axis. These ions improve apatite’s chemical stability, lower the rate of decomposition (Kim et al., 2009), and increase resistance to corrosion in the body fluids compared with HA (Rameshbabu et al., 2006). F-HA can be used to treat osteoporosis (Pak et al., 1989). The solubility and degradations of HA can be tuned by amount of substituted fluoride (Eslami et al., 2009). F-HA coating has improved bond strength with metallic implants (Lugscheider et al., 1991). Heat treatment of F-HA coating on titanium surface has improved the bonding strength and introduced hydrogen bonding, i.e., OHeFeOH. At higher concentration of fluoride in coated F-HA, OHeF hydrogen bonding has been observed (Tredwin et al., 2013; Wu et al., 2010). F-HA is less soluble in acidic media compared with HA, which makes it a suitable candidate for dental applications. Dental paste containing F-HA has significantly improved the hardness of the tooth enamel and prevented its damage (Wang et al., 2012a). F
substitution into HA has enhanced its biocompatibility and cellular attachments (Wang et al., 2007; Peng et al., 2019; Eslami et al., 2009). By increasing the concentration of F
substitution in HA, thermal stability of apatite has been increased (Eslami et al., 2009; Bianco et al., 2010). F
substitution has also improved mechanical properties of HA (Bianco et al., 2010; Gross and Bhadang, 2004).

5.6.2

Chloride substitution

The chlorapatite (Cl-HA) is 6.8 wt% of natural apatite. Larger ionic size of Cl
(0.167 nm) restricts its frequent substitution in bioapatite. Cl
substitutions in HA can expand the unit cell volume (Kannan et al., 2006b) and lattice parameter such as a-axis (Zhao et al., 2014) and c-axis of HA (Kannan et al., 2006b). Cl
ions generate acidic environment that stimulates osteoclasts activities and bone resorption process. Therefore, pure chlorapatite is not considered a suitable candidate for hard tissue regeneration.

5.6.3

Carbonate substitution

Bioapatite is also called carbonated apatite as carbonates (CO2
3 ) ions are the most abundant substitutions in it. Approximately, 5e8 wt% CO2
3 is present in the bone (Zapanta-LeGeros, 1965) and 3.5 wt% in tooth enamel (Gross and Berndt, 2002; Elliott, 2002; LeGeros, 1991). There are two possible positions in apatite that can
be substituted by CO2
3 . It can either substitute at the OH site called as “A-type” sub3
stitution or in the PO4 position called “B-type” substitution. Planar structure of CO2
3 2þ and weaker bond between CO2
than calcium phosphate bond introduce 3 and Ca many changes in apatite structure and properties, i.e., decrease a-axial length, increase c-axial length (Wopenka and Pasteris, 2005), improve mechanical strength (Merry et al., 1998), alter solubility of apatite, and install mesoporosity. C-HA has been reported as good stimulator and facilitator for bone growth (Mostafa et al., 2011).

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Mesoporosity of C-HA has shown better bioactivity, cell proliferation, osteoblast attachments, and osteointegration compared with nonporous C-HA (Mohammad et al., 2016). The C-HA implants have displayed good biocompatibility, osteointegration, osteoconduction, and fast bioresorption compared with HA implant showed by in vivo results (Landi et al., 2003).

5.6.4

Silicon substitution

Silicon (Si) is nonmetallic element found in bone and connective tissues (Carlisle, 1980; Jugdaohsingh, 2007). It is required for production and function of glycosaminoglycan. Therefore, it helps in bone formation and maintenance of cardiovascular health and wound healing (Nielsen, 2009). Si is an abundant element, and its deficiency is very rare in human body. Experimental results of animal models reveal that Si dietary deficiency has deformed bone growth and taken to low collagen formation (Carlisle, 1980). In biomineralization process, significant amount of Si (0.5 wt%) was present at the earliest stages of calcification (Carlisle, 1970). The orthosilicic acid (Si(OH)4) facilitates apatite mineralization process even in the presence of proteins that normally inhibit apatite formation (Tanizawa and Suzuki, 1995; Damen and Ten Cate, 1992; Pietak et al., 2007). Si accelerates osteoblast proliferation and osteogenic differentiation (Rodrigues et al., 2017) as well as possessing angiogenic potential. Calcium silicate material expresses VEFG expression in human dermal fibroblasts regulating the angiogenesis (Li and Chang, 2013). Calcium silicate graft in a rabbit’s defective femur shows the angiogenic potential of silicon (Wang et al., 2013).
3 Tetravalent silicate ions, SiO4
4 , can substitute PO4 group of HA and generate hydroxide vacancies. In the literature, a report of up to 5 wt% of Si substitution into HA was reported, and 1 wt% substituted amount was enough to improve bioactivity (Vallet-Regí and Arcos, 2005; Gasqueres et al., 2008). The carbonate groups in apatite facilitate incorporation of silicates into HA lattice during thermal treatment. Silicon amount of more than 1.0 mol in HA can form a-TCP during calcination at 700⁰C (Palard et al., 2008). According to Rietveld refinement method, SiO4
4 substitution of 0.4 wt% does not make major changes in crystal structure. The a-lattice parameter is contracted by w0.005%, while the c-parameter is expanded by w0.1% (Leventouri et al., 2003; Gibson et al., 1999). SiO4
4 substitution does not affect the interatomic distances; however, the angles of tetrahedrons are distorted from 109.47, and distortion in angle is increased by increasing the further SiO4
4 substitution. HA morphology and crystallinity are affected by SiO4
4 substitution (Leventouri et al., 2003). Si-HA promotes angiogenic potential (Magnaudeix et al., 2016) and has a stimulatory effect on osteoblast proliferation, differentiation, and mineralization process (Aminian et al., 2011; Izquierdo-Barba et al., 2019; Casarrubios et al., 2016; Balamurugan et al., 2008; Liu et al., 2017).

5.6.5

Boron substitution

Boron (B) is a beneficial element for human health, particularly for growth of bone and wound healing (Lakhkar et al., 2013). Boron-deficient diet markedly decreases bone

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formation capacity in mice (Gorustovich et al., 2008). In another study, its supplement improves bone mineral density and increase compressive and tensile strength of bone (Hakki et al., 2013). It stimulates the growth factor VEGF, which is vital for new bone formation and wound healing (Studer et al., 2012; Dzondo-Gadet et al., 2002). Its substitution into hard tissue regenerative materials such as CaP, HA, and bioglass has shown enhanced angiogenesis and osteogenesis capacity of the damaged tissue

3 (Yılmaz and Evis, 2016). Its ions (BO3
3 or BO2 ) can substitute PO4 group and/or
OH group in the HA lattice (Ternane et al., 2002b). Oxyboron ion substitutions can change lattice parameters by slightly decreasing a-axis and increasing c-axis (Barheine et al., 2011). When higher concentration of boron precursors (P/B ratio ¼ 7.22) was used for substitution, other phases of Ca3(BO3)2 and CaO were appeared along BHA (Ternane et al., 2002b). B-HA samples have shown better adhesion and proliferation rates with mesenchymal stem cells (Barheine et al., 2011).

5.6.6

Sulfate substitution

Sulfate anions are abundant in human plasma and are needed for cellular function. Sulfur performs important role in repairing of skin cells (Mol et al., 2009), nails (Forslind et al., 1976), hair, and cartilage. It protects cartilage from osteoarthritis (Alshemary et al., 2013). There are many reports mentioning potential applications of composite of HA and calcium sulfate for bone repair (Nilsson et al., 2004, 2013). A limited work is reported so far about the substitution of sulfate in HA. Needle-shaped, sulfate ionesubstituted HA particles were prepared by microwave-assisted ion exchange process (Alshemary et al., 2013). Sulfate substitution has resulted structural disorder into HA lattice as its ionic size (0.258 nm) is bigger than that of PO
3 4 ion (0.238 nm). Its substitution increases the lattice parameters and cell volume of HA (Alshemary et al., 2013).

5.6.7

Selenium substitution

2
Selenium in the forms of selenate (SeO2
4 ) or selenite (SeO3 ) oxyanions has been reported for chemopreventive properties (Wang et al., 2012b; Zeng and Combs, 2008; Monteil-Rivera et al., 1999). The selenate (0.249 nm) has a tetrahedral structure. Its size is a bit bigger than of phosphate ion (0.238 nm). On the other hand, size of selenite ion (0.239 nm) is like phosphate ion, but it has flat trigonal pyramid geometry. After substitution, it creates a positively charged vacancy that can compensate decalcification and dihydroxylation (Kolmas et al., 2014). Selenate- or selenite ionesubstituted HA samples were synthesized by coprecipitation method (Wang et al., 2012b; Kolmas et al., 2014; Ma et al., 2013) to treat and reduce chance of recurrence bone cancers (Wang et al., 2012b). Selenate substituted HA lattice has shown slight expansion in parameter a. However, the selenite substitution has resulted in significant expansion of a-axis and no change in c-axis parameters (Kolmas et al., 2014). Selenium-added HA coating has shown excellent proliferation of MC3T3-E1 preosteoblasts cells and no cytotoxicity when coating composite contains 0.6% of Se. The coating material also prevented the formation of biofilm (Rodríguez-Valencia et al., 2013).

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Selenium-added HA coating also showed antibacterial properties (Rodríguez-Valencia et al., 2013). Selenate-substituted HA showed nontoxic effect at low concentration in Spirotox and Microtox tests (Kolmas et al., 2014).

5.6.8

Tellurium substitution

Tellurium (Te) is a semimetal belonging to p-block elements. It possesses antioxidant and antibacterial properties (Zare et al., 2012; Zhong et al., 2013). There is one report so far about Te-doped HA synthesis. Its presence introduces antimicrobial and antifungal activities in the prepared HA samples. There is no reported data about biocompatibility of Te-doped HA (Yahia et al., 2017).

5.7

Significance of multiple substitutions

The preparation of a multifunctional apatite via multiple substitution of HA is an excellent strategy to mimic properties of bioapatite and expand its clinical applications. Multiionic substitution can improve HA bioactivity, osteogenesis, angiogenesis, crystallinity, solubility, structural properties, degradation kinetics, mechanical properties, magnetic properties, luminant properties, and antibacterial potential (Bose et al., 2013). Dual substitutions may be cationic or anionic or a combination of cation and anion in HA lattice. Bication-cosubstituted HA (Na/Mg-HA) sample and b-TCP mixtures improved thermal stability of crystalline phases of HA at 1400  C (Kannan et al., 2006a). Mg- and Zn-cosubstituted HA (Mg/Zn-HA) enhanced mechanical properties, i.e., compressive as well as surface hardness. A comparative antibacterial study of dual-substituted HA samples (Ti/Ag, Ti/Cu and Ti/Zn) has been conducted against E. coli and S. aureus. In these samples, Ti/Ag-HA and Ti/Cu-HA have exhibited better inhibition (Hu et al., 2007) Cosubstituted HA with Sr and Ce showed good antibacterial property as well as improved bioactivity. In vitro antibacterial evaluation of Mg/ Ag-cosubstituted HA has shown reduced cytotoxicity of Ag ions (Gopi et al., 2014). Cu- and F-cosubstituted HA showed antibacterial activity against E. coli (Shanmugam and Gopal, 2014). Dual doping of borane and cation, either Eu3þ or Ce3þ (Ternane et al., 2001, 2002a), installed luminescent properties of BHA. Lanthanide-doped, Eu- and Tb-codoped, spindle-shaped F-HA showed unique dual color emission potential for cell and tissue imaging applications (Ma et al., 2016). The growth of apatite has been improved by trications cosubstituted HA (Sr/Mg/Zn-HA) (Gopi et al., 2012).

5.8

Grafting of organic compounds/polymers on the surface of hydroxyapatite

The surface grafting of HA can be anticipated to tune its properties for targeted applications. This modification carried out with organic chelants may influence physiochemical properties of HA such as surface area, pore volume, pore size (Tabassum

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et al., 2016; Zarif et al., 2019a,b; Hussain et al., 2016), surface charge (Chen et al., 2011), hydrophobicity, or hydrophilicity (Lee et al., 2014). These properties play a significant role in biological functions, cell adhesions, binding strength with other surfaces, as well as delivery applications of drug, protein, and gene, etc. (Zarif et al., 2019a,b; Tabassum et al., 2016; Sharif et al., 2019; Chen et al., 2011; Lee et al., 2014; Hussain et al., 2016). The organic chelants used for tuning of HA surface include carboxylates (Tabassum et al., 2016; Zarif et al., 2019a,b), amino acids (Gonzalez-McQuire et al., 2004), alcohol (Borum-Nicholas and Wilson, 2003), amines (Zhang and Darvell, 2010), alkylphosphonic (D’Andre and Fadeev, 2003), and organosilanes (Cao et al., 2014; Wang et al., 2011). The nontoxic carboxylic acids and amino acids are regarded as superior modifiers when compared with silane coupling agents. The later have been known to exhibit some biotoxic effects (Dupraz et al., 1996). In our group, a series of carboxylic acid (adipic acid, malonic acid, succinic acid, and stearic acid)emodified HA particles has been synthesized to tailor the physiochemical properties of HA. In this study, ibuprofen was used as a model drug to determine the effect of functionalities and surface properties of modified HA. In vitro drug loading and delivery results have revealed the dependence on functional groups, surface area, and porosity of the modified material (Tabassum et al., 2016). In other studies, carried out in our group, HA was modified with citric acid, aspartic acid, and tartaric acid. It was shown that physiochemical properties of modified HA enhanced its electrostatic interaction with an antibiotic drug, moxifloxacin. The grafted HA showed better results for drug release (in vitro) as compared with pure HA (Zarif et al., 2019a,b). Grafting of HA with polymeric materials helps to improve HA adhesion to metallic implant. A series of polyvinyl alcohol decorated mesoporous HA (PVA-HA) composites were prepared via in situ coprecipitation method. Concentration of PVA in PVA-HA composite has impacted the coating adhesion of material on stainless steel plates (Hussain et al., 2016).

5.9

Conclusion and outlook

This chapter has provided an overview of how different substituents can be used to improve performance of HA. Furthermore, this study provides a background knowledge to design the new substituted versions of HA. Substitution of HA is a powerful tool to tune properties of HA. Structural parameters of HA allow single- as well as multiatomic substitution into apatite, and this freedom encourages the researchers to use HA for diverse biomedical applications. Biogenic apatite is usually integrated with living organs; however, inhomogeneities and impurities are main challenges discouraging its frequent use. In contrast, synthetic multiions-substituted HA can be prepared easily with predictable and reproducible properties by using controlled conditions. In addition, synthetic HA can also be substituted with ions that are not present in natural apatite and can improve functionalities of HA, i.e., antibacterial, angiogenic, osteogenic, magnetic, luminescent properties, etc. A plenty of work has been

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published related to substitution of HA and subsequent properties. However, there is a long journey to be traveled from lab to clinical use. This goal can be achieved by effective cooperation of material scientists, chemists, bioengineers, biologists, clinicians, and industrial partners.

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Carbonate substituted hydroxyapatite

6

Saadat Anwar Siddiqi, Usaid Azhar Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

6.1

Introduction

The human bone is a mixture of inorganic and organic components. The inorganic component of human bone comprises calcium and phosphorus ions, hydroxyapatite (HA), having molecular formula Ca10(PO4)6(OH)2 with 1.67 Ca/P ratio (Porter et al., 2004). Bone is also known as biological apatite and contains up to 2%e8% carbonate, and additionally many other ionic components such as Si, Mg, K, Sr, Zn, etc., may also be present, whose maximum weight composition is up to 4%. However, the capability to produce HA bone tissue is reduced if these substituted ions are not 2þ þ3 2þ þ 2þ þ present. The cationic (SiO4
4 , CO3 , B ) and anionic (Mg , Na , Sr , K , 2þ 2þ 3þ Zn , Ba , and Al ) replacement in the synthetic HA makes it similar to the natural apatite. These are either incorporated in the HA crystal lattice or adsorbed on the surface of HA (Bonfield and Gibson, 2003; Barralet et al., 2003). Ion substitutions can take place in HA for the calcium, phosphate, or for the hydroxyl ions. Cations (Naþ, Mg2þ, Kþ, Sr2þ, Agþ, Zn2þ, Ba2þ, Al3þ, Pbþ2, Laþ3, and Feþ2) substitute the calcium (Caþ2) group in the HA lattice, and anions (SiO4
4 ,
3
CO2þ ,F⁻, and Cl⁻) substitute from the OH or PO group. With each substitution, 3 4 the biological and physiological properties of HA get modified. Silicon, strontium, magnesium, zinc, and fluorine have been known to have the most significant effect on biological properties. Silicon (Si) substitution has been significantly used to increase the rate of bone growth due to Si release or change in the crystallographic orientation and surface chemistry of HA (Porter et al., 2003; Gibson et al., 1999). It is reported that bone resorption and bone formation can be improved by strontium (Sr2þ) substitution in the HA lattice. Magnesium (Mg2þ) substitution also plays a significant role in bone remodeling and helps to activate the osteoblast cells (Aina et al., 2012; Shepherd et al., 2012; Serre et al., 1998). Zinc (Zn2þ)-substituted HA shows antiinflammatory effects and has antibacterial/antifungal nature (Stani
c et al., 2010; Chung et al., 2006). Low amount of silver in silver-substituted HA is considered to be very promising as an antibacterial agent and is also used to improve the cell proliferation and differentiation of osteoblast precursor. Fluorine (F) substitution improves biological interaction by improving cell attachment (Chen and Miao, 2005; Chen et al., 2006).

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00006-9 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Handbook of Ionic Substituted Hydroxyapatites

The carbonate ion (CO2
3 ) has been found to be most abundant ion with calcium (Ca2þ) and phosphate (PO3
4 ) ions, which makes up 2%e8% by weight of inorganic natural component of bone [3]. This gives the carbonated hydroxyapatite (CHA), is more closer to the chemical composition of natural bone. CHA is confirmed as a biodegradable and osteoconductive bioceramic. Its composition is not only similar as a mineral but also similar to biological characteristics. Furthermore, it has a better resorption rate than pure HA particles. That’s why CHA is better suited for bone tissue engineering (Chang et al., 2016; Dutta et al., 2015). Artificial bone graft substitutes and porous scaffolds should be identical to the natural bone in terms of porosity, density, and biological properties. This area has been the subject of intense global research. Recently, bone tissue engineering provides an alternative approach for the repair of bone (Shen et al., 2017; Liu et al., 2016). To meet up the requirement of clinical applications, materials for bone tissue engineering develop a bioactive three-dimensional structure with a pore size of 100e400 microns to stimulate the cell adhesion and diffusion (Ostrowska et al., 2016; Chang et al., 2016). Moreover, to speedup the formation of new bone tissue, scaffolding should also possess positive osteoinductivity. Inorganiceorganic composite scaffold shows excellent osteoconductivity and biocompatibility; however, they lack osteoinductivity (Dutta et al., 2015; Guo et al., 2015; Guan et al., 2015). Two common approaches have been developed to improve the osteogenic differentiation; first, introducing the cytokines into the scaffold and second, incorporating trace elements into the bone scaffolds. The chemical compositions and the porous structure of the natural bone inspire us to fabricate different types of organic and inorganic bone scaffolds. Freeze-drying and electrospinning are simple techniques that produce complex-shaped and highly porous ceramic or polymeric scaffolds for the biomedical applications. This chapter focuses on the crystal structure and crystal chemistry of synthetic CHA and the use of different methods to prepare CHA and in particular, describing the porous composite scaffold for understanding the nature and inorganic chemistry of bone mineral. Our work describes various approaches to the determination of carbonate apatite structures, using single-crystal X-ray structure method in combination with Fourier-transform infrared (FTIR) spectroscopy. These materials are highly biocompatible and expected to yield important benefits for regeneration of bone by using resorbable scaffolds.

6.1.1

Hydroxyapatite

Hydroxyapatite (HA) is the key element of human’s hard tissues, i.e., bone and tooth. It is an ionic crystal with a hexagonal structure. The chemical formula of HA is Ca5(PO4)3(OH), but it is usually written as Ca10(PO4)6(OH)2, which indicates that HA unit cell consists of two units (Porter et al., 2004). HA is extracted from the collagen through noncollagenous proteins such as osteopontin, osteocalcin, and osteonectin. These proteins create about 3%e5% of the bone and also active sites for biomineralization and cellular attachment. Naturally occurring HA and synthetic

Carbonate substituted hydroxyapatite

151

Ca O P H P1 O2

Ca1

Ca2 P2

H2 P3

c

O1 H1

a

b

Figure 6.1 Schematic HA structure (Ren et al., 2013).

HA differ in their constitution mostly with respect to stoichiometry. Normally, synthetic HA is stoichiometric, whereas human bone is not pure or stoichiometric as it contains many other ions, carbonate being the mostly abundant one. Because of this compositional requirement, many researchers who are working on HA synthesis are trying to prepare HA powders with suitable ion substitution. However, how HA is loaded on collagen fibers with orientation in nanoscaled hard tissue is still unclear, and investigation of how to compensate for HA’s brittleness in biomedical applications is an ongoing research area. Fig. 6.1 shows schematic and crystal structure of HA. As it can be seen in Fig. 6.1, HA has a hexagonal crystalline structure, and it elongates along the c-axis. There are many phosphate ions in the a-face, whereas c-face is rich with calcium ions. There are two common processes to prepare HA; wet process in solution and dry process in high temperature for synthesizing (Levitt et al., 1969; Yong et al., 2004; Rhee, 2002; Zhang and Vecchio, 2007; Chaudhry et al., 2006). In a typical wet synthesis process, HA can be produced by dropping phosphoric precursor such as phosphoric acid into calcium precursor solution such as calcium hydroxide solution while maintaining a ratio of Ca/P 1.67 and pH basic with 2þ ammonia solution. Several combinations of PO3
precursors are examined 4 and Ca in many researches (Rhee, 2002; Chaudhry et al., 2006; Jillavenkatesa and Condrate Sr, 1998). Hydrolysis is also a common method to synthesize HA in wet condition. In this case, other forms of calcium phosphate such as dicalcium phosphate anhydrous (Wang et al., 2010), dicalcium phosphate dihydrate (Zhang and Vecchio, 2007), and tricalcium phosphate (Kannan et al., 2008) are hydrolyzed in alkali solutions. It is well known that these phase transitions and resulted morphology depend on the pH value, temperature, and existence of ions including calcium and phosphate (Liu et al., 2003; Sadat-Shojai et al., 2013). The solid-state transition is another way to fabricate HA in the dry mode. Stoichiometric ratios of calcium salt and phosphoric salt are mixed and then these powders are burned around 900e1200 C. Reactant products with high crystallinity can be obtained through this method; however, they often show heterogeneous phase in their composition (Palmer et al., 2008).

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Table 6.1 Possible substitution in apatite structure (A10(BO4)6X2). A10

BO4

X

Ca2þ

PO3
4

OH


Sr2þ

CO2
3

CO2
3

Mg2þ

HPO2
4

F


Kþ

VO2
3

Cl


Naþ

SO2
4

I


Mn2þ

CO3F3


Br


Pb2þ

O2


Zn2þ

6.1.2

Substitution of hydroxyapatite

The chemical characteristics of substituted HA apatite lattice will depend on the revised structure of HA. Ions may substitute for calcium, phosphorus, or hydroxyl groups. These substitutions can occur independently; only one ion or simultaneously more than one ion can be substituted. If these substitutions occur at the same time, they may have a synergistic or antagonistic effect on the properties of apatite. The apatite can be expressed as A10(BO4)6X2. Ions such as Ca2þ, Sr2þ, Mg2þ, Mn2þ, Naþ, and 2
2
Kþ can be replaced in A site and ions such as PO3
4 , CO3 , and SiO4 can be put in BO4 site. For the X site, no metallic ions can be substituted (Zilm et al., 2016). These substitutions may change the properties of the crystal structure including change in the lattice structure, solubility, and morphology. The possible substitutions for each group are shown in Table 6.1.

6.1.3

Carbonate substituted hydroxyapatite

The major composition of bone and tooth is CHA. The chemical formula is generally written as (Ca10(PO4)6
x(CO3)x(OH)2 (Cahyanto et al., 2015), indicating that the CHA unit cell consists of three units. It is phosphorus-, carbonate-, and calcium-based. CHA is a better biological apatite due to its chemical stability, crystallinity, and biodegradability as compared with pure HA. However, HA in bones and teeth is not pure HA, but it contains a few carbonate and some other ions. Human bones contain 7.4 wt% of carbonate ions and teeth have 5.6 wt% of them. The amount of carbonate ions in human is totally dependent on the source species (bone and tooth) and also the age, ranging all the way from 2% to 8% (Zhu et al., 2015; Boskey and Coleman, 2010). Many researchers have developed mechanisms for the synthesis of nanosized carbonate-substituted HA by substituting phosphate (PO4) or hydroxyl (OH) ions with carbonate (CO3) ions. Nanosized CHA powder shows increased biocompatibility and biodegradability, promoting interesting results in the study of in vivo bone defect (Calasans-Maia et al., 2015). It has a very promising application in bone tissue engineering due to osteoinductive effect and possesses highly bioresorbable properties

Carbonate substituted hydroxyapatite

153

Oa C

Ob

Oc

B-type

Ob

A-type c C

Oc

Oa

a

b

Figure 6.2 Crystal structure model of carbonated hydroxyapatite (CHA) showing OH and PO4 and CO3 positions (Ren et al., 2013; El Feki et al., 2000).

(Germaini et al., 2017; Nakamura et al., 2016). Furthermore, CHA improves mechanical strength and is used as in situ hardening bone fillers. Suchanek et al. (2002) prepared CHA powders via mixing CaCO3, Ca(OH)2, and (NH4)2HPO4 by using mechanicalechemical methods and concentration of carbonate ranging between 0.1 and 12 wt%. Kinoshita et al. (2005) used urea (CH4N2O) as a carbonate source to synthesize CHA powder and it is mixed in a solution of (NH4)2HPO4 and Ca(NO3)2 via precipitation method at 100 C. Toriyama et al. (1995) also synthesized CHA by an oxidative decomposition of Ca-EDTA in a phosphate solution with hydrogen peroxide (H2O2). Monica et al. (Calasans-Maia et al., 2015) prepared the CHA at low temperature range (5 C) by using Ca(NO3)2.4H2O, (NH4)2HPO4, and (NH4)2CO3 reagents. On the basis of these experimental results, the carbonate can be substituted from two anionic sites of HA lattice and hence CHA, Ca10(PO4)6
x(CO3)x(OH)2, can be classified into three types.
➢ A-type CHA (CO2
3 substitutes OH ) 2
3
➢ B-type CHA (CO3 substitutes PO4 )
3
➢ AB-type CHA (CO2
3 substitutes both OH and PO4 )


In HA lattice structure, CO2
3 ion is exchanged with the OH ions and is assigned 2
3
A-type carbonate substitution. The CO3 ions replaced by PO4 ions are assigned as B-type carbonate substitution. AB-type carbonated substitution is the replacement of 3

CO2
3 ions by both PO4 and OH ions. Fig. 6.2 illustrates a general lattice structure of CHA.

6.2

A-type carbonated hydroxyapatite


CO2
3 ion replacing OH ions is the type-A carbonate substitution, and it is chemically written as Ca10(PO4)6(OH)2
2y(CO3)y, where 0  y  1. The amount of CO3

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Handbook of Ionic Substituted Hydroxyapatites

substitution in synthetic A-type CHA is relatively low, 4e6 wt%, as compared with its substitution in B-type CHA. A-type carbonate substitution, formed synthetically, has been reported by many researchers (LeGeros et al., 1969; Gibson and Bonfield, 2002). In this type of substitution, an extension of apatite structure results in an increase in a lattice parameter, and the orientation of carbonate ions is generally parallel to the a/c plane (Fleet et al., 2004). Fleet et al. (Fleet and Liu, 2003) indicated that chemical symmetry of type-A CHA is p3, and carbonate ion is oriented with two oxygen atoms close to the c-axis. Honey et al. (Madupalli et al., 2017) also reported the greater disruption with in the c-axis channels in A-type substitution as compared with other types. Suetsugu et al. (2000) showed that the CO3 ions are unevenly occupied with six equivalent places around the c-axis and its triangular plane parallel to the c-axis. High temperature is required to prepare the A-type CHA. Elliott et al. (1985) reported the high-temperature synthesis of type-A CHA, who reacted Ca3(PO4)2 and CaCO3 at 900 C in dry CO2. Bonel et al. (Labarthe et al., 1973) also investigated the synthesis of type-A CHA, when HA is heated at about 900 C in a stream of dry CO2 at atmospheric pressure for several days. As a result, there is progressive increase in a unit cell edge and decrease in c-axis. Under these conditions, type-A CHA is formed by a carbonate substitute for hydroxyl group in the c-axis channel, according to the reaction given below: Ca10(PO4)6(OH)2 þ xCO2 Ca10(PO4)6(OH)2
2x (CO3)x þ H2O Stephen et al. (2007) reported the substitution of A-type CO2
3 by providing the environment to pure HA at 1000 C for 3.5 days. Li et al. (2012) reported the nanostructured porous A-type carbonated CHA fabricated by using activated carbon microwave technique and employed CO2 as the carbonate source. The in vitro biological tests of A-type CHA (Chaudhry et al., 2006) showed that CHA is biodegradable and promotes cell proliferation and differentiation. This biomaterial has greater bioactivity than other HA ceramics and could be a much better candidate for clinical use.

6.3

B-type carbonated hydroxyapatite

3
When CO2
3 ion replaces PO4 ions in HA, it results in type-B CHA and is chemically written as Ca10
x(PO4)6
x(CO3)x(OH)2
x, where 0  x  2. Carbonated apatites of B-type CHA are generally described by reduction in a lattice parameter and an increase in c lattice parameter as PO4 tetrahedra are larger than CO3 triangles. In this type of substitution, CHA crystallinity gets reduced, but the solubility of CO3 ions in the apatite lattice (Anwar et al., 2016) gets increased. Ivanova et al. (2001) described that CO3 ions unevenly occupied outside (side face) of the tetrahedron, whereas Leventouri (2006) established the system of CO3 triangle substitution on the reflect plane of the PO4 tetrahedron, and the CO3 group is located almost flat of the b/c plane in apatite lattice with a small angle of q ¼ 88 degrees. The aqueous precipitation method of Zapanta-LeGeros (1965) and LeGeros (1967) has been widely adopted as a starting point for the synthesis of B-type CHA.

Carbonate substituted hydroxyapatite

155

The method involves the dropwise addition of calcium acetate solution into phosphatecarbonate solutions, maintained at a temperature between 25 and 100 C. ZapantaLeGeros experimented with molar ratios of carbonate to phosphate from 0 to 50. To ensure homogeneous precipitation, the amount of phosphate used was four times the stoichiometric amount needed to make HA. In a typical synthesis, 250 mL of 0.02 M calcium acetate was dropped at a regulated rate into a mechanically stirred mixture of 120 mL of 0.1 M Na2HPO4 and 600 mL of 0.1 M NaHCO3. LeGeros (1991) noted that B-type CHA is also prepared by the hydrolysis of CaHPO4, CaHPO4$H2O (brushite), or octacalcium phosphate in carbonatecontaining solutions. For monetite starting material, 2 g of reagent grade CaHPO4 is hydrolyzed in 1 L of 0.01e0.4 M NaHCO3 solution for 5 h at 95e100 C, with the precipitate filtered, washed, and dried at 60 C overnight. The amount of carbonate in the apatite increases with the concentration in solution, but only to a maximum of 22 wt%, as excess carbonate results in the formation of aragonite. B-type CHA was also synthesized by Wilson et al. (2004). Four liters of 0.04 mol/L calcium acetate solution was added over a period of 3 hours to 12 L of a solution of 0.022 mol/L Na2HPO4 and 0.54 mol/L NaHCO3 held at 95 C. The solution was kept around 90e95 C for 5 days, and then the apatite solution was filtered, washed, and dried at 80 C to obtained 18g yield of sample. Wilson et al. also (Wilson et al., 2006) prepared sodium-free type-B CHA by hydrolysis of monetite (CaHPO4). Seven nominally identical carbonate apatites were prepared by reacting the monetite with hot ammonium carbonate solution following the hydrolysis procedure. The amount of (NH4)2CO3 was selected just below the point at which calcite was consistently precipitated as examine by preliminary experiments, where they fixed the weight of CaHPO4 reacted with increasing weights of (NH4)2CO3. In the actual preparations, 20 g of CaHPO4 was added to 150 g of (NH4)2CO3 dissolved in 10 L of deionized water. The solution was maintained a final pH of 9e10 and kept at 70 C for 3 days. The solution was filtered off, washed well, and dried at 100 C for 24 h. Leventouri et al. (2001) prepared low-temperature B-type CHA by hydrolysis of 0.15 and 0.25 M solutions of sodium bicarbonate with the addition of 7.2 g calcium hydrogen phosphate. Each solution was cooked around 100 C for several hours. The CHA precipitates were then filtered, washed, and dried overnight at approximately 60 C. Tolga et al. (Demirtas¸ et al., 2015) reported the precipitation method that was advanced by microwave light at 600W. It is an ideal condition for the preparation of B-type CHA with 9.21% (wt.) carbonate content, 1.62 Ca/P molar ratio. As a product, bone-like highly pure CHA was produced. Guo et al. (2011) prepared mesoporous B-type carbonated substituted HA by using hydrothermal method at 120 C. This mesoporous material having w5 mm in diameter was composed of many nanoparticles inside the entire microspheres. Zou et al. reported the sonochemistry-assisted microwave method, which proved to be an effective, ultrafast, and simple technique to synthesize CHA nanopowders with Ca/P molar ratio near the theoretical value. Rodlike crystals were produced with a width of around 8 nm and length of around 30 nm (Zou et al., 2012).

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B-type carbonateesubstituted HA ceramic is used as a protective coating for dental and bone implants, also used in the form of granules for the insulation of bone defects (Panda et al., 2003; Huang et al., 2015; Tang et al., 2014). It is reported that 3e8 wt% of carbonate ions in B-type substitution can result in faster bone formation and cell classification as compared with pure HA (Suchanek and Yoshimura, 1998). The absence of B-type carbonate ions in the lattice apatite reduces the crystallinity and increases the solubility in both in vivo and in vitro conditions (Porter et al., 2005). CHA exhibited excellent osteinductive properties, and it does not prompt cytotoxicity in the prolonged incubation period and promotes the normal cell differentiation and proliferation (Mohammad et al., 2016).

6.4

Type-AB carbonated hydroxyapatite

3

In type-AB CHA, CO2
3 ion is replaced by both PO4 and OH ions. Biological apatite is mixed AB-type substitution in most of the cases, and the chemical formula is usually written as ((Ca10-x(PO4)6
x(CO3)x(OH)2-x-2y(CO3)y. It is nearly a fully dense ceramic having maximum density over 95%. Substitution of AB-type carbonated HA represents a more complex mode of substitution then single A-type and B-type. Both a and c lattice parameter were found to increase in the AB-type CHA, whereas c/a ratio shows a slight decrease (Sun et al., 2016). Furthermore, the decrease in the a lattice and an increase in the c lattice has also been reported (Gibson and Bonfield, 2002). Some attempts are made to prepare the AB-type carbonate HA powder. Barinov produce biomimic apatite from both A-type and B-type carbonate (CO2
3 ) substitution using the following reagents: calcium oxide (CaO), ammonium phosphate (NH3PO4), and ammonium hydrogen carbonate (NH3HCO3) via microwave irradiation technique. The synthesis reaction is (Barinov et al., 2006)

10CaO þ 6(NH4)2HPO4 þ (NH4)2CO3 þ 4H2O 0 Ca10(PO4)6
yCO3(OH)2
z þ 14NH4OH Barralet demonstrated the synthesis of AB-type carbonated CO2
3 substitution apatite, using various reagents such as (NH4)3PO4 triammonium orthophosphate, Ca(NO3)2 calcium nitrate tetrahydrate, and (NaHCO3) sodium bicarbonate by using coprecipitation technique (Barralet et al., 1998). In Fig. 6.3, it is noticed that with the increasing carbonate amount, the size of apatite crystallites was slightly decreased ranging from 10 to 160 mm and morphology of apatite looks like spheroidal shape (Barralet et al., 1998). Ivanova et al. (2001) reported the type-AB CHA, which was synthesis by reaction of CaCO3 (calcite) with (NH4)H2PO4 (ammonium phosphate monobasic) solution by using hydrothermal method. Fine calcite powder was treated with 0.2 M (NH4)H2PO4 solution in a stainless steel autoclave at 250 C and 1 kbar for 10 days. A minor amount of 2 M NH4OH was added to the starting solution to bring the pH to 9. The reaction product of white color with a grain size of about 10 mm was washed with distilled water and then dried in air.

157

(c)

222

HA JCPDS No. 09–432

330

211

Carbonate substituted hydroxyapatite

(b)

Intensity (a.u)

(a) 32

34

36

(c)

20

30

40

50

104

004

(b) 310 311 113 203 222 312 213 321

300 202 301

102 210

200 111

002

211

30

(a) 60

2θ (degree)

Figure 6.3 X-ray diffraction pattern of (a) hydroxyapatite (HA), (b) carbonated HA (CHA) 3wt% carbonate content, and (c) CHA 6wt% carbonate content (Venkateswarlu et al., 2014).

Single-phase AB-type CHA has also been prepared by direct reaction at high temperatures (Driessens et al., 1983). Appropriate mixtures of calcium hydrogen phosphate (CaHPO4) and calcium carbonate (CaCO3) were heated at 870 C in a CO2 atmosphere with a partial water vapor pressure of 5 mmHg. Physical and chemical analyses indicated that, at a constant CO3/OH ratio in the apatite channel, carbonate substitutes for phosphate on a one-to-one basis.

6.5

Synthesis methods

During the last decade, many different methods have been established. They produced carbonated HA nanoparticles with control of their microstructure. These methods include different types of chemical synthetic processes. The processing conditions of each synthesis can be different over wide ranges. Various synthesis methods have been exploited for preparation of carbonated HA including precipitation technique, hydrothermal technique, ultrasonic irradiation method, and solegel approach, etc.

6.5.1

Precipitation technique

This is the most commonly used method and has gained attraction because of the rapid production, simplicity, and control of the composition and particle size. The first step is to mix the anion solution, i.e., calcium source with cationic solution, which is phosphorus and carbonate solution followed by nucleation, growth, and filtration. Precipitation of HA contains mixing of calcium and phosphate reagents at less than 100 C temperature.

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Handbook of Ionic Substituted Hydroxyapatites

Slosarczyk et al. (Rapacz-Kmita et al., 2004) employed the precipitation method to prepare carbonated HA powder. In their work, calcium nitrate tetrahydrate (Ca(NO3)2$4H2O) was used as calcium source, diammonium hydrogen phosphate (DAHP) ((NH4)2HPO4) was used as the phosphate source, and ammonium bicarbonate (NH4HCO3) was used as the carbonate source. The pH of the solution was maintained at w11 by using ammonia solution. The CO3HA powder was obtained after aging for 24 h (Rapacz-Kmita et al., 2004). Similarly, Zhou and coworkers (Zhou et al., 2008) synthesized CO3HAp from an acetone solution of Ca(NO3)2$4H2O with a mixed solution of (NH4)2HPO4 and NH4HCO3. The solution pH was adjusted to w11 by using NH4OH. To precipitate CO3Hap, the prepared calcium phosphate solution was heated at 25, 37, or 55 C for only 0.5 min under magnetic stirring. The solutions were immediately filtered under vacuum conditions then washed by using deionized water. There are many other investigations that report the precipitation method to synthesize carbonated HA powder (Shu et al., 2007; Liao et al., 2005a; Othman et al., 2016; Bang et al., 2015). Mortier et al. (1989) also used a homogeneous precipitation process in which the fibrous CHA precipitates were fabricated from a solution of (NH4)2HPO4, Ca(NO3)2, and urea at 120e150 C.

6.5.2

Hydrothermal technique

In the 20th century, the hydrothermal method for preparation of materials was clearly recognized as an important technique (Suchanek and Riman, 2006). It is one of the simplest processes for the production of bioceramics, which is generally known as the reaction of chemical products in an aqueous solution at high temperature and pressure conditions. Hydrothermal synthesis is the simple chemical precipitation method in which the aging step is carried out at a high temperature, generally in an autoclave pressure vessel and treated above the boiling point of the water. The phase purity of carbonated HA precipitates is significantly improved by increasing the hydrothermal temperature. Elongated structures can be made by this process. Some studies show the carbonated HA rodlike shapes are produced in acidic conditions (Lin et al., 2008; Joki
c et al., 2011) for natural (Yoshimura et al., 2004) or synthesized, and also nanorods are synthesized under alkaline conditions (Loo et al., 2008; Jia et al., 2010; Abdal-hay et al., 2014). Guo et al. (2011) reported the hydrothermal method for the preparation of carbonated HA microspheres. In their experiments, 5.0 g of DAHP was dissolved into 20 mL deionized water, followed by the addition of 0.6 g of calcium carbonate microspheres. The mixture was sealed into the stainless steel vessel for hydrothermal treatment at 140 C for 12 h. Xue et al. (2015) fabricated highly crystalline carbonated-HA nanorods with different carbonate content via hydrothermal method. Ca(NO3)2.4H2O, NH4HCO3, and (NH4)2HPO4 were used as a calcium, carbonate, and phosphorus precursors, respectively. The carbonate and phosphorus solution was added dropwise to a solution of calcium, EDTA, and CTAB by keeping the pH around 9e11 by the addition of NH4OH. The suspension was poured into the stainless steel vessel for hydrothermal treatment at 180 C for 24 h and then cooled down to room temperature. CO3HA nanorods were washed from ethanol and deionized water.

Carbonate substituted hydroxyapatite

159

Suchanek et al. (2002) synthesized the CO3HA and NaCO3HA powders by a heterogeneous reaction between Ca(OH)2, Na2CO3, CaCO3, and (NH4)2HPO4 solution at room temperature by using the mechanoechemicalehydrothermal technique, where the carbonate concentration was in the range of 0.8 and 12 by weight%. AA Chaudhry et al. (2013) synthesized the B-type CHA nanoparticles by using continuous hydrothermal flow synthesis (CHFS) method. They used urea as carbonate source and prepared five solutions with varying carbonate contents. Slurries of urea and DAHP and that of calcium nitrate tetrahydrate were prepared separately and added to deionized water at 400 C in a reactor. The pH value was kept close to 10 by using ammonium hydroxide. These solutions are added with the help of HPLC pumps, and the CHFS system allows the precursor solutions flow with controlled speed. Both the slurries meet in a superheated stainless steel container at 400 C and 24 MPa pressure. The product slurry is then centrifuged, washed with deionized water, and freeze-dried (Chaudhry et al., 2013).

6.6

Characterization techniques

Many characterization techniques have been used for the identification of carbonated HA in the form of nanoparticles, powders, rods, and dicks, etc. These characterization techniques are usually classified as direct visualization and spectroscopic techniques. X-ray diffraction (XRD), FTIR, and Raman spectroscopy are regularly used techniques to characterize bioceramics. Spectroscopic techniques such as FTIR and Raman are used to confirm the attached functional groups and chemical composition of HA. Both techniques use specific properties of the material to determine the composition. X-ray photoelectronic spectroscopy determines the chemical structure and amount of carbonate ions into the apatite lattice CHA structure (Anwar et al., 2016; Xue et al., 2015). Solid-state nuclear magnetic resonance spectroscopy is a handy tool to find the crystallinity and structural information of carbonate substituted apatite and bone mineral (McElderry et al., 2013). The thermal behavior of CHA can be investigated by using thermogravimetric analysis/differential scanning calorimetry, which is attributed to endothermic and exothermic loss of weight at different temperature range (Karunakaran et al., 2019).

6.6.1

X-ray diffraction

X-ray powder diffraction is a technique that is used to calculate the atomic and molecular structure of the crystal. Many researchers have studied the effect of carbonate substitution on lattice parameters. These lattice parameter changes according to different types of CO3 substitution (A, B, and AB) (Krajewski et al., 2005). The XRD spectra of carbonates are similar to the spectra of HA. They show the different peaks, which are well matched with the diffraction pattern of the standard JCPDS card no 09e0432 of HA. The characteristic peak of HA and CHA fall in the region of 31e34 degrees. CHA is also apportioned to the hexagonal crystalline structure with a spatial group P63/m. The crystallinity is decreased when carbonate ions are added

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Handbook of Ionic Substituted Hydroxyapatites

into the HA lattices. Venkateswarlu et al. (2014) reported the XRD analysis of CHA as compared with HA in Fig. 6.3. Type-A carbonate substitution causes an extension of apatite structure, which results in an increase in a lattice parameter, and the orientation of carbonate ions is generally as parallel to the a/c plane. In case of B-type carbonate substitution, there is a reduction in a lattice parameter and an increase in c lattice parameter, whereas, in AB-type carbonate substitution, the lattice parameter (a and c) can increase or decrease. However, a certain disagreement is found in the literature about the lattice parameter of all types of carbonate-substituted HA.

6.6.2

Fourier-transform infrared spectroscopy

FTIR has proven to be one of the most useful molecular spectroscopy techniques for studying analog and biological apatite materials. FTIR spectroscopy is used routinely both to detect the presence of carbonate and to characterize synthetic carbonate apatite materials because A, B, and AB CHA are not readily distinguished using X-ray powder diffraction methods due to the low contents of carbonate. Although FTIR detects even minor amounts of carbonate in apatite, this method can yield ambiguous results for the proportion of type-A and type-B carbonate ions using the asymmetric stretch (n3) spectral region because of band overlap. Powdered FTIR spectra for the carbonate apatite crystals synthesized at high temperature and pressure can be seen in the studies of Fleet and coworkers (Fleet, 2014). The spectra are illustrated by a complex band of asymmetric stretching vibration of the phosphate group at around 1000e1100 cm
1. The characteristic bands for carbonate occur in the 1600e1400 cm
1 (asymmetric stretch, n3) and 880e862 cm
1 (out-ofplane bend, n2) spectral regions. Furthermore, a weak band for the stretch vibration of hydroxyl (OH) group may be present in CHA samples near around 3570 cm
1 and echoed by an OH liberation band near around 631 cm
1. The assignment of band is summarized in Table 6.2. Table 6.2 Carbonated hydroxyapatite Fourier-transform infrared peak assignment according to frequency range and functional groups. Wavenumber (cmL1)

Peak assignment

1494

n3 stretching mode of the CO2
3 in CHA

Koutsopoulos (2002)

1453

n4 or n3 bending mode of the (CeO bond) in A- and B-type CAP and n1 stretching mode of CO23 (CeO bond) in B-type CHA

Koutsopoulos (2002)

1544 and 1461

Asymmetric stretch (n3) of CO2
3 in AB-type CHA

Fleet and Liu (2003)

References

CO23

Carbonate substituted hydroxyapatite

161

Table 6.2 Continued Wavenumber (cmL1)

Peak assignment

References

1540 and 1449

CO2
3

Asymmetric stretch (n3) of assignments for A-type CHA

(Fleet and Liu, 2004; McClellan and Van Kauwenbergh, 1990)

1474 and 1409

Asymmetric stretch (n3) of CO2
3 assignments for B-type CHA

(Fleet and Liu, 2004; McClellan and Van Kauwenbergh, 1990)

566, 571

n4b triply degenerates bending mode of the PO3
4 (OePeO bond)

Panda et al. (2003)

1455, 1410 and 871

B-type v3 stretching mode of the CO2
3

(LeGeros et al., 1969; Nelson and Featherstone, 1982)

1050, 1088,1034

n3b triply degenerates asymmetric stretching mode of PO3
4 (PeO bond)

(Suchanek et al., 2002; Rehman and Bonfield, 1997)

762, 701, 759, 700

n4 in-plane deformation bending mode of CO2
3 (OeCeO bond)

3570

ns stretching mode of OeH

(Koutsopoulos, 2002; Neira et al., 2008)

1546

n4 bending mode of CO2
3 (A-type)

(Fleet and Liu, 2004, 2005; Neira et al., 2008)

1494

n3 stretching mode of CO2
3 in CAP (A-type of B-type)

Suchanek et al. (2002)

1455

n3 or n4 bending mode of CO2
3 (A- and B-type)

Youness et al. (2017)

1413

n3 stretching mode of CO2
3 (B-type)

Zhou et al. (2008)

961

n1 nondegenerates symmetric stretching mode of PO3
4 (PeO bond)

(Youness et al., 2017; Ulian et al., 2013)

879

n2 bending mode of CO2
3 (A-type)

Jokanovi
c et al. (2006)

633

nL, vibrational mode of OH (OeH bond)

Rehman and Bonfield (1997)

602

n4a triply degenerates bending mode of PO3
4 (OePeO bond)

Mostafa et al. (2011)

583, 572

n2 doubly degenerates bending mode of PO3
4 (OePeO bond)

Mostafa et al. (2011)

162

6.7

Handbook of Ionic Substituted Hydroxyapatites

Carbonated hydroxyapatite as a coating material

Because of diseases, accidents, and aging, orthopedic and dental biomedical materials are one of the most wanted materials to address these issues in the last few decades. Titanium and stainless alloys are mostly used when people get their bones fractured or broken or they need to have joints replaced that can support high loads. To repair a fractured bone, the essential requirement for the bone implant material is that it not only heals the bone defect but also is used for remodeling of bone. There are few difficulties in the production of materials that are suitable for the bone implant. For example, the first difficulty that occurs frequently is the synchronization between the implant material and the bone during the bone remodeling process (reabsorption and repair process). The implant material should also not affect the immune system. So far, bone implants produced with HA or with natural biocompatible substances are being widely used to treat bone defects. Furthermore, all these techniques have their limitations. For example, implant materials are sometimes reabsorbed before osteogenesis (Gristina, 1987; Hetrick and Schoenfisch, 2006). In other cases, infections due to implant result in the onset of diseases and in some cases even leads to death. For these reasons, it is necessary to find a way to design and produce implants suitable for the regeneration of bone tissue. Many methods have been used to make the implants biocompatible and safe for the patient. One of the most popular technique is to coat the implants with biocompatible and bioactive materials. For the regeneration and implantation of bone tissue, HA is the most commonly used material (Iqbal et al., 2012). Carbonated HA has been also used as a coating material because it is chemically similar to natural bone. A number of articles have been published on CHA as a coating material in the last few years (Li et al., 2018; Kwon et al., 2018; Graziani et al., 2017). Furthermore, the coating of the surfaces has also shown to improve osteointegration. Stigter et al. (2004) used a thin coating of CHA on a Ti implant surface by using a biomimetic precipitation method. The advantage of this method is that incorporation of antibiotics in CHA coatings might be used to prevent postsurgical infections and to promote bone bonding of orthopedic devices. Other known methods that have been used for the CHA coating on the surface of the materials include laser deposition, electrophoretic deposition, immersion coating, hot isostatic pressing, electrostatic spraying, and thermal spraying. Darr et al. (2004) for first time reported CHA coating by using metal organic chemical vapor deposition technique. They deposited carbonated HA powder onto the thick Ti6Al4V metal plates. Julietta et al. (Stigter et al., 2004) prepared the CHA films onto Ti substrates by using pulsed laser deposition (PLD) technique to improve its bioactivity and biocompatibility. Hidalgo-Robatto et al. (2018) fabricated the fluorine-incorporated CHA (FCHA) coatings onto the titanium alloy substrate by using PLD method. Fig. 6.4 illustrates the micrograph of FCHA powder coatings by varying the water vapor pressure (0.0 and 0.45 mbar) and time (30 and 120 mins). Furthermore, carbonated HA coating usually cannot be strongly bonded with metallic implants, which may get peeled off or removed because of nonuniform crystallinity and poor coatingesubstrate adherence (Hamdi et al., 2000; Yang et al., 2005). These defect can be resolved by adding some organic polymers such as

Carbonate substituted hydroxyapatite

163

(a)

(b)

(c)

(d)

(e)

(f)

Figure 6.4 Scanning electron microscopy images of titanium alloy coatings fluorineincorporated CHA coatings at (a) 0 mbar for 30 min, (b) 0.45 mbar for 30 min, (c) 0.45 mbar for 120 min, and (def) images were taken on higher magnifications.

Figure 6.5 Scanning electon microscopy cross-sectional image of chitosan/carbonated hydroxyapatite coating on Ti implant (Tang et al., 2014).

poly(methyl methacrylate) PMMA, chitosan, and gelatin, etc., with CHA powder. Tang et al. (2014) demonstrate the chitosan/CHA composite coatings on to Ti6Al4V substrate by electrophoretic deposition method. SEM images (Fig. 6.5) demonstrate that CHA particles exhibit a plate-like structure, and these plate agglomerated to form a microporous structure having a pore size ranging from 0.5 to 2 mm. They also discussed cross-sectional view of coatings; interestingly, the thickness of

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Handbook of Ionic Substituted Hydroxyapatites

chitosan/CHA is around 20 mm; 2 mm Ti oxide layer is produced between the substrate and coating area, which may improve the physiochemical properties and mechanical interlocking.

6.8

Carbonated hydroxyapatiteebased composite materials

It is well-known that bioceramics such as HA are hard but brittle. On the other hand, some biomedical devices such as joints receive high load, so they require high toughness. A Young’s modulus of natural compact bone is said to be 8e24 GPa, whereas the modulus of an inorganic component, HA, is 130 GPa. This is because an organic component, collagen, has high elasticity. Its Young’s modulus is only 1.25 GPa (Meyers et al., 2008; Chen et al., 2012). To fabricate real bone-like materials, it is difficult to obtain both toughness and pliability by using HA and CHA only. That is because it has higher Young’s modulus than natural bones and low fracture toughness (Miyazaki et al., 2013). Therefore, composites have attracted attention to reproduce artificial bones. Typical composite materials for biomedical application are metal-ceramic. They are already commercially used. Concerning some metallic joints, only the part which interacts to bone is covered with HAp to give a high wear resistance and osseoconductivity (Meyers et al., 2008; Bauer et al., 2013; Markovi
c et al., 2011). Polymer-ceramic composites are also well investigated and are very useful in the field of tissue regeneration. As CHA is too brittle and weak to bend or impact, polymers support these drawbacks. The biodegradable polymers are used in bone regeneration. Urist et al. concluded that polymers induced bone formation in animals (Urist, 1965; Urist and Strates, 1971). Similarly, Kulkarni et al. reported that the morphology of polymerbased implants and implant sites is essential to stimulate bone tissue regeneration. They suggested that tubular resorbable polymeric implants could improve bone tissue regeneration in long bones (Kulkarni et al., 1971). Several synthetic polymers were tested directly or in combination with HA or CHA to mimic the natural bone matrix. Among these, the most common are PLGA of polyester (Zhang et al., 2009; Flahiff et al., 1996), PLA (Higashi et al., 1986), and poly (caprolactone) (Zhao et al., 2008). In most polymeric compounds, CHA has been used as an enhancer to impart mechanical strength and bioactivity to the composite. Although research into the use of CHA compounds is progressing slowly, it is constantly becoming more prominent and significant. Researchers have fabricated the CHA scaffolds for use of drug delivery carrier, bone tissue regeneration, and the treatment of various bone defects such as osteoporosis. Liao et al. (2005b, 2007) fabricated the nano-CHA/collagen composite at room temperature via biomimetic self-assembly method. These composite is promising for hard tissue therapy. Rakmae et al. (2011, 2012) prepared CHA/PLA compound while Hong et al. synthesized a CHA/PLGA compound. They determined that these compounds had adequate mechanical strength to be used in bone regeneration. Mao et al. (2009) prepared a CHA/PMMA compound and showed that this material served as better osteointegrated properties and ideal

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biomaterial for bone repair. It has also been shown that chitosan and CHA improve the biocompatibility and mechanical strength of chitosan-based materials (Chen et al., 2018). A compound containing chitosan and high molecular weight CHA particles was prepared by electrospinning, freezing drying, and lyophilization (Tang et al., 2014; Tong et al., 2010; Park and Kim, 2017). The compound showed improved mechanical properties than medium molecular weight chitosan. Furthermore, it has been reported that different compositions of collagen with CHA have better biocompatibility and enzymatic degradability (Liao et al., 2007; Takallu et al., 2018).

6.9

Biological studies

CHA is chemically similar mineral phase of the bone, which provides excellent biological properties. The biodegradability of HA is low as compared with carbonated substituted HA. To achieve good biological properties, material must have an ability to degrade at suitable rate and maintain a balance between cell growth and also regeneration of new bone tissues (Calasans-Maia et al., 2015). Thereby, the incorporation of carbonate into HA increases its in vivo dissolution rate and allows the production of bone substitutes with increased resorptive properties (McElderry et al., 2013). Recent studies have been shown that CHA are biocompatible, osteoinductive, osteoconductive, and show increased biodegradation rate. It also promotes interesting results during in vivo bone defect studies (Calasans-Maia et al., 2015) and important part in reconstructive dentistry (Zhao et al., 2011; Aerssens et al., 1998). Many researcher reported the in vivo degradation of CHA spheres and granules, which promotes the bone formation. dos Anjos et al. (2019) reported that even the high dose of carbonate into HA did not induce significant cytotoxic in murine preosteoblasts cells. Rupani et al. (2012) have also demonstrated that CHA is a biocompactable material having increased biodegradation properties and osteogenic capability as compared with HA. Pieters et al. (2010) reported the fabrication of B-type CHA pellets with different values of carbonate content (3.4, 5.2, 6.9, 11.6, 16.0 wt%) and checked the MC3T3-E1 cells behavior. Low percentage of carbonate exhibits low cell adhesion, but the at higher percentage (11.6 and 16.0 wt%), cell adhesion is significantly higher. Kasai et al. (2010) reported fabrication of sintered porous CHA scaffolds and grow bone marrow mesenchymal stem cells (MCSc) derived from syngeneic rats. They establish the new bone formation within 4e8 weeks. Hesaraki et al. (2014) investigate the comparison of CHA with simple HA, and the Wistar ratederived MCS and fibroblasts L929 seeded on n-CHA revealed better cell proliferation and cell differentiation. Hollow HA and CHA microsphere having a high surface area also provide the good osteoconductivity and biocompatibility in bone repair. However, nonionic HA has slow resorption rate, and bone remodeling is limited. Xiao et al. prepared the hollow microsphere of HA and CHA and loaded them with bone morphogenetic protein-2 to enhance the bone healing defects. Fig. 6.6 showed the optical images of hematoxylin and eosine and von Kossae strained section of rat calvarial defects with the HA and CHA. It illustrates the some marrow-like tissues were present in the defected area of bone.

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CHA12

HA

(a)

(b)

(c)

(d)

(e)

(f)

Figure 6.6 Transmitted light images of (a, b) von Kossae and (c, d) hematoxylin and eosin (H&E)estained sections of rat calvarial defects implanted with hollow CHA and HA microspheres loaded with BMP2 at 12 weeks postimplantation; (e, f) higher magnification images of the boxed areas are given in (c, d).

Pezzatini et al. (2006) reported morphological behavior of both in vivo and in vitro; the osteoblast cells were attached and proliferated on the surface of CHA. So, incorporation of carbonate ions into the HA is an effective way to enhance cell growth and cellular adherence with the host cells.

6.10

Conclusion

The human bone may contain up to 8 wt% carbonate that can occupy phosphate or hydroxide positions. Carbonate containing HA are of three structural types: A, B, and AB types, depending on whether carbonate has occupied phosphate, hydroxyl, or both the positions. All the three types of CHA have been synthesized by various preparation methods. These different types of CHA exhibit slightly different kind of biological properties. The stability of CHA at higher temperatures is also a matter of concern for its scaffolds preparation needed for tissue engineering applications. This may be a reason that a relatively limited amount of literature is available for its in vivo studies as compared with other ionic substitutions in HA. However, all the studies of cell attachment and degradation reveal superior properties as compared with pure HA.

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Wang, J.-C., et al., 2010. Deriving fast setting properties of tetracalcium phosphate/dicalcium phosphate anhydrous bone cement with nanocrystallites on the reactant surfaces. J. Dent. 38 (2), 158e165. Wilson, R.M., et al., 2004. Rietveld structure refinement of precipitated carbonate apatite using neutron diffraction data. Biomaterials 25 (11), 2205e2213. Wilson, R.M., Dowker, S.E., Elliott, J.C., 2006. Rietveld refinements and spectroscopic structural studies of a Na-free carbonate apatite made by hydrolysis of monetite. Biomaterials 27 (27), 4682e4692. Xue, C., et al., 2015. Hydrothermal synthesis and biocompatibility study of highly crystalline carbonated hydroxyapatite nanorods. Nanoscale Res. Lett. 10 (1), 316. Yang, Y., Kim, K.-H., Ong, J.L., 2005. A review on calcium phosphate coatings produced using a sputtering processdan alternative to plasma spraying. Biomaterials 26 (3), 327e337. Yong, P., et al., 2004. Synthesis of nanophase hydroxyapatite by a Serratia sp. from waste-water containing inorganic phosphate. Biotechnol. Lett. 26 (22), 1723e1730. Yoshimura, M., et al., 2004. Hydrothermal conversion of calcite crystals to hydroxyapatite. Mater. Sci. Eng. C 24 (4), 521e525. Youness, R.A., et al., 2017. Molecular modeling, FTIR spectral characterization and mechanical properties of carbonated-hydroxyapatite prepared by mechanochemical synthesis. Mater. Chem. Phys. 190, 209e218. Zapanta-LeGeros, R., 1965. Effect of carbonate on the lattice parameters of apatite. Nature 206 (4982), 403. Zhang, X., Vecchio, K.S., 2007. Hydrothermal synthesis of hydroxyapatite rods. J. Cryst. Growth 308 (1), 133e140. Zhang, P., et al., 2009. In vivo mineralization and osteogenesis of nanocomposite scaffold of poly (lactide-co-glycolide) and hydroxyapatite surface-grafted with poly (L-lactide). Biomaterials 30 (1), 58e70. Zhao, J., et al., 2008. Preparation of bioactive porous HA/PCL composite scaffolds. Appl. Surf. Sci. 255 (5), 2942e2946. Zhao, J., et al., 2011. Amorphous calcium phosphate and its application in dentistry. Chem. Cent. J. 5 (1), 40. Zhou, W., et al., 2008. Synthesis of carbonated hydroxyapatite nanospheres through nanoemulsion. J. Mater. Sci. Mater. Med. 19 (1), 103e110. Zhu, Q.-X., Li, Y.-M., Han, D., 2015. Co-substitution of carbonate and fluoride in hydroxyapatite: effect on substitution type and content. Front. Mater. Sci. 9 (2), 192e198. Zou, Z., et al., 2012. Ultrafast synthesis and characterization of carbonated hydroxyapatite nanopowders via sonochemistry-assisted microwave process. Ultrason. Sonochem. 19 (6), 1174e1179. Zilm, M., et al., 2016. Hydroxyapatite substituted by transition metals: experiment and theory. Phys. Chem. Chem. Phys. 18 (24), 16457e16465.

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Fluoride-substituted hydroxyapatite

7

Sandleen Feroz 1 , Abdul Samad Khan 2 1 Department of Dental Materials, Foundation University Islamabad Campus, Islamabad, Pakistan; 2Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia

7.1

Structure of hydroxyapatite

The main constituent of vertebral mineralized tissue such as bone and teeth are calcium phosphate (CaP) salts (M€arten et al., 2010; Hu et al., 2010). Fig. 7.1 represents the hierarchical structure of typical cortical or compact bone at various length scales. The inorganic biomineral phase comprises of 60%e70% of bone, while the collagen that forms the organic phase along with water forms the remaining portion. Among the CaP salts, hydroxyapatite possesses the most similarity to the mineral component of the bone. In fact, hydroxyapatite (Ca10(PO4)6(OH)2, HA) crystalline phase presents as a thermodynamically most stable CaP salt in body fluid, with Ca/P ratio of 1.67 (M€arten et al., 2010, Vallet-Regi and Gonzalez-Calbet, 2004). Its crystal structure is  a hexagonal cylinder and each unit cell contains 10Ca2þ, 6PO3 4 , and 2OH . Its theo3 retical density value is as high as 3.156 g/cm , microhardness is 5.0, and refractive index is 1.64e1.65. However, synthetic HA in its pure form has lower symmetry, and it belongs to the monoclinic crystal system, space group P21/b, with unit cell dimensions a ¼ 9.4214(8) Å, b ¼ 2a, c ¼ 6.8814(7) Å, and g ¼ 120 degrees (Calderin et al., 2003). According to several studies (both in vitro and in vivo), the synthetic HA can promote the new bone growth through osteoconductive mechanism. Therefore, it is a material of choice because of its osteogenic potential without causing any local or systemic toxicity or inflammatory response (O’Hare et al., 2010; Gu et al., 2004; Kokubo and Takadama, 2006). After implantation, a tissue-free layer containing carbonated apatite forms on the surface of HA-based ceramic that helps in early stabilization of the implanted material to the surrounding tissues (Kokubo and Takadama, 2006). Recent studies have shown that it also inhibits the abnormal growth of cancerous cells (Rong et al., 2016; Dey et al., 2014; Qing et al., 2012). Among various hydroxyapatite structures, nanosized HA has grain size less than 100 nm in at least one direction and an ultrafine structure, whereby it can readily promote the bone formation as it mimics the natural bone mineral in structure and composition (Cai et al., 2007; Wang et al., 2010). The nanosized HA-based ceramic biomaterials showed enhanced bioactivity and resorbability as compare with micron-sized ceramics (Dong et al., 2009; Dorozhkin, 2012). Similarly, like the

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00007-0 Copyright © 2020 Elsevier Ltd. All rights reserved.

176

Handbook of Ionic Substituted Hydroxyapatites Whole bone

Tissue structure

Nanostructure

Microstructure

Osteab (haversian system) (∼ 200 μm)

Lamellae (∼ 7μm)

Hydroxyapatite (∼ 50 x 25 x 2 nm)

Collagen fibrils (∼ 50nm) Periosteum Spongy bone

Collagen fiber (∼ 5μm) Osteonic canal Collagen molecules

Compact bone Blood vessels

Hydroxyapatite crystals Microscopic view

Collagen triple helix (∼ 300 x 1.5nm)

Mineralized fibrils

Macro

Nano

Figure 7.1 The diagrammatic representation of typical bone hierarchical structure at various levels of magnification. Haversian systems forming the microstructure of cortical or compact bone. At the nanoscale are collagen fibers forming the structural framework (Sadat-Shojai et al., 2013).

biological apatite, the release of calcium ions from the nanosized HA is faster than coarser crystals. Thus, because of improved cellular proliferation, densification, sinterability, and other cellular activities related to bone formation bioceramic-based on nanosized HA is the material of interest for researchers for various biomedical applications (Bose et al., 2009; Bianco et al., 2009).

7.2

Presence of ions in biological apatite

Biological apatite is comprised of hydroxyapatite along with various minor groups, el2 2þ þ  þ  2þ ements, and trace elements such as HPO2 4 , CO3 , Mg , Na , F , K , Cl , and Sr (Skinner, 2005). Human bone specimen were analyzed to determine the ionic concentrations. There was significant increase in the concentration of Naþ and CO2 3 ions with age, whereas a decrease in ionic concentration of Mg2þ, Kþ, and Cl was observed (Handschin and Stern, 1995). The prospective view of HA crystal unit cell is well represented by Fig. 7.2 (Fihri et al., 2017). Four calcium ions per unit cell called Ca (I) are in columns parallel to the c-axis, and nine oxygen atoms surrounds them (Pasteris et al., 2008). The two equilateral triangles along the c-axis (at z ¼ 1/4 and 3/4) are formed by other calcium ions (six per unit cell) called Ca (II), and two monovalent anions occupied them, such 3 as OH, F and, Cl, but rarely by bivalent CO2 3 per unit cell. The six PO4 anions complete the structure, and each P atom is linked to four oxygen atoms. Similarly,

Fluoride-substituted hydroxyapatite

(a)

177

(b) Ca1 Ca2

Ca1 Ca2 P O OH

b

b c

P O OH

a

a

c

(c)

(d)

b c

b a

c

a

Figure 7.2 (a) Projection of the unit cell of HA according to plan (001); (b) projection showing the arrangement of octahedrons (Ca(1)O6) in the HA structure; (c) projection showing the sequence of octahedral (Ca(1)O6) and tetrahedral (PO4) in the HA structure; and (d) projection showing the sequence of octahedral, (Ca(1)O6) and (Ca(2)O6), and also tetrahedral (PO4) in the HA structure (Fihri et al., 2017).

OH groups are not present in some bony apatite and mostly present in low concentration (Ratnayake et al., 2017) (Bigi et al., 2016; Brown and Constantz, 2017).

7.3

Ionic substitution of hydroxyapatite

Synthetic HA is not considered as an ideal bone material substitute when implanted in load-bearing areas due to its brittle nature and lack of strength. Therefore, even minor substitutions have significant effects on mechanical properties, solubility, osteoblastic, and osteoclastic responses both in vivo and in vitro (Kim et al., 2004a; Basirun et al., 2018). There are two main type of substitutions, i.e.,; 1. Cationic substitutions: When the calcium site in HA is partially replaced with ions such as Zn2þ, Mg2þ, or Agþ cationic substitutions occur. 2. Anionic substitutions: Anionic substitutions can be divided into two types

Type A substitutions occur when the hydroxyl ion is replaced by a larger halide (F, Cl) ion as shown in Fig. 7.2 (Ratnayake et al., 2017).

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Handbook of Ionic Substituted Hydroxyapatites

Type B substitutions involve when the phosphate functional groups are replaced by the silicate functional groups (SiO4⁴) or carbonate (Gibson and Bonfield, 2002). Geological environment offers wide degree of temperature and elemental changes because of which these ionic substitutions can be observed well in geological apatite, which makes it more prone to substitution, e.g., F, Cl, and OH anions can be substituted for each other in almost any proportion at the channel sites (Hughes and Rakovan, 2002). The phosphate group can be replaced by anionic complexes, whereas the calcium ions can be substituted by metallic bivalent and trivalent cations (Lala et al., 2016). However, in synthetic HA, the laboratory conditions such as temperature and pressure do not allow the wide range of substitutions. In fact, when these ions are incorporated into the HA structure, it results in various changes in properties, which can be observed with the aid of different analytical techniques. Properties such as solubility level and crystallinity can be affected by these ionic substitutions (Ratnayake et al., 2017). Substitution of biocompatible bivalent cations, such as Mg2þ, Zn2þ, Sr2þ, can result in no charge imbalances in the apatite lattice; however, monovalent ions substitutions such as Kþ and Naþ can cause a charge imbalance. Monovalent anions such as F and Cl can substitute OH anions in the anion channel without causing any charge imbalance. These charge imbalances can be compensated either by the creation of supplementary vacancies or by substitutions of anions and cations simultaneously without formation of any vacancy (Sadat-Shojai et al., 2013). Among all these substitutions, when OH ions in apatitic structure is replaced by fluoride ions, a very interesting biomaterial having thermal and chemical properties far superior than hydroxyapatite is formed.

7.4

Fluoride substitution in hydroxyapatite

Fluoride plays an important role in the prevention of dental caries as it increases resistance of tooth surface mineral against acidic dissolution under low pH conditions in oral cavity (Feroz and Moeen, 2017). Fluoride ions have been used successfully in the treatment of osteoporosis and stimulate the proliferation of osteoblastic cells (Chavassieux et al., 1993; Kim et al., 2004b). Therefore, the fluoride-substituted hydroxyapatite (FHA) has been extensively investigated during the past few decades particularly in the field of biomedical sciences. FHA is obtained when OH ions in HA is substituted partially by F ions, whereas fluorapatite (FA) is formed if this substitution is completed (Bigi et al., 2016). Fluorapatite is most commonly found in enamel layer mineral, and because of its superior mechanical properties, it can be used in several dental applications. Both HA and FA have similar atomic structure, differing only in the substitution of fluoride ion for the hydroxyl group. However, contraction in the a  -axis dimensions to 0.9368 nm is observed as a result of this substitution, with no change in the c-axis dimensions, as F ions is smaller than OH (Sonamuthu et al., 2018). Both fluorohydroxyapatite (F-halfesubstituted hydroxyapatite) and fluorapatite (complete

Fluoride-substituted hydroxyapatite

179

substitution) have been successfully synthesized under control laboratory conditions. Increase in crystallinity, crystal size, fracture toughness, elastic modulus, and decrease in solubility were observed as a result of fluoride ions substitution (Fahami et al., 2016).

7.4.1

Method of preparation of fluoride-substituted hydroxyapatite

Different synthesis routes for the production of fluorohydroxyapatite (with varying degree of F ion concentration) and fluorapatite with different morphologies and crystallinity are as follow (Eslami et al., 2008, 2011; Rintoul et al., 2007; Bianco et al., 2010; Kim et al., 2004b; Tredwin et al., 2013b, 2014; Darroudi et al., 2010; Chen and Miao, 2005; Bir et al., 2012; Stanic et al., 2014; Jokanovic et al., 2013; Szczes et al., 2017). 1. 2. 3. 4. 5. 6. 7. 8. 9. 10.

Wet chemical methods Modified wet chemical process Sol gel Hydrothermal Wet Precipitation Dry solid-state method Multiple emulsion technique Electrode deposition technique Mechanochemical techniques Microwave synthesis

7.4.1.1

Modified wet chemical process

Synthesis of crystalline FHA occurred via pH cycling method by changing the concentration of sodium fluoride (NaF) in hydroxyapatite suspension as a modified wet chemical process (Eslami et al., 2008). This cyclic pH fluctuation was repeated three times. The reaction involved in the preparation of FHA can be explained by the following reactions: 1. Ca10 (PO4)6(OH)2 þ 2xNaF/Ca10 (PO4)6(OH) 2-2xF2x þ xNa2O þ xH2O. 2. Ca10 (PO4)6(OH)2 þ xF þ xHþ /Ca10 (PO4)6(OH) 2-2xF2x þ xH2O.

The obtained micro-FHA was rod shaped as shown in Fig. 7.3, where the X-ray diffraction (XRD) pattern revealed that pure apatite was prepared by this method, and trace amount of impurities of other calcium phosphate was not detected. Secondary phase calcium oxide was observed, and b-tricalcium phosphate phase was not detected showing its thermal stability (Rameshbabu et al., 2006). All the typical absorption characteristics of FHA were observed by Fourier-transform infrared analysis. Additionally, same contents of carbonate were also observed (CO2 3 peak around 1500 and 744 cm1). This clearly indicated the presence of carbonate in apatite crystal structure. Due to OH.F bond, the 3536 cm1 band indicated that some of the OH groups in the apatite crystal structure were replaced by F ions.

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Handbook of Ionic Substituted Hydroxyapatites

(b)

(a)

100 nm

100 nm

Figure 7.3 Transmission electron microscopy images of synthesized (a) fluorohydroxyapatite and (b) fluoroapatite (Rameshbabu et al., 2006).

Thus, multiple wet techniques have been used for the synthesis of FHA nanopar ticles, e.g., direct precipitation from the solutions containing Ca2þ, PO3 4 , and F (Chen and Miao, 2005; Qu and Wei, 2005; Rodrıguez-Lorenzo et al., 2003), by low molecular poly(acrylic acid) (Bertoni et al., 1998), by means of hydrolysis (Kurmaev et al., 2002), by immersion of HA in a sodium fluoride (NaF) solution (RodriguezLorenzo and Gross, 2003), and by surfactant-assisted preparation (Zhang and Zhu, 2005)

7.4.1.2

Dry solid-state method

FHA was prepared by a reaction of HA and CaF2 (0e50 wt.%) followed by the process of calcination at 700 C. Firstly, the OHeFeOH and then OHeF hydrogen bonds of sintered FHA composites were formed with increasing CaF2 wt.% (Wu et al., 2010). Another study showed that the thermal stability of HA crystal structure became improved as the percentage of F ion incorporation increased (Eslami et al., 2009). Tredwin et al. (Tredwin et al., 2013a) revealed that by increasing the F content at all calcination temperature results in increasing the bond strength, and this change was particularly noticeable at 800 and 1000 C. Thus, varying the amount of fluoride substitution can be fine-tuned. Cell culture tests also revealed that high concentration of F content results in low surface potential, which ultimately favors cellular attach ment (Supov a, 2015).

7.4.1.3

Solegel method

Solegel method was used to synthesize HA, FHA, and FA (Tredwin et al., 2014; Khan et al. 2008, 2013). In this method, precursors used were triethyl phosphite (TEP, [P(C2H₅O)3]) and calcium nitrate (Ca(NO3)2.4H2O) under an ethanol and waterbased solution. A general reaction is shown below: 6Ca(NO3)2 þ 6(C2H5O)3P(0)/Ca10(PO4)6(OH)2 þ by product

Fluoride-substituted hydroxyapatite

181

The FHA and FA sols were prepared by using different amounts of ammonium fluoride (NH4F, Aldrich, USA) in the P-containing solution. The molar ratios of P/F were 12, 6, 4, and 3 in correspondence to compositions of Ca10(PO4)6F0.5OH1.5, Ca10(PO4)6F1OH1, Ca10(PO4)6F1.5OH0.5, and Ca10(PO4)6F2 (by replacing OH ions with F ions in molar ratios of 0.25, 0.5, 0.75, and 1, respectively). This study concluded that the varying degree of F ion incorporation influences the cellular attachment, morphology, proliferation, and differentiation of osteoblastic cells, and a significant increase in metabolic activity with increasing concentration of fluoride content was observed. It was also revealed that increasing the incorporation of F ions into the apatitic structure would result in a significant decrease in the temperature of crystallization along with more compact unit cell structure. Furthermore, this incorporation of F also affects the rheological properties of the sol gel.

7.4.1.4

Other processes

Most wet chemical methods used for the synthesis of FHA nanoparticles need precise and careful monitoring of the whole process. If all the parameters are not strictly followed, the properties of the final products are greatly influenced. Additionally, to obtain final products with proper crystallinity, long postheat treatments are needed because of which wet methods are not suitable for mass production on industrial scale (Zhang et al., 2005; Kim et al., 2009; Fathi and Zahrani, 2009). Duff first developed the cyclic pH method to prepare the solid solution of HA, FA, and FHA samples (Duff, 1975). For production on a larger scale and nanostructured FHA, mechanochemical methods and combustion method can be used (Zhao et al., 2014; Wu et al., 2010; Kim et al., 2009). FHA nanopowders were prepared with 0%, 25%, 50%, 75%, and 100% replacement of OH group with F ions through a mechanical alloying method. During the synthesis process, calcium, phosphorous, and fluoride precursors were mixed with high-energy planetary ball mills for 6h at 300 rpm, using eight balls of 20 mm in diameter and a ball-to-powder weight ratio equal to 35:1. The prepared nanopowder Ca10(PO4)6F2 having mean particle size of 35e65 nm fulfilled the requirements of the ASTM 1185-88 standard specification for use as biomaterials (Fathi and Zahrani, 2009; Søballe and Overgaard, 1996). FHA was successfully synthesized by Zhang et al. (Zhang et al., 2005) through a combination of mechanochemical and hydrothermal process. The starting materials used were CaCO3, CaPHO4.2H2O, and CaF2. A comparison was also made between the two milling media, i.e., H2O and acetone for the synthesis of FHA, and it was revealed that apatite phases were formed at varied rates in both media (Zhang et al., 2005). The kinetics of reactions observed in deionized water was faster as compared with kinetics of reactions in acetone. The hydrothermal process induced by the mechanical collision besides the solidstate reaction resulted in fast reactions kinetics in deionized water. It was concluded that during the first phase, starting materials reacted to form a calcium-deficient, poorly crystallized apatite, and during the second stage of the reaction, F ions become completely incorporated into the structure of apatite.

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Handbook of Ionic Substituted Hydroxyapatites

Murugan et al. (Murugan et al., 2002) isolated bioapatite (BAP) by an alkaline hydrothermal process, followed by calcination at 700 C. The BAP was obtained from bovine bone. The fluoridation of BAP was done by using low-temperature method followed by high-temperature method. A stoichiometric amount of hydrofluoric acid was reacted with BAP in a hydrothermal apparatus at 250 C for 30 min in low-temperature method, whereas stoichiometric amounts of BAP and sodium fluoride were ground in an agate mortar and heated at 900 C for 12 h in high-temperature method.

7.4.2

Structure of fluoride-substituted hydroxyapatite

The main components of the apatitic family are hydroxyapatite, chlorapatite, carbonated apatites, and fluorapatite. Fluorapatite, Ca10 (PO4)6F2, is a widespread form of calcium phosphate containing fluoride as one of the main components, which plays an important therapeutic role in the prevention of caries. During the developmental phases of enamel and dentin by substituting OH in the apatite molecule, fluoride fixes calcium, increases the mineral resistance to acidic dissolution, and promotes the process of remineralization (Robinson et al., 2003).

7.4.2.1

Crystallographic study of fluorapatite structure

Fluorapatite ionic crystal structure belongs to the spatial group P63/m (C6h2 in the Sch€ oenflies notation), and its parameters are a ¼ b ¼ 9.462 Å and c ¼ 6.849 Å, a ¼ b ¼ 90 degrees, g ¼ 120 degrees (Hughes et al., 1989). The structure of FA crystal can be explained as hexagonal close-packing of the phosphate ions (where the phosphate ions can be considered as spheres) with channels of octahedral holes through the structure that is parallel to c-axis hexagonally; the “X-Channels”. Ca (1) and Ca (2) are the two types of calcium ions found in the structure. Ca (1) ions occupied two thirds of the channels within the structure, and these ions are linked to other neighboring calcium ions below and above by three shared oxygen atoms. Additionally, Ca (1) ions are also connected by three more distant oxygen atoms. Fluoride ions occupy one-third of the X-channel in the structure, and these fluoride ions are coordinated by three Ca (2) ions. The Ca (2) ions are sevenfold linked by one fluoride ion and six atoms of oxygen as shown in Tables 7.1e7.3. Table 7.1 Positions of nonequivalent atoms of FA (Hughes et al., 1989). Atom

Symmetry site

Crystallographic data

F

6 (C3h)

0

0

1/4

Ca(I)

3 (C3)

1/3

2/3

0.00010

Ca(II)

m (Cs)

0.0071

0.2423

1/4

P

m (Cs)

0.3690

0.3985

1/4

O(I)

m (Cs)

0.4849

0.3273

1/4

O(II)

m (Cs)

0.4667

0.5875

1/4

O(III)

1 (E)

0.2575

0.3421

0.0705

Fluoride-substituted hydroxyapatite

183

Table 7.2 Site symmetry of FA atoms in the space group P63/m, according to the International Tables of Crystallography (Glazer, 1986). Atom

Multiplicity and Wyckoff symbol

Crystallographic data

F

2a

(0; 0; 1/4 ), (0; 0; 3/4)

Ca(I)

4f

(1/3; 2/3; z), (2/3; 1/3; z), (2/3; 1/3; zþ1/2 ), (1/3; 2/3; 1/2 z)

Ca(II)

6h

(x; y; 1/4 ), (1y; xy; 1/4 ), (yx; 1x; 1/4 ) (1x; 1y; 3/4) (y; yx; 3/4), (xy; x; 3/4)

12i

(x; y; z) (1x; 1y; 1z) (1x; 1y; 1/2 þz) (x; y; 1 /2 z) (1-y; xy; z) (y; yx; 1z) (y; yx; 1/2 þz) (1y; xy; 1/2 z) (yx; 1x; z) (xy; x; 1z) (xy; x; 1/2 þz) (yx; 1x; 1/2 z)

P O(I) O(II) O(III)

Table 7.3 Estimation of F-content from Fourier transform infrared spectra in calc-FHA samples (Rintoul et al., 2007). HF/HLa

%FƄb

%Fc

%Fd

Sample

Absorption

OH stretching

OH libration

Found

calc-HAp

e

0

calc-FHAp1

0.881

calc-FHAp2

0

0

10e20

19

18

0.528

20e25

31

26

calc-FHAp3

0.096

50e75

50

58

calc-FHAp4

0.035

50e75

76

72

calc-FHAp5

0.034

50e75

76

73

a

ratio of high frequency (HF) and low frequency (LF) absorption of the OH stretching, % fluorine predicted from the ratio of HF and LF absorption, % fluorine predicted from the spectral signature of the OH liberation bands, d % fluorine determined by fluoride selective electrode. b c

A FA unit cell contains seven nonequivalent atoms: F, Ca (I), Ca (II), P, O (I), O (II), and O (III) as shown in Table 7.1. To show this property, chemical formula of FA can be written as: Ca(I)4Ca(II)6[PO(I)O(II)O(III)2]6 F2, and it takes into account the four nonequivalent ions as tabulated in Table 7.2. The description of atoms is always based on symmetrical site and coordinates of single series. Other positions can be found with different elemental symmetry.

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Handbook of Ionic Substituted Hydroxyapatites

Because of the presence of several ions, FA serves as a suitable host for various ionic substitutions. Despite the differences related to size or valence, FA structure accepts most substituents. Only the substitutions related to Ca2þ ion result in some modifications related to the structure of crystal (when the substitution is smaller than Ca2þ ion). Additionally, because of these modifications, along with the Ca2þ, other sites of unit cell are also affected. By graphical construction of FA unit cell using symmetrical operators of spatial group, P63/m is useful for better understanding of FA crystal structure and modifications induced by various substitutions.

7.4.2.2

Structural analysis of fluoride-substituted hydroxyapatite

Hydroxyapatite constitutes the mineral component of the bones and teeth and has been the subject of interest particularly in the field of biomaterials science (Gunawidjaja et al., 2012). In the recent years, FHA has shown enhanced biocompatibility, cellular proliferation, and bone growth (Harrison et al., 2004). FHA as bioactive ceramic coating combines the effects of bone-regenerative properties of bioactive glasses and fluorapatite (Brauer et al., 2012; Lynch et al., 2012; Pedone et al., 2012; Mneimne et al., 2011). During the preparation of FHA, OH ions are substituted partially by F ions to enhance the stability of materials. Similarly, Fluorapatite is known for lower rate of bioresorption and greater chemical stability than HA (Li et al., 2004). Recently, most studies focused on the different methods of FHA and FA preparation. However, the effects of fluoride contents on the structure have been studied seldomly. For structural analysis, FTIR spectroscopy is a powerful modern tool for studying FHA samples as it provides an accurate information regarding the fluoride ion content. According to literature, samples of FHA prepared by solid precipitation method displayed FTIR spectra identical to hydroxyapatite despite having a high content of fluoride (65%) (Rintoul et al., 2007). However, calcined samples of FHA, heated to 1000 C for 2 h, showed strong correlation between the OH bands and the fluoride content (Wei et al., 2003) as shown in Table 7.3. Detailed spectroscopic analysis showed that, depending on the degree of substitution, the OH libration and stretching vibrations shift and split it into a number of new bands (Freund and Knobel, 1977; Baumer et al., 1985). This has been attributed by the fact that hydroxyl ions orient themselves with neighboring fluoride ions in such a way that hydrogen atoms direct toward the fluoride ions. Thus, the hydrogen bonds formed by fluoride ions with hydroxyl ions are stronger than the hydroxyl ion itself, resulting in chain reversal in the X-channel. This correlation between the fluoride ions content, the position, and intensity of OH vibration bands have been observed using FTIR spectroscopy, thus making it an important tool for the determination of fluoride content in FHAp samples (Baumer et al., 1985). FTIR spectroscopy of FHA synthesized by a cyclic pH method for calcined samples showed a strong correlation between a-axis parameters and content of fluoride ions. However, fluoride ion content was found higher than estimated from the a-axis values for uncalcined samples. The results revealed that for calcined samples, OH band shifts and splits in accordance with FeHO interactions affecting the OH vibration.

Fluoride-substituted hydroxyapatite

185

Solid-state muclear magnetic resonance (NMR), Raman spectroscopy, and XRD have been used extensively in the past to study the structure of hydroxyapatite and substitution in hydroxyapatite by different anions and cations such as Pb2þ, Mg2þ, and F ions (Laurencin et al., 2011; Gunawidjaja et al., 2012; Xu et al., 2010; Chen et al., 2015). Fig. 7.4 showed the XRD pattern of FHA containing different fluoride contents, where typical peaks of apatites were observed at 28.1 , 28.9 , 31.7 , 32.8 , and 34.0 , which are corresponding to (102), (210), (211),(300), (202), (310), (222), (213), and (004) Miller’s planes, respectively (McCubbin et al., 2008). As the level of fluoride ions in fluorohydroxyapatite increased, small peaks became noisy and disappeared in the range of 45e50 degrees, and the apatite crystal structure peaks of (211), (300), and (200) gradually shifted to the right-hand side. These shifts were produced because of the decrease in a-axis length of the hexagonal crystal lattice, where by it was induced by the lower ionic radius of F ions. Literature showed that the H NMR peaks of HA usually appeared at 0.3 and 4.45 ppm corresponding to hydroxyl ions and water molecules, whereas for FHA samples, as the concentration of fluoride ions increased (Fig. 7.5(a)) within the apatite lattice, the chemical shift of OH ions causes the height of peak to decrease because of F ions substitution. This showed that OH ions substitute F ions within the lattice structure. The Ca NMR spectra (Fig. 7.5(b)) of FHA with different levels of fluorine content showed Ca (I) and Ca(II) peaks at 9.6 and 5.6 ppm, respectively. The increased fluorine content had minimal effect on Ca (I) environment, but the chemical environment of Ca (II) changed by decreasing the Ca (II) eO bond length as indicated by Ca (II) peak, which shifted downfield. Thus, the anionic substitution of F ions impart the Ca (II) ions as in Ca(II) polyhedron, Ca(II) ion bonds to six oxygen atoms and one

FHA6 Intensity

FHA5 FHA4 FHA3 FHA2 FHA1 HA 10

20

30 40 2θ (degree)

50

60

Figure 7.4 X-ray diffraction patterns of fluorohydroxyapatites containing different fluorine levels (HA: 0wt%, FHA1: 0.54wt%, FHA2:0.83wt%, FHA3:1.59wt%, FHA4: 1.93wt%, FHA5: 2.2wt%, FHA6: 2.94wt%) (Chen et al., 2015).

186

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(a)

(b)

Structure water .

OH

FHA6 FHA5

FHA6 FHA5

FHA4

FHA4 FHA3

FHA3 FHA2 FHA1 HA 12

10

FHA2 FHA1 HA 8

1

6 4 2 0 –2 H chemical shift (ppm)

–4

–6

(c)

100 80

60

40

43Ca

20 0 –20 –40 –60 –80 –100 chemical shift (ppm)

(d) FHA6

* *

*

*

FHA5

FHA6 FHA5 FHA4 FHA3 FHA2 FHA1 HA

10

Ca(II) Ca(I)

FHA4 FHA3 FHA2 FHA1 8

6

31

4 2 0 –2 P chemical shift (ppm)

–4

–40

–60 19

–80 –100 –120 –140 –160 F chemical shift (ppm)

Figure 7.5 (a) 1H NMR spectra, (b) 43Ca NMR spectra, (c) 31P NMR spectra, and (d) 19F NMR spectra of fluorohydroxyapatites containing different fluorine levels (HA: 0wt%, FHA1: 0.54wt%, FHA2:0.83wt%, FHA3:1.59wt%, FHA4: 1.93wt%, FHA5: 2.2wt%, FHA6: 2.94wt%). The acquisition time for each 1H, 43Ca and 31P spectrum were 5 min, 24 and 1 h, respectively. All 1H and 31P NMR spectra were obtained using a Varian VNMRS 400 MHz solid-state NMR spectrometer and a 6 mm double-resonance MAS probe with a spinning frequency of 8 kHz. All 43Ca spectra were obtained on a 830 MHz solid-state NMR spectrometer using a single-resonance 4 mm MAS probe with a spinning frequency of 10 kHz at room temperature 40 (25 C). Solid-state 19F NMR experiments were performed on a Bruker 600 MHz NMR spectrometer using a spin echo pulse excitation. A p/2 pulse length of 5.57 ms, a recycle delay of 2 s, a spinning rate of 14 kHz, and the echo time set to five rotor period were used for 19F MAS NMR experiments (Chen et al., 2015).

column anion. The P NMR spectra (Fig. 7.5(c)) of FHA exhibited a single well-resolved resonance at 2.8 ppm with various fluorine contents, and this resonance shifted upfield as fluorine content increased from 0 wt.% to 1.5 wt.% and then remained constant at 3.4 ppm from 1.93 wt.% to 2.94 wt.%. The increased incorporation of fluoride ions increased the crystallinity within the HA lattice, which was indicated by progressively narrow signal. The F NMR spectra (Fig. 7.5(d)) exhibited one peak at around 103 ppm and a few sidebands. As the hydroxyl ions in the apatite lattice were replaced by F ions, the height of fluorine signal also increased (Pizzala et al., 2009; Xu et al., 2010; McCubbin et al., 2008; Wilson et al., 2006; Cho et al., 2003).

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Raman spectroscopic analysis of studied samples with different content of fluorine 1 3 revealed strong PO3 4 band at̴960 cm . However, PO4 band associated with the PeO stretch shifts upfield as the concentration of fluorine increased, which is mainly due to PeO bond shortening that occurs progressively. This progressive replacement of OH ions by smaller ionic radius F ions significantly increased the electrostatic forces of attraction between the atoms of oxygen in phosphate tetrahedral, which results in PeO bonds shortening and an increase in vibrational frequency. The increased fluorine content in the samples resulted in intensity decrease of the OeH stretch at about 3580 cm1 (O’Donnell et al., 2009; McElderry et al., 2013; McCubbin et al., 2008; Jay et al., 2012).

7.4.3

Biomedical applications of fluoride-substituted hydroxyapatite

Bone is a highly vascularized, complex structure, which undergoes the process of healing and repairing throughout life (Ratnayake et al., 2017). To develop a biocompatible bone graft material, to restore the structural and functional integrity of the damaged hard tissues, the basic concept of the biological, chemical, and various mechanical processes surrounding the implanted structure is required. This ultimately not only minimizes the future complications such as failure of implanted structure but also improves the quality of life of a patient. Currently, researchers are facing challenges to develop an ideal bone graft material to restore large bony defects due to trauma or disease. In both developed and underdeveloped countries, there is a steady increase in the number of procedures to repair the bony defects in the field of neurosurgery, oral and maxilla facial surgery, and orthopedics (Giannoudis et al., 2005). Hydroxyapatite has been investigated extensively for use in various biomedical applications because of its similarity to the mineral component of the bone, however, due to its inherent brittleness, low fracture toughness, and slow rate of resorption limits its use in the load-bearing areas as a bulk material. Several in vitro and in vivo investigations have been made to determine the ways to modify the critical properties of hydroxyapatite by various levels of anionic and cationic substitutions (Huang et al., 2016; Brown and Constantz, 2017). The use of FHA and FAp has significantly effected the morphology, crystal structure, solubility, and chemical stability of the implanted material (Huang et al., 2016). The potential application of FHA and FAp for various biomedical applications such as bioactive coating on metallic osseous implants, dental restorative material, and a replacement for bony defects will be discussed in this section.

7.4.3.1

Restore tooth enamel defects

Dental enamel, which is the hardest calcified structure in the human body, forms the outer most layer of tooth surface (Taji and Seow, 2010). Enamel lacks the inherent ability to regenerate itself once lost like that of the bone and dentin. Along with dental caries, erosion of tooth surface from intrinsic or extrinsic sources is the most common

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tooth substance loss in the modern world (Feroz et al., 2017). The conventional treatment of these tooth surface lesions mainly involved the filling and drilling technique that results in the loss of the surrounding healthy tooth substance. However, minimal invasion is a key phrase of today’s clinical practice, which emphasized on the prevention, early lesion detection, and the use of minimally invasive treatment options to restore the lesions. To regenerate the enamel layer, many techniques were used recently such as plasma spray (Fogarassy et al., 2005), sol gel (Busch et al., 2001), electrodeposition (Liao et al., 2008), and apatite application (Chen et al., 2005). However, because of the remarkable ability of FA to decrease the acidic dissolution of tooth surface along with the control release of fluoride ions, biocompatibility, and bioactivity, it is now considered as an ideal material for enamel regeneration (Wegehaupt et al., 2009; Nakajo et al., 2009). The crystalline paste of FA containing tetracalcium phosphate (TTCP), dicalcium phosphate anhydrous (DCPA), and ammonium fluoride could be placed into the enamel surface defects directly to repair the lost layer of enamel (Wei et al., 2011). This FA cement contains calcium to phosphorous (Ca/P) atomic molar ratio of 1.67 and calcium to fluoride (Ca/F) ratio of 5. Similarly, this cement exhibits the lower solubility than natural enamel and good cytocompatibility. Additionally, these cements also showed good mechanical properties, hydrophilicity, and antibacterial adhesion. Mechanical properties of FA cement such as values of hardness, modulus of elasticity, and compressive strength were found to be close to that of natural enamel. According to results, the compressive strength, surface microhardness, and elastic modulus values of FA cement were 112 MPa, 3.8 GPa, and 87.1 GPa, respectively, compared with the 117 MPa, 4.1 GPa, and 92.6 GPa values, respectively, of natural enamel. Fluorapatite cement was developed to restore the enamel cavities resembled to natural enamel both structurally and chemically. It also provides considerable protection against caries or erosion as it is less prone to dissolution in acidic environment. Thus, it could be the material of choice to be used to treat the enamel caries instead of conventional materials used for restoration.

7.4.3.2

Dental reinforcing agent

Nanotechnology-based restorative materials have revolutionized the field of modern dentistry, as it is an innovative concept to improve the properties of materials mainly used in dental restorations (Melo et al., 2013). FA exhibits anticariogenic properties by incorporation of fluoride ions and chemical bonding to tooth surfaces. However, their brittleness and poor mechanical properties limit their use as a restorative material in the stress-bearing areas (Lohbauer, 2009). Nano-fluorapatite-incorporated restorative materials significantly modify its properties. Recently, glass ionomer cement (GIC) containing FA particles was formulated to improve the mechanical properties, while preserving the desirable clinical properties (Barandehfard et al., 2016). In other studies, FA nanoparticles were mixed into the powder component of the resinmodified cement. The GIC containing these nanoparticles showed improved

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mechanical properties and has potential to use as a restorative material in high-stress areas such as CLASS I and CLASS II restorations. As the rate of solubility of FA is lower as compared with HA, the FA-containing GICs exhibited greater compressive strength and hardness than HA-added GIC samples (Barandehfard et al., 2016; Moshaverinia et al., 2008; Milne et al., 1997).

7.4.3.3

Implant coatings

The purpose of hydroxyapatite-coated metallic implants is to combine the excellent bioactivity of HA along with the superior mechanical properties of metallic substructure. These coatings also form a protective layer against corrosion in the biological environment. However, when hydroxyapatite is used as a coating in pure form, it mostly results in loosening and failure of implanted structure due to its high rate of bioresorption (Darimont et al., 2004; Baltag et al., 2000). Fluoride ions present in teeth and bone inhabits the process of ionic dissolution and increases the mineral content of the hard tissues. This incorporation of fluoride ions in pure hydroxyapatite coatings not only reduces the biodegradation but also enhances the adhesion of the coatings of the metallic implant surface (Zhang et al., 2006; Lee et al., 2005). In vitro investigations to determine the biological performance of these FHA-coated implants revealed that this incorporation of fluoride not only improves the ionic deposition but also is osteoinductive in nature (Xiao et al., 2018; Wang et al., 2007a). Some researchers also suggested that these coatings also promote the cellular proliferation (Kim et al., 2004a; Bianco et al., 2010; Wang et al., 2007b). Thus, fluoride ion incorporation into the hydroxyapatite lattice structure results in decrease solubility of HA in physiological saline solution and promotes the cellular adhesion over the entire coated surface (Wang et al., 2007b; Huang et al., 2016). Alkaline phosphatase activities, cellular proliferation, and osteocalcin levels enhanced by these fluoridated hydroxyapatite coatings especially when the degree of fluoridation is in the range of 0.8%e1.1% (Wang et al., 2007b). Thus, FHA is considered as an ideal material for bone scaffolding due to their chemical stability, biocompatibility, and low solubility (Bianco et al., 2010). The FHA thermal and chemical stability in biological environments with respect to HA further increased its applications and can also be considered as positive factor, e.g., its use in bone drug delivery systems.

7.5

Concluding remarks

This chapter provides basic understanding regarding the structure, properties, method of preparation, and biomedical applications of FHA with varying degree of fluoride ion substitution. Despite these numbers of in vitro and in vivo analyses on the various aspects of FHA, there is still dearth of information related to the osseointegration, biodegradation, and cytotoxic properties of FHA, and therefore, further investigations should be conducted in this regard to produce better material as bone substitute.

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Kurmaev, E., Matsuya, S., Shin, S., Watanabe, M., Eguchi, R., Ishiwata, Y., Takeuchi, T., Iwami, M., 2002. Observation of fluorapatite formation under hydrolysis of tetracalcium phosphate in the presence of KF by means of soft X-ray emission and absorption spectroscopy. J. Mater. Sci. Mater. Med. 13, 33e36. Lala, S., Ghosh, M., Das, P., Das, D., Kar, T., Pradhan, S., 2016. Magnesium substitution in carbonated hydroxyapatite: structural and microstructural characterization by Rietveld’s refinement. Mater. Chem. Phys. 170, 319e329. Laurencin, D., Almora-Barrios, N., DE Leeuw, N.H., Gervais, C., Bonhomme, C., Mauri, F., Chrzanowski, W., Knowles, J.C., Newport, R.J., Wong, A., 2011. Magnesium incorporation into hydroxyapatite. Biomaterials 32, 1826e1837. Lee, E.-J., Lee, S.-H., Kim, H.-W., Kong, Y.-M., Kim, H.-E., 2005. Fluoridated apatite coatings on titanium obtained by electron-beam deposition. Biomaterials 26, 3843e3851. Li, L.-H., Kong, Y.-M., Kim, H.-W., Kim, Y.-W., Kim, H.-E., Heo, S.-J., Koak, J.-Y., 2004. Improved biological performance of Ti implants due to surface modification by micro-arc oxidation. Biomaterials 25, 2867e2875. Liao, Y.-M., Feng, Z.-D., Li, S.-W., 2008. Preparation and characterization of hydroxyapatite coatings on human enamel by electrodeposition. Thin Solid Films 516, 6145e6150. Lohbauer, U., 2009. Dental glass ionomer cements as permanent filling materials?eProperties, limitations and future trends. Materials 3, 76e96. Lynch, E., Brauer, D.S., Karpukhina, N., Gillam, D.G., Hill, R.G., 2012. Multi-component bioactive glasses of varying fluoride content for treating dentin hypersensitivity. Dent. Mater. 28, 168e178. M€arten, A., Fratzl, P., Paris, O., Zaslansky, P., 2010. On the mineral in collagen of human crown dentine. Biomaterials 31, 5479e5490. Mccubbin, F.M., Mason, H.E., Park, H., Phillips, B.L., Parise, J.B., Nekvasil, H., Lindsley, D.H., 2008. Synthesis and characterization of low-OH fluor-chlorapatite: a single-crystal XRD and NMR spectroscopic study. Am. Mineral. 93, 210e216. Mcelderry, J.-D.P., Zhu, P., Mroue, K.H., Xu, J., Pavan, B., Fang, M., Zhao, G., Mcnerny, E., Kohn, D.H., Franceschi, R.T., 2013. Crystallinity and compositional changes in carbonated apatites: evidence from 31P solid-state Nmr, Raman, and AFM analysis. J. Solid State Chem. 206, 192e198. Melo, M.A., Guedes, S.F., Xu, H.H., Rodrigues, L.K., 2013. Nanotechnology-based restorative materials for dental caries management. Trends Biotechnol. 31, 459e467. Milne, K., Calos, N., O’donnell, J., Kennard, C.L., Vega, S., Marks, D., 1997. Glass-ionomer dental restorative: part I: a structural study. J. Mater. Sci. Mater. Med. 8, 349e356. Mneimne, M., Hill, R.G., Bushby, A.J., Brauer, D.S., 2011. High phosphate content significantly increases apatite formation of fluoride-containing bioactive glasses. Acta Biomater. 7, 1827e1834. Moshaverinia, A., Ansari, S., Moshaverinia, M., Roohpour, N., Darr, J.A., Rehman, I., 2008. Effects of incorporation of hydroxyapatite and fluoroapatite nanobioceramics into conventional glass ionomer cements (GIC). Acta Biomater. 4, 432e440. Murugan, R., Kumar, T.S., Rao, K.P., 2002. Fluorinated bovine hydroxyapatite: preparation and characterization. Mater. Lett. 57, 429e433. Nakajo, K., Imazato, S., Takahashi, Y., Kiba, W., Ebisu, S., Takahashi, N., 2009. Fluoride released from glass-ionomer cement is responsible to inhibit the acid production of cariesrelated oral streptococci. Dent. Mater. 25, 703e708.

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O’hare, P., Meenan, B.J., Burke, G.A., Byrne, G., Dowling, D., Hunt, J.A., 2010. Biological responses to hydroxyapatite surfaces deposited via a co-incident microblasting technique. Biomaterials 31, 515e522. O’donnell, M., Hill, R., Law, R., Fong, S., 2009. Raman spectroscopy, 19F and 31P MAS-NMR of a series of fluorochloroapatites. J. Eur. Ceram. Soc. 29, 377e384. Pasteris, J.D., Wopenka, B., Valsami-Jones, E., 2008. Bone and tooth mineralization: why apatite? Elements 4, 97e104. Pedone, A., Charpentier, T., Menziani, M.C., 2012. The structure of fluoride-containing bioactive glasses: new insights from first-principles calculations and solid state NMR spectroscopy. J. Mater. Chem. 22, 12599e12608. Pizzala, H., Caldarelli, S., Eon, J.-G., Rossi, A.M., Laurencin, D., Smith, M.E., 2009. A solidstate NMR study of lead and vanadium substitution into hydroxyapatite. J. Am. Chem. Soc. 131, 5145e5152. Qing, F., Wang, Z., Hong, Y., Liu, M., Guo, B., Luo, H., Zhang, X., 2012. Selective effects of hydroxyapatite nanoparticles on osteosarcoma cells and osteoblasts. J. Mater. Sci. Mater. Med. 23, 2245e2251. Qu, H., Wei, M., 2005. Synthesis and characterization of fluorine-containing hydroxyapatite by a pH-cycling method. J. Mater. Sci. Mater. Med. 16, 129e133. Rameshbabu, N., Kumar, T.S., Rao, K.P., 2006. Synthesis of nanocrystalline fluorinated hydroxyapatite by microwave processing and its in vitro dissolution study. Bull. Mater. Sci. 29, 611e615. Ratnayake, J.T., Mucalo, M., Dias, G.J., 2017. Substituted hydroxyapatites for bone regeneration: a review of current trends. J. Biomed. Mater. Res. B Appl. Biomater. 105, 1285e1299. Rintoul, L., Wentrup-Byrne, E., Suzuki, S., Grøndahl, L., 2007. FT-IR spectroscopy of fluorosubstituted hydroxyapatite: strengths and limitations. J. Mater. Sci. Mater. Med. 18, 1701e1709. Robinson, C., Kirkham, J., Brookes, S., Bonass, W., Shore, R., 2003. The chemistry of enamel development. Int. J. Dev. Biol. 39, 145e152. Rodriguez-Lorenzo, L., Gross, K., 2003. Encapsulation of apatite particles for improvement in bone regeneration. J. Mater. Sci. Mater. Med. 14, 939e943. Rodriguez-Lorenzo, L., Hart, J., Gross, K., 2003. Influence of fluorine in the synthesis of apatites. Synthesis of solid solutions of hydroxy-fluorapatite. Biomaterials 24, 3777e3785. Rong, Z.-J., Yang, L.-J., Cai, B.-T., Zhu, L.-X., Cao, Y.-L., Wu, G.-F., Zhang, Z.-J., 2016. Porous nano-hydroxyapatite/collagen scaffold containing drug-loaded ADMePLGA microspheres for bone cancer treatment. J. Mater. Sci. Mater. Med. 27, 89. Sadat-Shojai, M., Khorasani, M.-T., Dinpanah-Khoshdargi, E., Jamshidi, A., 2013. Synthesis methods for nanosized hydroxyapatite with diverse structures. Acta Biomaterialia 9, 7591e7621. Skinner, H., 2005. Biominerals. Mineral. Mag. 69, 621e641. Søballe, K., Overgaard, S., 1996. The current status of hydroxyapatite coating of prostheses. J. Bone Joint Surg. Br. 689e691. Sonamuthu, J., Samayanan, S., Jeyaraman, A.R., Murugesan, B., Krishnan, B., Mahalingam, S., 2018. Influences of ionic liquid and temperature on the tailorable surface morphology of Fapatite nanocomposites for enhancing biological abilities for orthopedic implantation. Mater. Sci. Eng. C 84, 99e107.

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Stanic, V., Dimitrijevic, S., Antonovic, D.G., Jokic, B.M., Zec, S.P., Tanaskovic, S.T., Raicevic, S., 2014. Synthesis of fluorine substituted hydroxyapatite nanopowders and application of the central composite design for determination of its antimicrobial effects. Appl. Surf. Sci. 290, 346e352.  Supov a, M., 2015. Substituted hydroxyapatites for biomedical applications: a review. Ceram. Int. 41, 9203e9231. Szczes, A., Hołysz, L., Chibowski, E., 2017. Synthesis of hydroxyapatite for biomedical applications. Adv. Colloid Interface Sci. 249, 321e330. Taji, S., Seow, W., 2010. A literature review of dental erosion in children. Aust. Dent. J. 55, 358e367. Tredwin, C.J., Georgiou, G., Kim, H.-W., Knowles, J.C., 2013a. Hydroxyapatite, fluorhydroxyapatite and fluorapatite produced via the solegel method: bonding to titanium and scanning electron microscopy. Dent. Mater. 29, 521e529. Tredwin, C.J., Young, A.M., Georgiou, G., Shin, S.-H., Kim, H.-W., Knowles, J.C., 2013b. Hydroxyapatite, fluor-hydroxyapatite and fluorapatite produced via the solegel method. Optimisation, characterisation and rheology. Dent. Mater. 29, 166e173. Tredwin, C.J., Young, A.M., Neel, E.A.A., Georgiou, G., Knowles, J.C., 2014. Hydroxyapatite, fluor-hydroxyapatite and fluorapatite produced via the solegel method: dissolution behaviour and biological properties after crystallisation. J. Mater. Sci. Mater. Med. 25, 47e53. Vallet-Regi, M., Gonzalez-Calbet, J.M., 2004. Calcium phosphates as substitution of bone tissues. Prog. Solid State Chem. 32, 1e31. Wang, Y., Liu, L., Guo, S., 2010. Characterization of biodegradable and cytocompatible nanohydroxyapatite/polycaprolactone porous scaffolds in degradation in vitro. Polym. Degrad. Stab. 95, 207e213. Wang, Y., Zhang, S., Zeng, X., Cheng, K., Qian, M., Weng, W., 2007a. In vitro behavior of fluoridated hydroxyapatite coatings in organic-containing simulated body fluid. Mater. Sci. Eng. C 27, 244e250. Wang, Y., Zhang, S., Zeng, X., Ma, L.L., Weng, W., Yan, W., Qian, M., 2007b. Osteoblastic cell response on fluoridated hydroxyapatite coatings. Acta Biomater. 3, 191e197. Wegehaupt, F.J., Solt, B., Sener, B., Wiegand, A., Schmidlin, P.R., Attin, T., 2009. Influence of fluoride concentration and ethanol pre-treatment on the reduction of the acid susceptibility of enamel. Arch. Oral Biol. 54, 823e829. Wei, J., Wang, J., Shan, W., Liu, X., Ma, J., Liu, C., Fang, J., Wei, S., 2011. Development of fluorapatite cement for dental enamel defects repair. J. Mater. Sci. Mater. Med. 22, 1607. Wei, M., Evans, J., Bostrom, T., Grøndahl, L., 2003. Synthesis and characterization of hydroxyapatite, fluoride-substituted hydroxyapatite and fluorapatite. J. Mater. Sci. Mater. Med. 14, 311e320. Wilson, E.E., Awonusi, A., Morris, M.D., Kohn, D.H., Tecklenburg, M.M., Beck, L.W., 2006. Three structural roles for water in bone observed by solid-state NMR. Biophys. J. 90, 3722e3731. Wu, C.-C., Huang, S.-T., Tseng, T.-W., Rao, Q.-L., Lin, H.-C., 2010. FT-IR and XRD investigations on sintered fluoridated hydroxyapatite composites. J. Mol. Struct. 979, 72e76. Xiao, S., Wang, M., Wang, L., Zhu, Y., 2018. Environment-friendly synthesis of trace element Zn, Sr, and F codoping hydroxyapatite with non-cytotoxicity and improved osteoblast proliferation and differentiation. Biol. Trace Elem. Res. 1e14.

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Xu, J., Zhu, P., Gan, Z., Sahar, N., Tecklenburg, M., Morris, M.D., Kohn, D.H., Ramamoorthy, A., 2010. Natural-abundance 43Ca solid-state NMR spectroscopy of bone. J. Am. Chem. Soc. 132, 11504e11509. Zhang, H.G., Zhu, Q., 2005. Surfactant-assisted preparation of fluoride-substituted hydroxyapatite nanorods. Mater. Lett. 59, 3054e3058. Zhang, H.G., Zhu, Q., Xie, Z.H., 2005. Mechanochemicalehydrothermal synthesis and characterization of fluoridated hydroxyapatite. Mater. Res. Bull. 40, 1326e1334. Zhang, S., Xianting, Z., Yongsheng, W., Kui, C., Wenjian, W., 2006. Adhesion strength of solegel derived fluoridated hydroxyapatite coatings. Surf. Coat. Technol. 200, 6350e6354. Zhao, J., Dong, X., Bian, M., Zhao, J., Zhang, Y., Sun, Y., Chen, J., Wang, X., 2014. Solution combustion method for synthesis of nanostructured hydroxyapatite, fluorapatite and chlorapatite. Appl. Surf. Sci. 314, 1026e1033.

Further reading Leroy, N., Bres, E., Jones, D., Downes, S., 2001. Structure and substitutions in fluorapatite. Eur. Cells Mater. 2, 36e48.

Magnesium-substituted hydroxyapatite

8

Ume Omema, Hamad Khalid, Aqif Anwar Chaudhry Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

8.1

Biological importance of the magnesium ion and its relevance to calcium phosphates

The eighth most prevalent ion in the earth crust is magnesium (Gr€ober et al., 2015), and it is considered to be very important among all other bivalent ions (Bhutto et al., 2005; Andres et al., 2017). It is the fourth most common mineral in the human body. Most of the magnesium in the human body is found in the skeleton. It is also found in muscles, cells, enamel, and dentine (Gr€ober et al., 2015; Bhutto et al., 2005; Posner, 1969). Magnesium is involved in almost 300 metabolic reactions, where it is involved in the synthesis of adenosine triphosphate (ATP) (Bhutto et al., 2005). Magnesium ion (Mg2þ) deficiency has an adverse effect on all steps of skeletal metabolism, causing interruption in bone growth and decrease in bone fragility (Andres et al., 2017; Rajendran et al., 2018). It is also involved in the regulation of osteoblast and osteoclast activity and in bone growth. Magnesium plays an important role during the arrangement of cytoskeleton during cell division, which leads to mitotic spindle formation and cytokinesis (Maguire and Cowan, 2002). Lu et al. studied the effects of magnesium on the osteogenesis of human osteoblasts. They concluded that 2 mM magnesium ions showed optimal results, and higher concentrations showed an inhibitory effect on osteoblast activity (Lu et al., 2017). Magnesium is also responsible in controlling the mitosis and DNA synthesis step, which are the rate-limiting steps during the cell cycle (Walker, 1986). Low extracellular magnesium ion concentration can influence the stimulation of osteoblast proliferation and migration (Abed and Moreau, 2009). Deficiency of magnesium in the human body can cause serious health issues that include heart failure, hypertension, muscles diseases, nervous system problems, and atherosclerosis (Mackie, 2003). Hydroxyapatite or HA (Ca10(PO4)6(OH)2) is similar to biological apatite, the mineral component of bone. HA is bioactive and is used to repair bone in the form of blocks, granules, as additives in polymers, as coatings on metallic implants, and in the form of dense bodies. Despite being similar, biological apatite (found in mineralized bone, enamel, and dentine) differs slightly from synthetic HA. It is poorly crystalline and contains different ions (e.g., such as magnesium, zinc, carbonate, and silicate, etc.). These ions are essential for the metabolism, growth of bone, and absorption of nutrients. Foŕ this reason, HA with ionic substitutions is more bioactive.

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00008-2 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Therefore, there is strong rationale behind development of magnesium-substituted hydroxyapatite (Mg-HA) (Farzadi et al., 2014; Ziani et al., 2014). Mg-HA has promising applications in bone repair and regeneration. Incorporation of magnesium ions in the HA lattice leads to enhanced bioactivity. Most importantly, the magnesium ion interacts with integrins of osteoblast cells, which are responsible for cell adhesion and stability (Stipniece et al., 2014). Use of Mg-HA promotes osteoblast activity. The magnesium ion acts similar to growth factors during early stage of osteogenesis, which leads to the formation of bone (Cacciotti et al., 2009). Mg-HA has been found to promote osteogenic differentiation of bone-forming cells (preosteoblasts) and thus improves the osteointegrative potential of the implant material during early implantation times reported by Zhao et al. (Zhao et al., 2013). Gibson et al. also reported a biomimetic HA implant material substituted with both Mg2þ (0.5 wt%) and CO2 3 (1 wt%), which exhibited enhanced bioactivity as compared with HA alone (Gibson and Bonfield, 2002).

8.2

Synthesis of magnesium-substituted hydroxyapatite

The magnesium ion substitutes the calcium ion in the HA lattice. The following equation is based on the assumption that calcium nitrate, magnesium nitrate, and diammonium hydrogen phosphate are used as calcium ion, magnesium ion, and phosphate ion precursors, respectively. Nonetheless, it represents a classical magnesium substitution synthesis reaction (Mehrjoo et al., 2015). 10xCa(NO3)2$4H2O þ xMg(NO3)$6H2Oþ6(NH4)2HPO4þ8NH4OH /Ca10xMgx(PO4)6(OH)2þ20NH4NO3þ(6 þ 4(10x)þ6x)H2O

8.2.1

Coprecipitation method

Coprecipitation is the most commonly used synthesis method for Mg-HA. In this process, two or more compounds are precipitated simultaneously in a solvent because of supersaturation. Coprecipitation involves nucleation, growth, and agglomeration of particles in a solution. Particle size distribution, morphology, and particle sizes are dependent on reaction conditions such as rate of reactant addition and stirring. Rate of nucleation is inversely proportional to the particle size. Usually, the precipitation is done by increasing the pH of reaction mixture up to 10; however, proper aging is required to attain stoichiometric ratio (Cushing et al., 2004). In this method, ammonium dihydrogen phosphate ((NH4)2HPO4, as source of 2þ PO3 4 ), calcium nitrate tetrahydrate (Ca(NO3)2$4H2O, as source of Ca ), and mag2þ nesium nitrate hexahydrate (Mg (NO3)2$6H2O, as source of Mg ) are often used as reactants for synthesis of Mg-HA. The pH of starting solutions and the resulting reaction mixture is kept above 10 by adding ammonia solution (NH4OH). Use of other bases is possible, but this is linked with the possibility of additional ion incorporation in HA (e.g., if KOH is used). Typical steps involved in this method are addition of magnesium ionecontaining calcium precursor solution to the phosphate

Magnesium-substituted hydroxyapatite

199

ionecontaining solution followed by aging of precipitates, filtration, and washing with distilled water (for removal of unreacted species). This is generally followed by drying and heat treatment. To avoid the inclusion of carbonates, some scientists have used N2 atmosphere for their experiments. The magnesium ion substitutes Ca(II) positions in the HA lattice (Ren et al., 2010; Hanafi et al.; Kim et al., 2012; Laurencin et al., 2011; Kannan et al., 2005). Mayer et al. reported on development of magnesium ione and carbonate ionecosubstituted apatites. In their study, refluxing (heating) was used after the precipitation reaction. They reported w1.6 wt% magnesium substitution in the HA lattice (along with 5.4 wt% carbonate ion substitution) (Mayer et al., 1997). Magnesium content in the 0.6e2.4 wt% range has been reported by Cacciotti et al. (Cacciotti et al., 2009). In coprecipitation, often long aging times are required for precipitating calcium phosphates to attain stoichiometry and become HA, which has a Ca:P molar ratio of (or close to in practice) of 1.67. Microwave irradiation has been used to speed up the aging process and thereby reducing overall time taken to synthesize HA (Nazir et al., 2011). Iqbal et. al. used microwave irradiation of preprecipitated suspensions to synthesize HA, Mg-HA, and magnesium-containing calcium phosphates. In doing so, they reduced the traditional aging times from up to a day to 2 h. They also reported on the effects of increasing magnesium concentration on thermal stability (which is reduced), loss of phase purity, and changing particle sizes and surface area (Iqbal et al., 2015).

8.2.2

Solegel method

Solegel synthesis route is a technique in which reactants are converted in a gel form in any stage of reaction scheme. This technique has been used abundantly to synthesize the metal oxides, glass, and ceramics. Researchers have explained the solegel method in a series of steps. In the first step, sol is formed form the stable solution of reactants. In second step, a gel is formed because of the oxide or alcohol bridge network by polycondensation or polyesterification, which increases the viscosity of the solution. The third step is called aging, which brings the solidification of reactants. After aging, drying and dehydration is carried out. In final step, the product is treated with high temperature (>800 C) (Cushing et al., 2004). Researchers have used solegel method to synthesize HA and Mg-HA preparation with 0%e10% of substituted Mg2þ ions, using different chemicals as source of 2þ Ca2þ, PO4 3 , and Mg . This method involves a transition stage in which a colloidal solution is formed, which is converted into a white sol gel. The produced gel is then further processed by aging at room temperature for 24 h, followed by drying to remove solvent and gaseous byproducts, and then subjected to sinter at 500e800 C (Ziani et al., 2014; Gozalian et al., 2011; Abinaya et al., 2014).

8.2.3

Batch and flow hydrothermal synthesis

Usually, in chemical reactions, the temperature of a reactor can be increased maximum up to the boiling point of the solvent, at which (or below) the reaction could be set on

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reflux. However, the boiling point of the solvent can be elevated by increasing the autogenous pressure of reaction in a close container. Materials that are insoluble in a certain solvent may get dissolved and recrystallized using this technique. Hydro is derived from the word hydra which means water and thermal means related to heat (Cushing et al., 2004; Byrappa and Adschiri, 2007). This method typically produces HA with good dispersion and high degree of crystallinity. Mg-HA whiskers with different molar ratios of 0, 0.025, 0.05, and 0.1 were prepared by Liangzhi et al. using hydrothermal synthesis at 180 C for 10 h. Ca(NO3)2.4H2O, Mg(NO3)2, and NH4H2PO4 were the raw materials used in the experiment, and acetamide was used for better dispersion. Hydrothermal treatment was followed by filtration and washing steps with distilled water and ethanol before drying at 200 C for 48 h (Liangzhi et al., 2016). Recently, researchers have reported on the use of continuous hydrothermal flow synthesis (CHFS) route to synthesize HA, substituted HA, and biphasic mixtures. Unlike hydrothermal synthesis in containers (which are batch processes), CHFS relies on flowing solutions reacting at a specific point and undergoing rapid nucleation and limited growth by passing through a heated zone before cooling. The high diffusivity and limited exposure to high temperatures results in nanosized and submicron particles in a very short amount of time eliminating the need for subsequent heat treatment steps (Chaudhry et al., 2006, 2011, 2012a,b). CHFS has also been used to synthesize Mg-HA in a direct continuous manner with tailorable properties (Chaudhry et al., 2008).

8.2.4

Solid-state methods

Solid-state methods utilized to synthesize phase-pure and ion-substituted HA are also referred to as mechanochemical methods. This route is a simple method for powder synthesis, wherein drying, filtration, and postsynthesis heat treatments are often not required. Adzila et al. used Ca(OH)2, (NH4)2HPO4, and Mg(OH)2 as precursors for synthesis of Mg-HA with an intended molar ratio of 1.67 (Ca þ Mg)/P. These reagents were mixed in a ball mill for 15 h. Incorporation of the magnesium ion in HA lattice led to synthesis of calcium-deficient HA (Adzila et al., 2013). Imrie et al. reported the synthesis of calcium phosphate cosubstituted with strontium and magnesium, wherein reagents were ground using a mortar and pestle followed by pressing into disc form. These discs were then heat-treated (held at 1100 C for 16 h) and eventually ground into fine powders. They also reported on the reduced stability of the HA phase on incorporation of magnesium (it stabilized the b-tricalcium phosphate [b-TCP] phase) (Imrie et al., 2013). Imrie et al. used CaCO3, NH4H2PO4, and Mg(NO3)2.6H2O to synthesize magnesium-containing biphasic calcium phosphates (containing apatitic and b-TCP phases). Precursors were mixed in an agate mortar. The mixture was then sintered at 1000 C for 10 h (Webler et al., 2015). Suchanek et al. have reported on the use of mechanochemicalehydrothermal method that combines the use of wet-chemical synthesis and use of ball milling (Suchanek et al., 2004a,b). This involves a precipitation reaction followed by ball milling to produce Mg-HA powders. Phase-pure Mg-HA containing 0.24e28.4 wt% of Mg can be synthesized by using this method (Suchanek et al., 2004a). Table 8.1.

Magnesium-substituted hydroxyapatite

201

Table 8.1 Mg-HA synthesis methods.

No.

Synthesis method

1

Coprecipitation

(Ren et al., 2010; Hanafi et al.; Kim et al., 2012; Laurencin et al., 2011; Kannan et al., 2005; Mayer et al., 1997)

2

Solegel method

(Ziani et al., 2014; Gozalian et al., 2011; Abinaya et al., 2014)

3

Hydrothermal synthesis

(Liangzhi et al., 2016; Chaudhry et al., 2006, 2008, 2011, 2012a, 2012b; Suchanek et al., 2004a, 2004b)

4

Solid-state methods

(Adzila et al., 2013; Imrie et al., 2013; Webler et al., 2015)

8.3

References

Magnesium-substituted hydroxyapatite coatings

HA is brittle and mechanically weak, which limits its application in load-bearing biological applications. This shortcoming can be overcome by using metallic implants and coating them with HA. This provides a bioactive surface and changes the mode of fixation from mechanical to biological as HA bonds to surrounding tissue (Miranda et al., 2019; Ibrahim et al., 2017). Despite significant interest in this area, there is significant space for innovations that address quick bone healing, strong bone/implant interfaces, and improved corrosion resistance. Titanium (Ti) and its alloys are one of the most extensively used metallic implants for biomedical applications in orthopedics and dentistry because of their exceptional mechanical properties and corrosion resistance (Healy and Ducheyne, 1992). However, the problem with metallic implants is the formation of a fibrous capsule around the defect area, and long-term persistence could lead to failure and/or complications (also attributed to the resulting stress shielding). It is also possible that the metallic surfaces change in vivo and this can affect biocompatibility of the implant negatively. Therefore, to address these shortcomings, it is desirable that the surface be modified/improved to be more bioactive (Rau et al., 2013; Lacefield, 1999). Cai et al. used solegel method to coat a titanium substrate with magnesium- and fluorine-substituted HA. It was observed that the magnesium ions improved bioactivity and promoted bone formation. However, increased concentration of magnesium stabilized the b-TCP phase (Cai et al., 2009). O’Neill et al. used Coblast technique to coat Mg-HA on Ti substrate. In this technique, the coating material was mixed with an abrasive material and then bombarded at the metal surface (O’Neill et al., 2009). Jiao and Wang coated titanium substrate with Mg-HA (0e10 mol% magnesium as a substituent of calcium) using electrodeposition (Jiao and Wang, 2009). Yan et al. deposited a 21.2  1.7 mm thick Mg-HA coating onto pure titanium with intermediate phase of TiO2 nanotubes using electrochemical

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Handbook of Ionic Substituted Hydroxyapatites

deposition method. In vitro studies showed that Mg-HA coating was biocompatible and without any adverse effects. Mg-HA coatings showed improved ability of apatite formation as well as higher corrosion resistance than HA alone (Yajing et al., 2014). Electrochemically deposited Mg-HA coatings on pure titanium surfaces were reported to promote osteogenic differentiation of MC3T3-E1 (preosteoblasts) in vitro. Mg-HAecoated substrates improved osseointegration of the implant during early stages of bone repair process as compared with implants coated with pure HA (Zhao et al., 2013). In another study conducted by Mr oz et al., Mg-HA was coated on porous Ti implants surface by pulsed laser deposition (PLD) method. The coated titanium implants were then introduced into a rabbit femoral defect model and analyzed by microcomputed tomography, which showed considerable enhancement in bone volume for implants coated with Mg-doped HA compared with uncoated implants (Mr oz et al., 2015). Liquid precursor plasma spraying technique can also be used to coat Mg-HA on a titanium surface. Huang et al. used this technique for Mg-HA coating on metallic surface (titanium sheets). It was observed that crystal structure of Mg-HA was unchanged after coating (Huang et al., 2011). Onder et al. used cathodic arc physical vapor deposition technique to deposit Mg-HA on a titanium surface (Onder et al., 2013). Sharifnabi et al. used solegel method to coat Mg-HA on medical grade 316L stainless steel. Magnesium was successfully incorporated in HA coating on metallic surface and crystallinity was determined to be about 70%. The coatings were observed to be smooth and without cracks. The coated samples released much less ions in physiological media, which shows that these coatings can potentially improve the biocompatibility and anticorrosion properties (Sharifnabi et al., 2014). Loszach et al. used an induction suspension plasma spray process to deposit Mg-HA on a titanium alloy. They used a solegel method to produce multiionesubstituted HA (magnesium was one of them). The chemical composition of this coating was very close to that of bone (Loszach and Gitzhofer, 2015). Gopi et al. used a different approach. They coated surgical grade stainless steel (316L SS) first with poly(3,4-ethylenedioxythiophene) using electropolymerization followed by electrodeposition of Mg-HA. These bilayercoated samples were bioactivedmaking them potential candidates for use as implants (Gopi et al., 2014). Bao et al. deposited Mg-HA on magnesium alloy using a sol-gel method. The coatings improved corrosion resistance of the magnesium substrate, which was determined by hydrogen evolution testing in Hanks’s solution (Bao et al., 2015). Polycaprolactone (PCL) is another biocompatible and biodegradable polyester that is used to fabricate 3D porous scaffolds with precise internal architecture and accurate distinct dimensions using rapid prototyping methods. To make PCL scaffolds suitable for orthopedic applications, Vandrovcova et al. deposited pure calcium phosphate and calcium phosphate substituted with 0.6% w/w magnesium on PCL using PLD. In vitro studies on osteoblast-like Saos-2 cells showed that after 7 days, magnesium-containing coatings amplified the gene expression of alkaline phosphatase activity and hence facilitated the osteoblastic differentiation process (Vandrovcova et al., 2015).

Magnesium-substituted hydroxyapatite

8.4

203

Characterization of magnesium-substituted hydroxyapatite

8.4.1 8.4.1.1

Electron microscopy Transmission electron microscopy

Gayathri et al. used transmission electron microscope (TEM) to determine the morphology and particle size of dried powder of pure HA and Mg-HA. Fig. 8.1 represents the TEM images and selected area electron diffraction (SAED) patterns. TEM images showed needle-like morphology of 10e20 nm width and 50e100 nm length. SAED patterns showed the powders have grains with polycrystallinity (Cacciotti et al., 2009).

8.4.1.2

Scanning electron microscopy

Scanning electron microscopy (SEM) is useful in observing sintered surfaces of calcium phosphates. Cacciotti et al. synthesized Mg-HA nanopowders using varying concentration of magnesium using the coprecipitation method (0.6e2.4 wt% magnesium content). The powder of Mg-HA was sintered at 1250 C for 1 h, which resulted in the densification of powders. SEM images of fractured surfaces of sintered HA and Mg-HA are shown in Fig. 8.2. The images show that for phase-pure HA lesser

HAp

0.5 Mg-HAp

200 nm

100 nm

0.25 Mg-HAp

200 nm

1.0 Mg-HAp

200 nm

Figure 8.1 Transmission electron microscope (TEM) and selected area electron diffraction (SAED) images of pure hydroxyapatite (HA) and magnesium-substituted HA (Mg-HA) (Cacciotti et al., 2009). Copyright 2009 reproduced with permission from Elsevier.

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Handbook of Ionic Substituted Hydroxyapatites

HAp

20 µm

HAp

5 µm

0.25 Mg-HAp

20 µm

0.25 Mg-HAp

5 µm

0.5 Mg-HAp

20 µm

0.5 Mg-HAp

5 µm

1.0 Mg-HAp

20 µm

1.0 Mg-HAp

5 µm

Figure 8.2 Scanning electron microscopy images of fractured surfaces of sintered hydroxyapatite (HA) and magnesium-substituted HA (Mg-HA) (with varying concentration of HA) (Cacciotti et al., 2009). Copyright 2009 reproduced with permission from Elsevier.

Magnesium-substituted hydroxyapatite

205

porosity is observed as compared to magnesium containing samples. The images presented suggest that porosity increases with magnesium content. The particles shown in TEM images shown in Fig. 8.1 were needle-like and on sintering at 1250 C for 1 h they coalesced to form polygonal grains (Cacciotti et al., 2009). SEM has also been used to assess coatings and their physical attributes. Onder et al. studied the deposition of Mg-HA (using simulated body fluids) on (Ti,Mg)N and TiN coatings (on the titanium surface). Using SEM, they were able to assess the apatitic deposition (and related features) on the coated surfaces.

8.4.2

Surface area

Batra et al. reported that Mg-HA showed a significantly higher BrunauereEmmette Teller (BET) surface area than HA. It was also observed that heat treatment decreased surface area. Crystallite sizes of 24 and 18 nm were determined (using BET theory) for HA and Mg-HA, respectively (Batra et al., 2013). Iqbal et al. reported that for increasing magnesium content (in calcium phosphates), the surface area initially increased but later decreased (possibly because of precipitation of nonapatitic phases). Surface areas of all samples decreased after heat treatment. The variation in surface area as a function of magnesium content is show in Fig. 8.3 (Iqbal et al., 2015). 160

As precipitated Heat treated

140

Surface area (m2/g)

120 100 80 60 40 20

0 0.1

0.5

1

1.5

2

4

5

10

16

22

29

Weight % magnesium (theoretical)

Figure 8.3 Surface area of magnesium-containing calcium phosphates in as-precipitated and heat-treated forms (at 900 C for 1 h) as a function of in-solution magnesium content (theoretical) (Iqbal et al., 2015). Copyright 2015 reproduced with permission from Elsevier.

206

8.4.3

Handbook of Ionic Substituted Hydroxyapatites

X-ray diffraction

X-ray diffraction (XRD) allows determination of phase purity and crystallinity of calcium phosphate materials. Advanced techniques linked to XRD such as Reitveld refinement facilitate identification of crystal structures, atomic positions, substituent contents, lattice parameters, etc. Iqbal et al. developed a range of magnesium-substituted calcium phosphates using the formula Ca10xMgx(PO4)6(OH)2 for x ranging from 0.04 to 8 (i.e., 0.1 to 22 wt%) by using microwave irradiation of preprecipitated suspensions. In XRD patterns for as-precipitated samples, Mg-HA with low crystallinity was observed for magnesium contents of up to 2 wt%. A further increase in magnesium content resulted in poorly resolved peaks possibly due to much decreased crystallinity and possible stabilization of nonapatitic phases. After heat treatment at 900 C for 1 h, it was observed that peaks were better resolved owing to improved crystallinity. For magnesium contents of 0.5e2 wt% (in HA, based on theoretical calculations), biphasic mixtures of magnesium-stabilized b-TCP and Mg-HA were observed after heat treatment. This highlights the decrease in thermal stability of the HA lattice. Further increase led to phase-pure magnesium stabilized b-TCP phase. For samples with more than 10 wt% or higher magnesium content, other magnesium-rich phases such as stanfieldite and farringtonite were observed (after heat treatment). Figs. 8.4 and 8.5 show XRD patterns of samples before and after heat treatment (Iqbal et al., 2015). At lower levels of magnesium, Chaudhry et al. reported CHFS of Mg-HA (0.7 wt% magnesium). An increase in magnesium content led to precipitation of biphasic mixtures of Mg-HA and Mg-whitlockite identified using XRD. Further increase in magnesium content led to phase-pure crystalline Mg-whitlockite precipitation. This highlighted that there is a threshold for magnesium incorporation into HA. Furthermore, using Rietveld refinement the authors reported an overall decrease in the unit cell volume. They attributed this decrease to the calcium ions being replaced by the smaller magnesium ions in the crystal structure (Chaudhry et al., 2008).

8.4.4

Fourier-transform infrared spectroscopy

FTIR spectroscopy allows identification of PeO, CeO, OeH, OePeO, and OeCeO linkages in calcium phosphates. Incorporation of substituting ions in the HA lattice often leads to phase stabilization of other phases, e.g., magnesium contents above a certain threshold lead to stabilization of the b-TCP phase. These changes and affects can be well detected and understood using FTIR spectra. Cacciotti et al. reported on the synthesis and property determination of Mg-HA. Typical HA and Mg-HA spectra from their report are shown in Fig. 8.6. Fig. 8.1(a) shows FTIR spectra of phase-pure HA and Mg-HA (2.4 wt% magnesium) in as dried form. Fig. 8.6(b) shows FTIR spectra of samples containing 0.6 wt%, 1.2 wt%, and 2.4 wt%, respectively, (after heat treatment at 1100 C). Bands in the 1410e1460 cm1 range attributed to carbonate groups (having substituted phosphate groups) are visible in Fig. 8.6(a). Peaks/bands at 361 and 3570 cm1 were attributed to OH groups in HA. Decreased intensity of peaks attributed to OH vibrations (at 360 and 3570 cm1) and broadening of peaks attributed to

Magnesium-substituted hydroxyapatite

207

31.9° – main HA peak (a) (b) (c)

Intensity

(d) (e) (f) (g) (h) (i) (j) (k) (l) 20

25

30

35 2-theta

40

45

50

Figure 8.4 X-ray diffraction patterns of as-precipitated samples containing (a) no magnesium (hence only HA), (b) 0.1 wt% magnesium, (c) 0.5 wt% magnesium, (d) 1.0 wt% magnesium, (e) 1.5 wt% magnesium, (f) 2.0 wt% magnesium, (g) 4 wt% magnesium, (h) 5 wt% magnesium, (i) 10 wt% magnesium (g) 16 wt% magnesium, (k) 22 wt% magnesium, and (l) no calcium (referred to as a magnesium phosphate). The magnesium contents refer to in-solution amounts based on theoretical calculations (Iqbal et al., 2015). Copyright 2015 reproduced with permission from Elsevier.

phosphate groups in HA were attributed to the presence of magnesium. This was possibly due to increase lattice distortion as a result of presence of HPO42 (which increases with magnesium content). In Fig. 8.6(b), it was observed that peaks/bands corresponding to carbonate groups disappear. Bands attributed to water (at 1636 and 3430 cm1) also decreased in intensity. For all magnesium-containing samples, the intensity of peaks linked to OH vibrations (632 and 3572 cm1) decreased with increasing magnesium content. Broadening of peaks/bands associated to phosphate groups (700e1700 cm1 range) was also observed with increasing magnesium content. The authors reported extra peaks in the 942e1120 cm1 range attributed to phosphate groups of TCP. This was due to decrease in thermal stability as a result of magnesium incorporation, which resulted in decomposition of Mg-HA into b-TCP (as a second phase in the resulting biphasic mixture) (Cacciotti et al., 2009). Iqbal et al. in their report on synthesis of magnesium-substituted calcium phosphates using microwave irradiation of preprecipitated suspensions used FTIR spectroscopy to determine that for theoretical substitution levels of up to 2 wt% the resulting phases were apatitic in nature. They attributed the peak at 632 cm1 to the librational mode of the hydroxyl group and linked it to the crystallinity of resulting calcium phosphates (Iqbal et al., 2015).

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Hydroxyapatite Mg-BTCP Stanfieldite Farringtonite

(a) (b) (c)

Intensity

(d) (e) (f) (g) (h) (i) (j) (k) (l) 20

25

30

35 2-theta

40

45

50

Figure 8.5 X-ray diffraction patterns of samples heat-treated at 900 C for 1 h containing (a) no magnesium (hence only HA), (b) 0.1 wt% magnesium, (c) 0.5 wt% magnesium, (d) 1.0 wt% magnesium, (e) 1.5 wt% magnesium, (f) 2.0 wt% magnesium, (g) 4 wt% magnesium, (h) 5 wt% magnesium, (i) 10 wt% magnesium, (g) 16 wt% magnesium, (k) 22 wt% magnesium, and (l) no calcium (referred to as a magnesium phosphate). The magnesium contents refer to in-solution amounts based on theoretical calculations (Iqbal et al., 2015). Copyright 2015 reproduced with permission from Elsevier.

8.4.5

Raman spectroscopy

Raman spectroscopy augments FTIR spectroscopy data, and it has been used extensively to study phase-pure and ion-substituted HA and calcium phosphates. Chaudhry et al. used this technique to evaluate a range of their flow-synthesized, magnesiumsubstituted HA and calcium phosphates. For theoretical in-solution contents ranging between 0.5 and 2 wt%, the Raman spectra were like those observed for HA. Peaks corresponding to asymmetric stretching (n3) of PeO bonds in phosphate groups were observed around 1083 and 1054 cm1. The peak corresponding to symmetric stretching (n1) of PeO was also observed at 963 cm1. Peaks due to bending vibrations (n4 and n2) were observed around 615 and 437 cm1. The authors attributed a peak at 633 cm1 to bending mode (of OePeO) linked to phosphate groups in a magnesium-stabilized whitlockite phase. Presence of HPO 4 groups was also confirmed, and this was linked to the presence of magnesium-stabilized whitlockite phase. Overall, Raman spectroscopy supported findings from FTIR spectroscopy, and XRD that increases in magnesium content led to stabilization of and hence direct precipitation of magnesium stabilized whitlockite phases above a certain threshold. The corresponding Raman spectra are shown in Fig. 8.7 (Chaudhry et al., 2008).

Magnesium-substituted hydroxyapatite

209

(a) 80

Transmittance (%)

70

HAp 632

60 1.0Mg-HAp 50

3572

40 30

500

1000 Wavelength

3500

4000

(cm–1)

(b) 90 0.25Mg-HAp

Transmittance (%)

80

0.5Mg-HAp

70

1.0Mg-HAp

60

3572

50 632 40 30 500

1000

3500

4000

Wavelength (cm–1)

Figure 8.6 Fourier-transform infrared spectra of (a) hydroxyapatite (HA) and magnesiumsubstituted HA (Mg-HA) (2.4 wt% Mg) as dried powders and (b) Mg-HA with varying concentrations of magnesium (0.6e2.4 wt%) heat-treated at 1100 C (Cacciotti et al., 2009). Copyright 2009 reproduced with permission from Elsevier Ltd.

8.4.6

Thermogravimetric analysis

Thermogravimetric analysis (TGA) is often carried out from room temperature to an elevated temperature to assess thermal stability of phase-pure and Mg-HA. Bauer et al. developed Mg-HA scaffolds from cuttlefish using a hydrothermal process. They carried out TGA with a heating rate of 10 C/min from 40 to 1200 C. There was a marked increase in amount of total weight loss with an increase in magnesium content (from 0% to 10%, wherein 10% magnesium refers to 10% of calcium ions

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P-O, PO43– O-P-O, PO43– O-P-O, HPO42– (a) (b) (c)

Intensity

(d) (e) (f) (g) (h) (i) (j) (k)

1250

950

Raman shift

650

350

Figure 8.7 Raman spectra corresponding to samples containing magnesium contents of (a) 0.5 wt%, (b) 1 wt%, (c) 1.5 wt%, (d) 2 wt%, (e) 4 wt%, (f) 6 wt%, (g) 8 wt%, (h) 10 wt%, (i) 12 wt%, (j) 14 wt%, and (k) no magnesium (referring to magnesium phosphate phase) (Chaudhry et al., 2008). Copyright 2008 reproduced with the permission from Royal Society of Chemistry.

being substituted). A sudden weight loss step becomes visible between 700 and 800 C, which is possibly due to decreased thermal stability of magnesium HA resulting in decomposition into Mg-HA and magnesium-containing b-TCP (Bauer et al., 2018).

8.5

Assessing biological response to magnesiumsubstituted hydroxyapatite

Bioactive materials are often assessed for their biological response. In vitro assessments include checking their antibacterial nature and response to cells. Similarly, in vivo models focus on assessment in animals. Following thorough research and development, ethical and regulatory compliance/procedure materials are often used in clinical trials (in human patients).

Magnesium-substituted hydroxyapatite

8.5.1 8.5.1.1

211

In vitro analysis Cellular response

Cell-based biological response studies are often carried out to evaluate how magnesium effects osteogenic potential, cell viability, and cytotoxicity. Y. Yajing et al. investigated the cellular response of mouse calvarial cells (MC3T3-E1) to HA and Mg-HA coatings. Their results showed that incorporation of magnesium into HA did not have any negative effect. It was proposed that Mg-HA surfaces can promote in vitro osteogenic differentiation of preosteoblasts (Yajing et al., 2014). Zhao et al. used MC3T3-E1 preosteoblasts to compare pure HA and Mg-HA coatings. Growth of cells, alkaline phosphatase activity, and secretion of osteocalcin were determined at different periods of time. The authors concluded that their Mg-HA coating surface promoted in vitro osteogenic differentiation of the cells used (Zhao et al., 2013). Serre et al. used two kinds of sponges (made up of collagen and apatites). One composition included an apatite additive that contained a low level of magnesium, whereas the other contained a higher amount of magnesium. It was observed that for sponge containing apatite with higher levels of magnesium, a toxic effect was observed, whereas the sponge with apatite containing lower amount of magnesium decreased the osteoconductive properties of the sponge. They did mention that at physiological levels magnesium ions promote mineralization of bone (Serre et al., 1998). Cellular response of Mg-HA in powdered form has also been evaluation by researchers. In their report on linking magnesium substitution to properties of HA, Mehrjoo et al. used human osteoblast cells (MG-63). Responses to the powders were assessed using 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) assay and Alkaline Phosphattase (ALP) activity was also determined. Using quantitative Real Time Polymerase Chain Reaction (qRT-PCR), the researchers also evaluated expression of genes. Within the scope of the study, higher levels of magnesium led to higher cell proliferation, ALP activity, and gene expression, enabling the authors to confirm the positive role of magnesium ions in bone regeneration applications (Mehrjoo et al., 2015). Determination of interaction with cells has been carried out on different forms of materials containing Mg-HA. These are either powders, additives in polymeric scaffolds, and/or coatings on metallic implants. These highlight different potential applications based on the positive influence of magnesium ion incorporation (also referred to as substitution or doping interchangeably in literature) into the HA lattice.

8.5.1.2

Antibacterial response

The key focus of apatitic materials (i.e., HA and substituted HAs) is bone repair and regeneration. Their application therefore includes sites that are accessed surgically possible due to trauma, disease, or for esthetic purposes. It is therefore desirable that bone regenerative materials offer dual functionality, i.e., they also offer antibacterial properties. This can help in mitigating the risks of infection during and after surgery.

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The main three hypothetical mechanisms proposed in literature in relation to the antibacterial activity of ions are listed below (Kolmas et al., 2014; Simchi et al., 2010; Dastjerdi and Montazer, 2010; Díaz-Visurraga et al., 2011; Hajipour et al., 2012); • • •

Permeation of the ions in consideration into the bacterial cells and disruption in the formation of ATP leading to disruption of the DNA replication process. Addition of the substituted ions in the bacterial cell membrane changes the structure and permeability of the cell membrane, which in turn effects the nutrient transport through the cell membrane finally resulting in damage to the cell membrane and cell death. Involvement of reactive oxygen species due to presence of ions, which react with components of the cell membrane and cell wall causing permanent change in their structure and cell death.

Sivaraj et al. deposited composite coatings on 316L stainless steel substrates, wherein they utilized multiwalled carbon nanotubes (MWCNT) with HA and Mg-HA, respectively. Based on bigger zones of inhibition observed, they concluded Mg-HA/MWCNT had higher antibacterial activity against Escherichia coli bacteria (Sivaraj and Vijayalakshmi, 2018). Gayathri et al. reported in his study the antibacterial activity of different concentrations of Mg-HA against Staphylococcus aureus and E. coli by using the agar diffusion procedure. According to their results, no zone of inhibition was observed against S. aureus; however, in case of E. coli, antibacterial activity was observed for MgHA (Gayathri et al., 2018). Andres et al. reported assessed antibacterial response of Mg-HA based on S. aureus, Pseudomonas aeruginosa, and E. coli strains. They attributed the antibacterial effect of Mg-HA to the surface topography of the material used and its electroactive nature. They reported optimum antibacterial response for 2.23 wt% magnesium content. On the basis of their findings, Andres et al. proposed Mg-HA as a suitable candidate material for implants and materials focusing on prevention of bone-related infections and promotion of wound healing (Andres, 2018).

8.5.2

In vivo analysis (animal model)

Magnesium plays a very important role in the in vivo bone formation and calcification of the bone (Oguchi et al., 1995). Wuthier et al. reported in his study the highest amount of magnesium ion was present in proliferating cartilage of chicks (Wuthier, 1971). In a report by Jones et al., mice were given magnesium supplemented feed. Results showed increase in calcification with an increased in magnesium content, while the rats that were magnesium deficient showed abnormal mineralization of bone (Jones et al., 1980). Brautbar et al. reported in his study that magnesium deficiency caused decalcification of bone in rats and mouse pups (Brautbar and Gruber, 1986). Animal models are therefore often used to assess the bone regenerative ability of bioactive materials. Because of the importance of the magnesium ion to bone development and health, some studies have been carried out using Mg-Ha in the animal models as well.

Magnesium-substituted hydroxyapatite

213

Landi et al. evaluated bone regenerative ability of Mg-HA (containing 5.7 mol% magnesium). Mg-HA was developed in the form of granules (400e600 microns in size). These granules were used to fill defects created in the femurs of New Zealand white rabbits. Based on comparison with stoichiometric HA (which did not contain magnesium), Mg-HA was found to be more osteoconductive (Landi et al., 2008). Ahmadzadeh et al. developed polyvinyl alcohol compositeecontained, carbonatesubstituted HA and zincemagnesium-substituted HA nanoparticles. They evaluated these materials in the tibia of rabbits and observed no inflammation or infection after 4 weeks of implantation (Ahmadzadeh et al., 2016).

References Abed, E., Moreau, R., 2009. Importance of melastatin-like transient receptor potential 7 and magnesium in the stimulation of osteoblast proliferation and migration by platelet-derived growth factor. Am. J. Physiol. Cell Physiol. 297 (2), C360eC368. Abinaya, R., et al., 2014. Comparative study on the Mg doped hydroxyapatite through sol-gel and hydrothermal techniques. Int. J. Innov. Res. Sci. Eng. 1, 1e6. Adzila, S., et al., 2013. Mechanochemical synthesis of magnesium doped hydroxyapatite: powder characterization. In: Applied Mechanics and Materials. Trans Tech Publ. Ahmadzadeh, E., et al., 2016. Osteoconductive composite graft based on bacterial synthesized hydroxyapatite nanoparticles doped with different ions: from synthesis to in vivo studies. Nanomedicine 12 (5), 1387e1395. Andres, N.C., 2018. Electroactive Mg(2þ)-hydroxyapatite nanostructured networks against drug-resistant bone infection strains, 10 (23), 19534e19544. Andres, N.C., et al., 2017. Manipulation of Mg2þeCa2þ switch on the development of bone mimetic hydroxyapatite. ACS Appl. Mater. Interfaces 9 (18), 15698e15710. Bao, Q.H., et al., 2015. Dip coated magnesium-substituted hydroxyapatite coatings on magnesium alloy for biomedical applications. J. Biomim. Biomater. Biomed. Eng. 25, 83e89. Batra, U., Kapoor, S., Sharma, S., 2013. Influence of magnesium ion substitution on structural and thermal behavior of nanodimensional hydroxyapatite. J. Mater. Eng. Perform. 22 (6), 1798e1806. Bauer, L., Ivankovic, M., Ivankovic, H., 2018. Magnesium substituted hydroxyapatite scaffolds hydrothermally synthesized from cuttlefish bone. In: MATRIB. Bhutto, A., et al., 2005. Magnesium and its Essential Role in Health. JLUMHS, pp. 33e34. Brautbar, N., Gruber, H.E., 1986. Magnesium and bone disease. Nephron 44 (1), 1e7. Byrappa, K., Adschiri, T., 2007. Hydrothermal technology for nanotechnology. Prog. Cryst. Growth Charact. Mater. 53 (2), 117e166. Cacciotti, I., et al., 2009. Mg-substituted hydroxyapatite nanopowders: synthesis, thermal stability and sintering behaviour. J. Eur. Ceram. Soc. 29 (14), 2969e2978. Cai, Y., et al., 2009. Improvement of bioactivity with magnesium and fluorine ions incorporated hydroxyapatite coatings via solegel deposition on Ti6Al4V alloys. Thin Solid Films 517 (17), 5347e5351. Chaudhry, A., et al., 2006. Instant nano-hydroxyapatite. In: A Continuous and Rapid Hydrothermal Synthesis, vols. 2286e2288, pp. 2286e2288.

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Chaudhry, A.A., et al., 2008. Synthesis and characterisation of magnesium substituted calcium phosphate bioceramic nanoparticles made via continuous hydrothermal flow synthesis. J. Mater. Chem. 18 (48), 5900e5908. Chaudhry, A., et al., 2011. High-Strength Nanograined and Translucent Hydroxyapatite Monoliths via Continuous Hydrothermal Synthesis and Optimized Spark Plasma Sintering, vol. 7, pp. 791e799. Chaudhry, A., et al., 2012. Rapid Hydrothermal Flow Synthesis and Characterisation of Carbonate- and Silicate-Substituted Calcium Phosphates, vol. 28. Chaudhry, A.A., et al., 2012. Phase stability and rapid consolidation of hydroxyapatite-zirconia nano-coprecipitates made using continuous hydrothermal flow synthesis. J. Biomater. Appl. 27 (1), 79e90. Cushing, B.L., Kolesnichenko, V.L., O’Connor, C.J., 2004. Recent advances in the liquid-phase syntheses of inorganic nanoparticles. Chem. Rev. 104 (9), 3893e3946. Dastjerdi, R., Montazer, M., 2010. A Review on the Application of Inorganic Nano-Structured Materials in the Modification of Textiles: Focus on Anti-microbial Properties, vol. 79, pp. 5e18. Díaz-Visurraga, J., et al., 2011. Metal Nanostructures as Antibacterial Agents. Farzadi, A., et al., 2014. Magnesium incorporated hydroxyapatite: synthesis and structural properties characterization. Ceram. Int. 40 (4), 6021e6029. Gayathri, B., et al., 2018. Magnesium incorporated hydroxyapatite nanoparticles: preparation, characterization, antibacterial and larvicidal activity. Arab. J. Chem. 11 (5), 645e654. Gibson, I.R., Bonfield, W., 2002. Preparation and characterization of magnesium/carbonate cosubstituted hydroxyapatites. J. Mater. Sci. Mater. Med. 13 (7), 685e693. Gopi, D., et al., 2014. Development of strontium and magnesium substituted porous hydroxyapatite/poly(3,4-ethylenedioxythiophene) coating on surgical grade stainless steel and its bioactivity on osteoblast cells. Colloids Surfaces B Biointerfaces 114, 234e240. Gozalian, A., et al., 2011. Synthesis and thermal behavior of Mg-doped calcium phosphate nanopowders via the sol gel method. Sci. Iran. 18 (6), 1614e1622. Gr€ ober, U., Schmidt, J., Kisters, K., 2015. Magnesium in prevention and therapy. Nutrients 7 (9), 8199e8226. Hajipour, M.J., et al., 2012. Antibacterial properties of nanoparticles. Trends Biotechnol. 30 (10), 499e511. Hanafi, M.E.S., Sallam S.M., Mohamed F.A., Characterization of Hydroxyapatite Doped with Different Concentrations of Magnesium Ions. Healy, K.E., Ducheyne, P., 1992. Hydration and preferential molecular adsorption on titanium in vitro. Biomaterials 13 (8), 553e561. Huang, T., et al., 2011. Nanostructured Si, Mg, CO32 substituted hydroxyapatite coatings deposited by liquid precursor plasma spraying: synthesis and characterization. J. Therm. Spray Technol. 20 (4), 829e836. Ibrahim, M.Z., et al., 2017. Biomedical materials and techniques to improve the tribological, mechanical and biomedical properties of orthopedic implants e a review article. J. Alloy. Comp. 714, 636e667. Imrie, F.E., et al., 2013. Synthesis and characterisation of strontium and magnesium cosubstituted biphasic calcium phosphates. In: Key Engineering Materials. Trans Tech Publ. Iqbal, N., et al., 2015. Microwave Assisted Synthesis and Characterization of Magnesium Substituted Calcium Phosphate Bioceramics, vol. 56. Jiao, M.-J., Wang, X.-X., 2009. Electrolytic deposition of magnesium-substituted hydroxyapatite crystals on titanium substrate. Mater. Lett. 63 (27), 2286e2289.

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Jones, J., Schwartz, R., Krook, L., 1980. Calcium homeostasis and bone pathology in magnesium deficient rats. Calcif. Tissue Int. 31 (1), 231e238. Kannan, S., et al., 2005. Synthesis and characterization of magnesium substituted biphasic mixtures of controlled hydroxyapatite/b-tricalcium phosphate ratios. J. Solid State Chem. 178 (10), 3190e3196. Kim, T.-W., et al., 2012. In situ synthesis of magnesium-substituted biphasic calcium phosphate and in vitro biodegradation. Mater. Res. Bull. 47 (9), 2506e2512. Kolmas, J., Groszyk, E., Kwiatkowska-Roz_ ycka, D., 2014. Substituted hydroxyapatites with antibacterial properties. BioMed Res. Int. 2014. Lacefield, W.R., 1999. Materials characteristics of uncoated/ceramic-coated implant materials. Adv. Dent. Res. 13, 21e26. Landi, E., et al., 2008. Biomimetic Mg-substituted hydroxyapatite: from synthesis to in vivo behaviour. J. Mater. Sci. Mater. Med. 19 (1), 239e247. Laurencin, D., et al., 2011. Magnesium incorporation into hydroxyapatite. Biomaterials 32 (7), 1826e1837. Liangzhi, G., Weibin, Z., Yuhui, S., 2016. Magnesium substituted hydroxyapatite whiskers: synthesis, characterization and bioactivity evaluation. RSC Adv. 6 (115), 114707e114713. Loszach, M., Gitzhofer, F., 2015. Induction suspension plasma sprayed biological-like hydroxyapatite coatings. J. Biomater. Appl. 29 (9), 1256e1271. Lu, W.C., Pringa, E., Chou, L.S., 2017. Effect of magnesium on the osteogenesis of normal human osteoblasts. Magnes. Res. 30 (2), 42e52. Mackie, E.J., 2003. Osteoblasts: novel roles in orchestration of skeletal architecture. Int. J. Biochem. Cell Biol. 35 (9), 1301e1305. Maguire, M.E., Cowan, J.A., 2002. Magnesium chemistry and biochemistry. Biometals 15 (3), 203e210. Mayer, I., Schlam, R., Featherstone, J.D.B., 1997. Magnesium-containing carbonate apatites. J. Inorg. Biochem. 66 (1), 1e6. Mehrjoo, M., et al., 2015. Effect of magnesium substitution on structural and biological properties of synthetic hydroxyapatite powder. Mater. Express 5 (1), 41e48. Miranda, G., et al., 2019. Surface design using laser technology for Ti6Al4V-hydroxyapatite implants. Opt. Laser. Technol. 109, 488e495. Mroz, W., et al., 2015. In vivo implantation of porous titanium alloy implants coated with magnesium-doped octacalcium phosphate and hydroxyapatite thin films using pulsed laser depostion. J. Biomed. Mater. Res. B Appl. Biomater. 103 (1), 151e158. Nazir, R., et al., 2011. Rapid Synthesis of Thermally Stable Hydroxyapaptite, vol. 38. O’Neill, L., et al., 2009. Deposition of substituted apatites onto titanium surfaces using a novel blasting process. Surf. Coat. Technol. 204 (4), 484e488. Oguchi, H., et al., 1995. Long-term histological evaluation of hydroxyapatite ceramics in humans. Biomaterials 16 (1), 33e38. Onder, S., et al., 2013. Magnesium substituted hydroxyapatite formation on (Ti,Mg)N coatings produced by cathodic arc PVD technique. Mater. Sci. Eng. CMater. Biol. Appl. 33 (7), 4337e4342. Posner, A.S., 1969. Crystal chemistry of bone mineral. Physiol. Rev. 49 (4), 760e792. Rajendran, A., et al., 2018. Multi-element substituted hydroxyapatites: synthesis, structural characteristics and evaluation of their bioactivity, cell viability, and antibacterial activity. J. Sol. Gel Sci. Technol. 1e18. Rau, J.V., et al., 2013. Nanostructured Si-substituted hydroxyapatite coatings for biomedical applications. Thin Solid Films 543, 167e170.

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Ren, F., et al., 2010. Synthesis, characterization and ab initio simulation of magnesiumsubstituted hydroxyapatite. Acta Biomater. 6 (7), 2787e2796. Serre, C., et al., 1998. Influence of magnesium substitution on a collageneapatite biomaterial on the production of a calcifying matrix by human osteoblasts. J. Biomed. Mater. Res. 42 (4), 626e633. Sharifnabi, A., et al., 2014. The structural and bio-corrosion barrier performance of Mgsubstituted fluorapatite coating on 316L stainless steel human body implant. Appl. Surf. Sci. 288, 331e340. Simchi, A., et al., 2010. Recent Progress in Inorganic and Composite Coatings with Bactericidal Capability for Orthopaedic Applications, vol. 7, pp. 22e39. Sivaraj, D., Vijayalakshmi, K., 2018. Substantial effect of magnesium incorporation on hydroxyapatite/carbon nanotubes coatings on metallic implant surfaces for better anticorrosive protection and antibacterial ability. J. Anal. Appl. Pyrolysis 135, 15e21. Stipniece, L., et al., 2014. Characterization of Mg-substituted hydroxyapatite synthesized by wet chemical method. Ceram. Int. 40 (2), 3261e3267. Suchanek, W.L., et al., 2004. Preparation of magnesium-substituted hydroxyapatite powders by the mechanochemicalehydrothermal method. Biomaterials 25 (19), 4647e4657. Suchanek, W.L., et al., 2004. Mechanochemical-hydrothermal synthesis of calcium phosphate powders with coupled magnesium and carbonate substitution. J. Solid State Chem. 177 (3), 793e799. Vandrovcova, M., et al., 2015. Pulsed laser deposition of magnesium-doped calcium phosphate coatings on porous polycaprolactone scaffolds produced by rapid prototyping. Mater. Lett. 148, 178e183. Walker, G.M., 1986. Magnesium and cell cycle control: an update. Magnesium 5 (1), 9e23. Webler, G.D., et al., 2015. Mg-doped biphasic calcium phosphate by a solid state reaction route: characterization and evaluation of cytotoxicity. Mater. Chem. Phys. 162, 177e181. Wuthier, R., 1971. Zonal analysis of electrolytes in epiphyseal cartilage and bone of normal and rachitic chickens and pigs. Calcif. Tissue Res. 8 (1), 24e35. Yajing, Y., et al., 2014. Magnesium substituted hydroxyapatite coating on titanium with nanotublar TiO2 intermediate layer via electrochemical deposition. Appl. Surf. Sci. 305, 77e85. Zhao, S.f., et al., 2013. Effects of magnesium-substituted nanohydroxyapatite coating on implant osseointegration. Clin. Oral Implant. Res. 24, 34e41. Ziani, S., Meski, S., Khireddine, H., 2014. Characterization of magnesium-doped hydroxyapatite prepared by sol-gel process. Int. J. Appl. Ceram. Technol. 11 (1), 83e91.

Zinc-substituted hydroxyapatite

9

Kashif Ijaz, Hamad Khalid, Aqif Anwar Chaudhry Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

9.1

The biological importance of zinc ion and its relevance to calcium phosphates

Zinc is the most abundant trace element in the bone. It is vital for human and animal metabolic processes. Its quantity ranges from 0.012 to 0.0250 wt% in human bone. Furthermore, zinc promotes bone growth and mineralization. Zinc deficiency can also be associated with retardation of bone growth (Ito et al., 2000; Yamaguchi et al., 1987). Presence of zinc has a positive effect on osteoblast cells, and it also restricts bone resorption by osteoclast cells (Yamaguchi, 1998). This forms the fundamental desire for substitution of zinc ions into the hydroxyapatite (HA) lattice. It has been demonstrated that incorporation of zinc into HA led to decreased production of inflammatory cytokines and increased production of antiinflammatory cytokines (by monocytes) (Grandjean-Laquerriere et al., 2006). Zinc in lower doses has been reported to increase alkaline phosphatase (ALP) activity and DNA content in bone tissue. Deficiency of zinc can therefore retard bone formation. Loss of zinc is caused by zinc-deficient diets, aging, and/or skeletal unloading. Long periods of bed rest in humans are linked to skeletal unloading, which in turn speeds up the excretion of zinc. This effects bone mineral density adversely. Therefore, presence of additional zinc ions (after implantation) can promote bone growth around the implant and also accelerate the bone growth process in elderly humans. However, additional zinc should be given via a slow release mechanism because elevated levels of zinc can cause adverse reactions (Ito et al., 2000). ALP activity on the osteoblasts cells increases with different amounts of zinc. Zinc deficiency, on the other hand, affects because of cell-mediated immune response, and, in elders, it has been linked to the development of osteoporosis (Gomes, 2011). Biological apatite (the mineral component of bone) is a calciumdeficient, nonstoichiometric apatite, which contains either cationic substitutions (replacing calcium), i.e., magnesium (Mg2þ), sodium (Naþ), zinc (Zn2þ), strontium (Sr2þ), aluminum (Al3þ), potassium (Kþ), and silver (Agþ), or anionic substitutions 2 4  (replacing OH or PO3 4 ), i.e., carbonate (CO3 ), silicate (SiO4 ), fluoride (F ), and  chloride (Cl ) (Shepherd et al., 2012; Camaioni et al., 2015; Khan et al., 2014). These substitutional ions impart significantly improved biological characteristics to synthetic HA by affecting surface charge, crystallinity, thermal stability, surface

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00009-4 Copyright © 2020 Elsevier Ltd. All rights reserved.

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chemistry, solubility, and crystal morphology. These all subsequently influence biological activity by increasing osteoblast cell proliferation and hence enhanced osteointegration (Shepherd et al., 2012; Camaioni et al., 2015; Pietak et al., 2007). Because of the natural occurrence of zinc in bone (and other ions) and important biological functions associated to it, ionic substitutions into HA are the subject of significant research and development activities (Pietak et al., 2007). The ability of the HA structure to accept ions allows design and development of new and better performing calcium phosphates tailored for specific applications (Pietak et al., 2007; Cacciotti, 2014). The positive effects of zinc are evident in Zn-incorporated biomaterials, which also include bone cements, bioactive glassebased materials, and coatings (Qiao et al., 2014). Substitutions of more than one ion/s have also been attempted. For example, cosubstitutions of zinc and fluorine ions into HA resulted in an increase in cell proliferation and ALP activity of cells (and also improved mechanical properties) (Uysal et al., 2014).

9.2

Synthesis of zinc-substituted hydroxyapatite

Several synthesis methods can be used to develop zinc-substituted HA (Zn-HA). Table 9.1 summarizes these methods with corresponding references.

9.2.1

Solegel methodology

Metal oxides, ceramics, and glasses are often synthesized using solegel methodology. In this method, the reactants are dissolved in appropriate solvents which results in formation of a gel. This gel is then processed into the final product. There are series of discrete steps in the solegel method. First, stable solutions of precursors are prepared. These are then transformed into a gel for via oxide or alcohol bridging due to polyesterification or polycondensation. In the third step, the gel is solidified. This step is also Table 9.1 List of zinc-substituted hydroxyapatite (Zn-HA) synthesis methods with references. Number

Synthesis method

References

1

Solegel methodology

(Gomes et al., 2012; Kaygili and Tatar, 2012; EshtiaghHosseini et al., 2007; Naqshbandi et al., 2014; Negrila et al., 2018)

2

Coprecipitation method

(Ito et al., 2000; Kaygili and Tatar, 2012; Ashuri et al., 2012; Fujii et al., 2006; Thian et al., 2013a; Ren et al., 2009; Cox et al., 2014)

3

Hydrothermal synthesis

(Thian et al., 2013a; Xiao et al., 2008; Radovanovic et al., 2012; Stojanovic et al., 2009)

4

Solid-state methods

(Boyer et al., 1997; Leshkivich and Monroe, 1993; Hahn et al., 2010)

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219

called aging. After aging, dehydration and heat treatment are carried out (Cushing et al., 2004). A typical solegel reaction for Zn-HA synthesis involves preparation of calcium and phosphorous solutions (while maintaining a Ca:P ratio of 1.67). Calcium nitrate tetrahydrate is typically used as the calcium source. Phosphorus pentoxide (P2O5), triethyl phosphate ((C2H5O)3PO), and diammonium hydrogen phosphate have been used as sources of phosphate ions. The zinc source (commonly zinc nitrate hexahydrate) is often dissolved in the calcium precursor solution before mixing. As highlighted earlier, these solutions are then mixed to obtain a gel, which is then dried before heat treatment (typically between 900 and 1200 C for 1e15 h) (Gomes et al., 2012; Kaygili and Tatar, 2012; Eshtiagh-Hosseini et al., 2007; Naqshbandi et al., 2014; Negrila et al., 2018).

9.2.2

Coprecipitation method

Coprecipitation refers to the simultaneous occurrence of nucleation, growth, coarsening, and/or agglomeration of particles from precursor solutions. Products of precipitation reactions are relatively insoluble species formed under high supersaturation. Because of a chemical reaction, supersaturation occurs. The key and primary step in coprecipitation is nucleation because of which large numbers of particles are formed. After nucleation, secondary processes such as Ostwald ripening and aggregation are started. These secondary processes are responsible for determining morphology and size of precipitates. Usually, to attain stochiometric ratio of HA, the precursors are coprecipitated by the addition of basic solution, which helps in proper aging (Cushing et al., 2004). A typical reaction involves dropwise addition of calcium and phosphorous and zinc-containing precursors into each other, which results in immediate precipitation. Typical calcium ion sources are calcium nitrate tetrahydrate or calcium hydroxide. Typical phosphate ion sources are diammonium hydrogen phosphate or phosphoric acid. Zinc nitrate is one of the most common sources of zinc ion and is generally added to the calcium precursor before reaction with the phosphate ion source solution. Typical addition rates of reactants vary between 5 and 30 mL/min. Often pH is maintained above 10 by addition of a base (sodium hydroxide or ammonium hydroxide). This is done to ensure that the resulting phase is apatitic. An aging stage follows addition, which may or may not employ heating. Aging times vary typically between few hours to several days. Use of heating reduces the need for longer aging times. After aging, washing is done to wash out unreacted precursors (if any) before oven dry typically between 60 and 140 C and often up to 24 h. This is often followed by heat treatments typically between 900 and 1200 C for 1e15 h to improve crystallinity (Ito et al., 2000; Kaygili and Tatar, 2012; Ashuri et al., 2012; Fujii et al., 2006; Thian et al., 2013a; Ren et al., 2009; Cox et al., 2014). A typical reaction for synthesis of Zn-HA can be represented by the following equation (Fadeeva et al., 2012); 1 (10-x)Ca2þ þ xMnþ þ 6PO3 / Ca10-xMx(PO4)6(OH)2 4 þ 2OH

where Mnþ represents Zn2þ.

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9.2.3

Handbook of Ionic Substituted Hydroxyapatites

Hydrothermal synthesis

Hydrothermal method (a batch process) is a type of chemical reaction in which the temperature of the solvent is increased above its boiling point at elevated pressure in a sealed container (Teflon or metallic container). During a batch hydrothermal reaction, ceramic sols are produced by chemical reactions under pressure and heat generally in the presence of an alkali or acid (Rao et al., 2017). This route offers an environment-friendly (green) and low-temperature direct route to submicron or nanosized particles without the need of a subsequent heat treatment step. Indeed, reports of nanoparticle synthesis using batch hydrothermal processing have increased in recent years. In hydrothermal synthesis, reaction involves mixing of calcium source (usually calcium nitrate hexahydrate), phosphorous source (commonly diammonium hydrogen phosphate), and a zinc source (mostly zinc nitrate hexahydrate) in a sealed Teflon container. This results in immediate precipitation. Pressure, temperature, pH, and aging time affect the composition, crystallinity, and morphology of Zn-HA particles. As compared with other methods, batch hydrothermal synthesis produces crystalline, well-defined, and phase-pure Zn-HA in a single step. Aging/vessel heating times can vary from a few hours to 24 h. Synthesis temperatures reported in literature vary between 90 and 200 C. Use of templating agents (e.g., poly(amidoamine), CTAB, PEG) has also been reported. Additionally, few scientists have reported on posthydrothermal synthesis heat treatments between 400 and 800 C for 1e4 h (Thian et al., 2013a; Xiao et al., 2008; Radovanovic et al., 2012; Stojanovic et al., 2009).

9.2.4

Solid-state methods

Solid-state methodology involves the synthesis of Zn-HA by forcing a reaction between solid zinc, calcium, and phosphate precursors. Hattori et al. synthesized Zn-HA by using calcium oxide, calcium hydrogen phosphate, and zinc oxide powder. These powder were mixed in purified deionized water and mixed in a ball mill. These mixed powder were heated to 400 and 800 C. Based on X-ray diffraction (XRD) and infrared spectroscopy data, the authors concluded that after 2.5 h of grinding, the individual nature of precursors was lost and an amorphous apatitic phase was obtained (Hattori et al., 2016). The synthesis of apatites substituted with silicates, sulfates, lanthanum, and fluorine has also been reported in literature (Boyer et al., 1997; Leshkivich and Monroe, 1993; Hahn et al., 2010).

9.3

Zinc-substituted hydroxyapatite coatings

Zn-HA can also be deposited as a coating on surfaces useful for biomedical applications. The demand for metal prostheses in orthopedic, maxillofacial, and dental surgery has tremendously increased during the last few decades. The main crucial aspect behind the success or failure of any metallic implant lies in its efficiency to

Zinc-substituted hydroxyapatite

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bond quickly with the existing surrounding bone. Hence, to accelerate the process of osteointegration, the ideal metallic implant material should display compositional and structural similarity to those of native biological infrastructure. To achieve this aim, one of the most commonly used approaches is to coat the implant surface with a coating similar to the main constituent of mineralized bone with improved biological activity. One of the most commonly used metal implants is titanium (Ti). Ti and its alloys display exceptional mechanical properties, good biocompatibility, good fatigue resistance, and high corrosion resistant properties in comparison with other metals (Vladescu et al., 2014; Graziani et al., 2017; Rau et al., 2015). The challenge with Ti implants is their limited osteogenic properties, which can consequently result in premature failure at implantation site. One solution to address this problem is to modify Ti implant surfaces with bioactive coatings (Ho and Ding, 2015). Guanfie et al. used the electrophoretic technique for deposition of Zn-HA. Their results suggested good adhesion between coatings and substrates. Formation of new apatite layer on the coatings was observed after 7 days of immersion in a simulated body fluid (suggesting apatite layer formation ability). The authors concluded that for the Zn/Ca ratio of 5%, dense and uniform Zn-HA coatings were observed (Sun et al., 2014). A report on electrophoretic deposition utilized 20  20 mm stainless steel plates that were used as a cathode (with a carbon rod serving as an anode). Zn-HA powder was dispersed in n-butanol containing w2% triethanolamine (TEA). This was followed by ultrasonication and overnight aging. Adsorption of TEA onto the HA particles made them positively charged. Acid (HCl) was added to achieve homogeneous dispersion. The distance between the two electrodes (stainless steel and carbon rod) was 10 mm. A constant voltage of 30 V was used for deposition. This led to deposition of coatings and was followed by drying via heating (Sun et al., 2014).

9.4

Characterization of zinc-substituted hydroxyapatite

9.4.1 9.4.1.1

Electron microscopy Scanning electron microscopy

Predoi et al. synthesized Zn-HA (Ca10xZnx(PO4)6(OH)2 with 0.01  x  0.05) using coprecipitation method (Predoi et al., 2017). They observed the morphology of synthesized Zn-HA using scanning electron microscopy (SEM) and assessed elemental composition using energy-dispersive X-ray (EDX) analysis. The SEM images showed that with an increase in zinc concentration, the particle size decreased; however, the morphology remained similar (see Fig. 9.1(a)). These results were in agreement with a study conducted by Ramya et al. (Ramya et al., 2014). EDX analysis confirmed the presence of zinc, calcium, phosphorus, and oxygen. The molar ratio of (calcium þ zinc)/phosphorus was almost 1.66  1, which was close to theoretical stoichiometric ratio (see Fig. 9.1(b)) (Predoi et al., 2017).

222

(a)

Handbook of Ionic Substituted Hydroxyapatites

(b) Zn: HAp (xZn=0.01)

Zn: HAp (xZn=0.03)

Zn: HAp (xZn=0.05)

Figure 9.1 Scanning electron microscopy (SEM) images (a) and energy-dispersive X-ray (EDX) analysis (b) of zinc-substituted hydroxyapatite (Zn-HA) synthesized by precipitation method (Predoi et al., 2017). Open access.

Ma et al. utilized a field emission electron microscope (FESEM) for observing the morphology of Zn-HA obtained (Ma and Qin, 2015). Fig. 9.2 shows the FESEM images of Zn-HA. The morphology was referred to as wrinkled flowers by the authors. When sodium alginate was used as controlling agent, some large sheets were also observed (see Fig. 9.2(c)). These sheets were not visible in transmission electron microscopy (TEM) as the small grid of TEM was able to capture only small particles.

Zinc-substituted hydroxyapatite

223

(a)

(b)

(c)

(d)

Figure 9.2 Field emission electron microscope (FESEM) of zinc-substituted hydroxyapatite (Zn-HA), (a) Zn/(Zn þ Ca) ¼ 5%, controlled by silk fibroin only; (b) Zn/(Zn þ Ca) ¼ 10%, controlled by silk fibroin only; (c) Zn/(Zn þ Ca) ¼ 5% controlled by silk fibroin and alginate; and (d) Zn/(Zn þ Ca) ¼ 10%, controlled by silk fibroin and alginate. Reprinted from the permission from Ma, J., Qin, J., 2015. Graphene-like zinc substituted hydroxyapatite. Cryst. Growth J. Des. 15(3), 1273e1279. Copyright 2015, American Chemical Society.

9.4.1.2

Transmission electron microscopy

Ma et al. influenced the precipitation mechanism by using silk fibroin and sodium alginate (Ma and Qin, 2015). The TEM images showed that when HA was incorporated with zinc, the morphology of apatite changed to graphene-like sheets from rectangle flakes (see Fig. 9.3). According to the TEM images, the length and width of HA was measured as 97  19 and 19  7 nm, respectively. The thickness was measured between 5 and 10 nm. Most of the particles of HA had sharp edges (see Fig. 9.3(b)). Similar morphology was also reported by Wang et al. (Wang et al., 2010). Most of the sheets of Mg-HA ((Zn/Zn þ Ca) ¼ 5) were w3 nm in thickness

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(a)

(b)

(c)

(d)

Figure 9.3 Transmission electron microscopy (TEM) images of hydroxyapatite (HA) (a, b) and zinc-substituted HA (Zn-HA). (c) (Zn/Zn þ Ca) ¼ 5% and (d) (Zn/Zn þ Ca) ¼ 10%. Reprinted from the permission from Ma, J., Qin, J., 2015. Graphene-like zinc substituted hydroxyapatite. Cryst. Growth J. Des. 15(3), 1273e1279. Copyright 2015, American Chemical Society.

(see Fig. 9.3(c)) and their size was more than 100 nm. The dark shades on the surface of sheets could be folds of sheets, which gives a wrinkled paper-like look to these sheets. Increase in zinc content did not affect the morphology. Sodium alginate was also used in combination with silk fibroin. The alginate did not change the structure of sheets, but the crystallinity of Zn-HA increased.

9.4.2

Surface area analysis

Surface area analysis using gas adsorption offers key insights into physical characteristics of a powdered sample. Predoi et al. used a coprecipitation method to develop Zn-HA (ranging between 1 and 5 at% Zn as Zn/(Zn þ Ca)). They determined BrunauereEmmeteTeller (BET) surface area (using N2 adsorption). The surface area increased from w100 to 115 m2/g when the zinc content increased from 1 to 5 at%. In this report, the authors also conducted BarreteJoynereHalenda N2 adsorption/desorption analyses to measure pore size and volume. They reported a decrease in pore size and volume with an increase in zinc content (Predoi et al., 2017).

Zinc-substituted hydroxyapatite

225

Yu et al. developed composites based on collagen scaffolds, which contain microspheres of mesoporous Zn-HA. For samples containing no zinc, 2 at%, and 5 at% zinc, the BET surface area increased from w84, 106, and 131 m2/g, respectively (Yu et al., 2017).

9.4.3

X-ray diffraction

XRD is used for assessment of phase purity and crystallinity of inorganic materials. Ren et al. developed a range of Zn-HA compositions and evaluated them using XRD. They recorded the diffraction patterns in 3e70 degrees (2q) range using Cu Ka radiation (0.05 degrees step size and 1 s per step). To carry out Rietveld refinement analysis, longer and more intensive patterns were obtained (10e100 degrees 2q range, 0.03 degrees step size, and 10 s per step). It was reported that the broadness of the peaks in XRD patterns obtained increased possibly because of decreasing crystallinity as a function of increasing zinc content (5e20 mol% Zn). The authors concluded that for 20 mol% (referring to Zn/(Ca þ Zn)), the apatitic structure was lost. XRD patterns representing this are show in Fig. 9.4. It was reported that further increase in zinc content led to appearance of different calcium zinc phosphate phases. Using Rietveld refinement analysis, the authors concluded that the lattice parameters decreased with increase in zinc content up to 10 mol% (Ren et al., 2009). Thian et al. observed broad XRD peaks of Zn-HA and suggested that these were due to a smaller crystal size. Furthermore, zinc incorporation did not affect the phase puritydi.e., it gave a good match to XRD patter of phase pure HA. Substitution of zinc ion into the apatite structure was also reported to result in a decrease of a and c

211 002

Zn20HA

Intensity (a.u.)

300

310

213

Zn15HA Zn10HA

202

222

Zn5HA 004

10

20

30

40 2θ (degree)

50

HA

60

70

Figure 9.4 X-ray diffraction patterns of zinc-substituted hydroxyapatite (Zn-HA). From Ren et al. (2009), copyright 2015 reproduced with the permission of Elsevier.

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axes. The authors attributed this to the radius of zinc ion being smaller than calcium ion (0.074 nm as compared with 0.099 nm, respectively) (Thian et al., 2013b).

9.4.4

Fourier-transform infrared spectroscopy

Fourier-transform infrared (FTIR) spectroscopy is an important technique for identifying key groups in the HA (and related calcium phosphates) lattice. Ionic substitution influences the crystal lattice, which in term influences thermal stability and sometimes the quantity of ions (which are being substituted). These can be closely observed at times using FTIR. The peak of PO3 4 is a characteristic peak of HA (Rehman and Bonfield, 1997). Thian et al. compared FTIR spectra of HA and Zn-HA. According to this report by Thian et. al. peaks corresponding to n4 bending vibrations of the OePeO 1 linkage in PO3 4 at 560 and 601 cm . Peaks related to the asymmetric stretching 1 (n3) vibrations of the PeO bond in PO3 4 were observed at 1088 and 1036 cm . A 1  broad band at 3750 cm (corresponding to OH ) overlapped with the band representing water present in free form (3310 and 3750 cm1). CO2 3 stretching vibrations (n3) at 1415 and 1640 cm1 were also observed (Thian et al., 2013a). Guerra-Lopez et al. discussed the effect of varying zinc concentrations in HA using a coprecipitation method. It was observed that with increase in the concentration of zinc, the intensity of water bands between 3452 and 1631 cm1 increased (see Fig. 9.5), suggesting that there was an increase in water absorption/retention. The stretching OH vibrations

Zn25

Transmittance (%)

Zn20 Zn15 Zn10 Zn5

961

CaHap

3600

3200

1600

1200

800

Wave number (cm–1)

Figure 9.5 Fourier-transform infrared (FTIR) spectra of hydroxyapatite (HA) (Ca-HAp) and zinc-substituted HA (Zn-HA) with 5 wt% (Zn5) to 25 wt% (Zn25) of zinc. From Guerra-Lopez et al., (2015), copyright 2015 reproduced with the permission of Elsevier.

Raman intensity (a.u.)

Zinc-substituted hydroxyapatite

227

Zn20 Zn15 Zn10 Zn5 CaHap 1100 1050 1000 990

980

970

960

950

940

600

550

450

400

Wave number (cm–1)

Figure 9.6 Raman spectra of phase pure hydroxyapatite (HA) and zinc-substituted HA (Zn-HA). From Guerra-Lopez et al., (2015), copyright 2015 reproduced with the permission of Elsevier.

shifted to lower wave number (3569 cm1), and the intensity of the corresponding peaks reduced with an increase in zinc content. However, an opposite trend was observed for liberational mode, where the values shifted to a higher wave number (633 cm1). Increased concentration of zinc in Zn-HA introduces another phase, which was confirmed by the appearance of a band at 941 cm1 and a shoulder at 619 cm1 (in Zn-HA with 25 wt% zinc). This second phase could be zinc ammonium phosphate, which possibly crystalized during reaction due to high concentration of zinc. Appearance of a weak band at 1430 cm1 (denoted to NHþ 4 ) also supports this suggestion (Guerra-L opez et al., 2015).

9.4.5

Raman spectroscopy

Comparisons between Zn-HA and HA using Raman spectroscopy were reported by Guerra-L opez et al. (see Fig. 9.6) (Guerra-L opez et al., 2015). It was observed that when zinc was incorporated in HA, a small shift of phosphate peak (961e963 cm1). However, this shift did not increase with an increase in zinc concentration. The PeO showed very weak antisymmetric stretching vibrations between 1020 and 1100 cm1. A peak at 1054 cm1 was the most intense among the phosphate peaks, which showed slight shift to higher wave number. Other bending modes (n2 and n4) of phosphate were observed between 600 and 400 cm1, which shifted toward higher wave number by the addition of zinc. For CeO vibrations, a strong n1 peak at 1080 cm1 and weak n3 peaks/bands between 1400e1470 cm1 and 850e890 cm1 (characteristic peak of carbonate) were generally observed in FTIR spectra; however, neither of them were observed in HA and Zn-HA in this case. From this data, it was concluded that no carbonate was present in the samples (Guerra-L opez et al., 2015; De Aza et al., 1997).

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Table 9.2 Thermogravimetric analysis (TGA) data of zinc-substituted hydroxyapatite (Zn-HA) (Zn3eZn20 refer to Zn wt% from 3 to 20). Mass fraction loss (%) Sample

(0e2008C)

(200e5008C)

(500e8008C)

(800e10008C)

Total loss (%)

Zn3

7.4

1.8

1.6

1.0

10.0

Zn5

10.6

3.1

1.7

0.9

13.2

Zn10

13.3

3.4

1.6

0.9

15.7

Zn20

14.3

3.7

1.9

0.5

16.7

From: Guerra-L opez et al., (2015), copyright 2015 reproduced with the permission of Elsevier.

9.4.6

Thermogravimetric analysis

Thermogravimetric analysis (TGA) is used for assessing thermal properties of a material. In ion-substituted HAs, this is a critical assessment because ion substitution alters the crystal structure of HA, and in doing so, the thermal stability can be lost. This results in formation of different phases when substituted HAs are exposed to elevated temperatures. Guerra-L opez et al. preformed TGA of Zn-HA (with 3, 5, 10, and 20 wt% zinc with respect to calcium) synthesized using a precipitation method. In TGA studies, they observed that there were three main temperature regions of weight loss, first around 200 C, second between 200 and 500 C, and third one between 700 and 890 C. In first two temperature regions, the weight loss of samples increased with an increase in zinc content determined to be w7.4%, 10.6%, 13.3%, and 14.3% for 3%, 5%, 10%, and 20% zinc-containing Zn-HA. Weight loss at higher temperatures did not display any association with zinc content (see Table 9.2). Table 9.2 represents TGA data of the aforementioned Zn-HA (Guerra-Lopez et al., 2015). Weight lost trends in first temperature zone (0e200 C) show that the increase of zinc content led to increase in water content. This is a finding endorsed by reported Zn-HA FTIR spectra as well (Ren et al., 2009; Miyaji et al., 2005). It is already reported that HA decomposes in to b-tricalcium phosphate on heating above 800 C for substituted HAs and/or for HA with decreased thermal stability (Resende et al., 2006; Guerra-L opez et al., 2001).

9.5 9.5.1

Biological performance of zinc-substituted hydroxyapatite Cell response

Presence of zinc influences both cells related to bone deposition and resorption, i.e., osteoblasts and osteoclasts, respectively. Therefore, assessment of response of cells

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to Zn-HA is of key importance. In a report on assessment of bioactivity of Zn-HA, Thian et al. reported increased growth of human adipose-derived mesenchymal stem cells, which was accompanied by an increase in bone cell differentiation markers (type I collagen) (Thian et al., 2013b). Miao et al. determined bioactivity of HA films (which were deposited on to Ti6Al4V substrates using solegel and dipping method), which contained both fluorine and zinc. Using human osteoblast-like cells (MG63), they reported better cell attachment and spread of the cells on coatings, which contained both zinc and fluorine as compared with fluorine alone. This was attributed to the possible release of zinc ions from the surface and positive interactions with the surrounding cells (Miao et al., 2007). Monocytes, which are white blood cells linked to the immune system, get influenced by presence of zinc (from Zn-HA) to release antiinflammatory cytokines (GrandjeanLaquerriere et al., 2006). In a study on Zn-HA and its effect on MC3T3-E1 cells (mouse osteoblastic cells), Sogo et al. showcased that the optimum zinc content for ZnHA ceramics (for promoting proliferation of osteoblast cells) was 0.34 wt%. This conclusion was possibly due to higher incorporation of zinc leading to formation of a zinc releasing calcium phosphate material with higher levels of zinc release not being desirable (Sogo et al., 2004). Other cells lines such as HeLa (derived from cervical cancer cells) have also been used to assess cell viability. Predoi et al. highlighted that Zn-HA coatings on a silicon substrate did not show any toxicity toward HeLa cells (Predoi et al., 2019). Wang et al. investigated biological activities of codoped/substituted HA scaffolds. In this study Zn-doped/substituted scaffolds promoted proliferation of bone marrow stromal cells (Wang et al., 2018). Uptake and signaling of receptor activator of nuclear factor (NF)-kb ligand (RANKL) is reduced because of RANK suppression by zinc. This restricts the differentiation of preosteoclasts into osteoclasts (Yamaguchi and Uchiyama, 2004). Shepherd et al. reported inhibition of osteoclast-like cell formation by Zn-HA. They also proposed Zn-HA to be a more suitable alternative to unsubstituted HA in the light of reduced resorption and osteoclast formation that was observed in their study (Shepherd et al., 2014).

9.5.2

Antibacterial response

Zn-HA has also proven to be a useful antibacterial agent in bone-associated infections. It has also been considerably used as therapeutically drug delivery carrier in pharmaceutical industry alongside its mechanical supportive and osseointegrative stimulatory role. Regarding this, mesoporous silica may be introduced into the Ap crystal lattice because of its capability of increasing the surface area, hence resultantly has the tendency to attach more drug molecules for delivery. Furthermore, peptides, i.e., osteostatin (pentapeptide with both in vitro and in vivo stimulative osteogenic potential) as well as insulin can also be attached through

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covalent bonding or adsorption to Zn-HA scaffolds. Zinc HAefilled composite resins with reported excellent biocompatibility and no toxicity have also been fabricated by Chadda et al. for dental applications (Chadda et al., 2016). Sych et al. successfully used HA substituted with 3 and 7 wt% zinc that has also been used for delivery of antibiotic “rifampicin,” which has established activity against leprosy, tuberculosis, and bacterial osteomyelitis (Sych et al., 2015). HA and mesoporous silica have also been employed to deliver bisphosphonates, i.e., alendronate and zoledronic acid; both suppress the osteoclasts activity to avoid bone resorption and stimulate bone development by osteoblasts. Drug release studies as well as in vitro biological activity studies on bone marrow mesenchymal stem cells and osteoclast cell line showed more gradual release from the composite in contrast to the pure silica (Huang et al., 2011; Zhu et al., 2014; Szurkowska and Kolmas, 2017). When associated with glass-based cements and coatings for bone regeneration, Zn has proven activity against Staphylococcus aureus and other bacteria associated with acute or chronic bone infections (Gomes, 2011). Anwar et al. studied the antibacterial activity of zinc-incorporated HA against Bacillus subtilis, S. aureus, Enterobacter aerogenes, and Escherichia coli. It was found that 2 wt% doping of Zn was more lethal to all the bacterial strains than 0 wt% and 4 wt% of Zn doping (Anwar et al., 2016).

9.5.3

Animal model

Cruz et al. published a detailed systematic review focused on answering the question whether zinc incorporation into calcium phosphate improve bone repair. In several animal studies, the form in which zinc has been utilized includes Zn-HA alone or in addition to HA. In several cases, Zneb-TCP has also been utilized. In most of the cases, the synthesis routes of these materials were wet precipitation routes (which include coprecipitation, solegel techniques, hydrothermal synthesis, and microwave-assisted wet chemical synthesis). Materials were implanted in the form of cylinders, granules, powder, or microspheres. The sizes of these implants varied between 2 and 8 mm (diameter) and 4e10 mm (length) for cylindrical forms. These materials and forms were obtained often after thermal treatments at elevated temperatures (between 1000 and 1200 C for 2e27 h). Table 9.3 summarizes the kinds of animals used and implantation sites selected. Canines, rabbits, and rats have been used. Implantation sites include the tibia, calvaria, femur, and even intramuscular sites. The experimental periods vary between 7 days and 78 weeks. The zinc content in materials used varied between 0.063 wt% and 5 wt%. Key conclusions of most of the studies were that inclusion of zinc did improve bone repair in certain cases (Cruz et al., 2018).

Animal model/Local of implantation

Experimental periods

Rabbits n ¼ 5/tibia

Biomaterials

Zn %

Main conclusion

References

7, 14, 28 days

HA/Zn-HA (cylinder 2  6 mm, sintered)

0.3 wt%

Zn-doping did not improve bone repair

Calasans-Maia et al. (2014)

Rat n ¼ 5/calvaria (8 mm)

30, 90, 180 days

Autogenous bone/HA/Zn-HA (powder)

3.16 wt% ND

Zn-doping improved bone repair

Maia et al. (2008a)

Rabbits n ¼ 5/tibia

7, 14, 28 days

HA/Zn-HA (cylinder 2  6 mm)

3.16 wt%

Zn-doping improved bone repair

Maia et al. (2008b)

Rat n ¼ 5/calvaria (8 mm)

30, 90, 180 days

Autogenous bone/HA/Zn-HA (powder)

0.5%

Zn-doping did not improve bone repair

Fernandes et al. (2009)

Rabbits n ¼ 5/ intramuscular

2, 4, 12 wks

HA/Zn-HA (granules, sintered)

2 wt%

Zn-doping did not induce osteoinduction

Nascimento et al. (2011)

Rabbits n ¼ 6/ calvaria (8 mm)

12 wks

nHA/Zn-HA (8 mm in diameter disc, calcinated)

0.000, 0.063, 0.316, 0.633 wt%

Zn-doping did not improve bone repair

Suruagy et al. (2016)

Rabbits n ¼ 6/femur

4 wks

ZnTCP/(ZnTCP/HA) (cylinder 2.55  10 mm, sintered)

0.316 wt%

Zn-doping improved bone repair

Kawamura et al. (2000a)

Rabbits n ¼ 5/femur

2, 4, 6, 12, 24 wks

(TCP/HA)/(ZnTCP/HA) (cylinder 2.55  10 mm, sintered)

0.316 wt%

Zn-doping improved bone repair

Kawamura et al. (2000b)

Rabbits n ¼ 5/femur

2, 4, 6, 12, 24, 60 wks

(TCP/HA)/(ZnTCP/HA) (cylinder 2.55  10 mm, sintered)

5%

Zn-doping improved bone apposition

Kawamura et al. (2003) 231

Continued

Zinc-substituted hydroxyapatite

Table 9.3 Animal models used, sites of implantation, experimentation duration, and composition of the calcium phosphate biomaterials used.

232

Table 9.3 Continued Animal model/Local of implantation

Experimental periods

Biomaterials

Zn %

Main conclusion

References

Rat n ¼ 10/tibia

8 wks

Empty/b-TCP/ZnTCP (powder)

0, 0.3, 0.9, 2.7 wt%

Zn-doping improved bone repair

Chou et al. (2013)

Canines n ¼ 4/ intramuscular

12 wks

TCP/ZnTCP (granules, sintered)

0, 3 wt%

Incorporation of zinc showed osteoinduction

Luo et al. (2014)

Rabbits n ¼ 5/tibia

1, 2 ,4 wks

HA/ZnHA (cylinder 2  6 mm)

0316 wt%

Zn-doping improved bone repair

Maia et al. (2010)

Rat n ¼ 6/femur

8 wks

1) Coll scaffold 2) MHMs/Coll scaffold 3) Zn5-MHMs/Coll scaffold

3.16 wt% ND

Zn-doping improved bone repair

Yu et al. (2017)

Handbook of Ionic Substituted Hydroxyapatites

Yu et al., 2017; Cruz et al., 2018; Calasans-Maia et al., 2014; Maia et al., 2008a, 2008b, 2010; Fernandes et al., 2009; Nascimento et al., 2011; Suruagy et al., 2016; Kawamura et al., 2000a, 2000b, 2003; Chou et al., 2013; Luo et al., 2014, copyright 2018 reproduced with the permission of Elsevier.

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References Anwar, A., et al., 2016. Novel continuous flow synthesis, characterization and antibacterial studies of nanoscale zinc substituted hydroxyapatite bioceramics. Inorg. Chim. Acta 453, 16e22. Ashuri, M., et al., 2012. Development of a composite based on hydroxyapatite and magnesium and zinc-containing solegel-derived bioactive glass for bone substitute applications. Mater. Sci. Eng. C 32 (8), 2330e2339. Boyer, L., Carpena, J., Lacout, J., 1997. Synthesis of phosphate-silicate apatites at atmospheric pressure. Solid State Ion. 95 (1e2), 121e129. Cacciotti, I., 2014. Cationic and anionic substitutions in hydroxyapatite. In: Handbook of Bioceramics and Biocomposites, pp. 1e68. Calasans-Maia, M., et al., 2014. Short-term in vivo evaluation of zinc-containing calcium phosphate using a normalized procedure. Mater. Sci. Eng. C Mater. Biol. Appl. 41, 309e319. Camaioni, A., et al., 2015. Silicon-substituted hydroxyapatite for biomedical applications. In: Hydroxyapatite (HAp) for Biomedical Applications. Elsevier, pp. 343e373. Chadda, H., et al., 2016. Cytotoxic evaluation of hydroxyapatite-filled and silica/ hydroxyapatite-filled acrylate-based restorative composite resins: an in vitro study. J. Prosthet. Dent. 116 (1), 129e135. Chou, J., et al., 2013. Bone regeneration of rat tibial defect by zinc-tricalcium phosphate (ZnTCP) synthesized from porous foraminifera carbonate macrospheres. Mar. Drugs 11 (12), 5148e5158. Cox, S.C., et al., 2014. Preparation and characterisation of nanophase Sr, Mg, and Zn substituted hydroxyapatite by aqueous precipitation. Mater. Sci. Eng. C 35, 106e114. Cruz, R., et al., 2018. Does the incorporation of zinc into calcium phosphate improve bone repair? A systematic review. Ceram. Int. 44 (2), 1240e1249. Cushing, B.L., Kolesnichenko, V.L., O’Connor, C.J., 2004. Recent advances in the liquid-phase syntheses of inorganic nanoparticles. Chem. Rev. 104 (9), 3893e3946. De Aza, P., et al., 1997. Vibrational properties of calcium phosphate compounds. 2. Comparison between hydroxyapatite and b-tricalcium phosphate. Chem. Mater. 9 (4), 916e922. Eshtiagh-Hosseini, H., Housaindokht, M.R., Chahkandi, M., 2007. Effects of parameters of solegel process on the phase evolution of solegel-derived hydroxyapatite. Mater. Chem. Phys. 106 (2), 310e316. Fadeeva, I., et al., 2012. Zinc- and Silver-Substituted Hydroxyapatite: Synthesis and Properties, vol. 442, pp. 63e65. Fernandes, G., et al., 2009. Histomorphometric analysis of bone repair in critical size defect in rats calvaria treated with hydroxyapatite and zinc-containing hydroxyapatite 5%, 396e398, 15e18. Fujii, E., et al., 2006. Selective protein adsorption property and characterization of nanocrystalline zinc-containing hydroxyapatite. Acta Biomater. 2 (1), 69e74. Gomes, P., 2011. Biological Evaluation of Biomaterials for Bone Tissue Regeneration. Gomes, S., Nedelec, J.-M., Renaudin, G., 2012. On the effect of temperature on the insertion of zinc into hydroxyapatite. Acta Biomater. 8 (3), 1180e1189. Grandjean-Laquerriere, A., et al., 2006. Influence of the zinc concentration of solegel derived zinc substituted hydroxyapatite on cytokine production by human monocytes in vitro. Biomaterials 27 (17), 3195e3200.

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Graziani, G., et al., 2017. Ion-substituted calcium phosphate coatings deposited by plasmaassisted techniques: a review. Mater. Sci. Eng. C 74, 219e229. Guerra-Lopez, J., et al., 2001. Influence of nickel on hydroxyapatite crystallization. J. Raman Spectrosc. 32 (4), 255e261. Guerra-Lopez, J.R., et al., 2015. Synthetic hydroxyapatites doped with Zn (II) studied by X-ray diffraction, infrared, Raman and thermal analysis. J. Phys. Chem. Solids 81, 57e65. Hahn, B.-D., et al., 2010. Aerosol deposition of silicon-substituted hydroxyapatite coatings for biomedical applications. Thin Solid Films 518 (8), 2194e2199. Hattori, Y., et al., 2016. Mechanochemical synthesis of zinc-apatitic calcium phosphate and the controlled zinc release for bone tissue engineering. Drug Dev. Ind. Pharm. 42 (4), 595e601. Ho, C.-C., Ding, S.-J., 2015. Novel SiO2/PDA hybrid coatings to promote osteoblast-like cell expression on titanium implants. J. Mater. Chem. B 3 (13), 2698e2707. Huang, W., et al., 2011. Alendronate decorated nano hydroxyapatite in mesoporous silica: cytotoxicity and osteogenic properties. Appl. Surf. Sci. 257 (23), 9757e9761. Ito, A., et al., 2000. Preparation, solubility, and cytocompatibility of zinc-releasing calcium phosphate ceramics. J. Biomed. Mater. Res. 50 (2), 178e183. Kawamura, H., et al., 2000. Stimulatory effect of zinc-releasing calcium phosphate implant on bone formation in rabbit femora. J. Biomed. Mater. Res. 50 (2), 184e190. Kawamura, H., et al., 2000. Effects of zinc-releasing calcium phosphate ceramics implanted in rabbit femora. In: Bioceramics, S.G., Moroni, A. (Eds.), pp. 387e390. Kawamura, H., et al., 2003. Long-term implantation of zinc-releasing calcium phosphate ceramics in rabbit femora. J. Biomed. Mater. Res. A 65A (4), 468e474. Kaygili, O., Tatar, C., 2012. The investigation of some physical properties and microstructure of Zn-doped hydroxyapatite bioceramics prepared by solegel method. J. Sol. Gel Sci. Technol. 61 (2), 296e309. Khan, A.F., et al., 2014. Bioactive behavior of silicon substituted calcium phosphate based bioceramics for bone regeneration. Mater. Sci. Eng. C 35, 245e252. Leshkivich, K., Monroe, E., 1993. Solubility characteristics of synthetic silicate sulphate apatites. J. Mater. Sci. 28 (1), 9e14. Luo, X., et al., 2014. Zinc in calcium phosphate mediates bone induction: in vitro and in vivo model. Acta Biomater. 10 (1), 477e485. Ma, J., Qin, J., 2015. Graphene-like zinc substituted hydroxyapatite. Cryst. Growth Des. 15 (3), 1273e1279. Maia, M., et al., 2008. Effect of hydroxyapatite and zinc-containing hydroxyapatite on osseous repair of critical size defect in the rat calvaria, 361e363, 1273e1276. Maia, M., et al., 2008. Stimulatory effect on osseous repair of zinc-substituted hydroxyapatite. In: Histological Study in Rabbit’s Tibia, vols. 361e363, pp. 1269e1272. Maia, M., et al., 2010. Bone Implant Interface Investigation by Synchrotron Radiation X-Ray Microfluorescence, vol. 1221, pp. 160e163. Miao, S.D., et al., 2007. In vitro bioactivity and osteoblast-like cell test of zinc containing fluoridated hydroxyapatite films. J. Mater. Sci. Mater. Med. 18 (10), 2101e2105. Miyaji, F., Kono, Y., Suyama, Y., 2005. Formation and structure of zinc-substituted calcium hydroxyapatite. Mater. Res. Bull. 40 (2), 209e220. Naqshbandi, A., et al., 2014. Sol-Gel Synthesis of Zn Doped HA Powders and Their Conversion to Porous Bodies, vol. 493, pp. 603e608. Nascimento, L., et al., 2011. Osseoinduction Evaluation of Hydroxyapatite and Zinc Containing Hydroxyapatite Granules in Rabbits, vols. 493e494, pp. 252e257.

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Negrila, C.C., et al., 2018. Development of zinc-doped hydroxyapatite by sol-gel method for medical applications. Molecules 23 (11). Pietak, A.M., et al., 2007. Silicon substitution in the calcium phosphate bioceramics. Biomaterials 28 (28), 4023e4032. Predoi, D., et al., 2017. Textural, structural and biological evaluation of hydroxyapatite doped with zinc at low concentrations. Materials 10 (3), 229. Predoi, D., et al., 2019. Zinc doped hydroxyapatite thin films prepared by sol-gel spin coating procedure. Coatings 9 (3), 18. Qiao, Y., et al., 2014. Stimulation of bone growth following zinc incorporation into biomaterials. Biomaterials 35 (25), 6882e6897.  et al., 2012. Biocompatibility and antimicrobial activity of zinc (II)-doped Radovanovic, Z., hydroxyapatite, synthesized by a hydrothermal method. J. Serb. Chem. Soc. 77 (12), 1787e1798. Ramya, J.R., et al., 2014. Physicochemical and biological properties of iron and zinc ions codoped nanocrystalline hydroxyapatite, synthesized by ultrasonication. Ceram. Int. 40 (10), 16707e16717. Part B). Rao, B.G., Mukherjee, D., Reddy, B.M., 2017. Chapter 1 e novel approaches for preparation of nanoparticles. In: Ficai, D., Grumezescu, A.M. (Eds.), Nanostructures for Novel Therapy. Elsevier, pp. 1e36. Rau, J.V., et al., 2015. Bioactive, nanostructured Si-substituted hydroxyapatite coatings on titanium prepared by pulsed laser deposition. J. Biomed. Mater. Res. B Appl. Biomater. 103 (8), 1621e1631. Rehman, I., Bonfield, W., 1997. Characterization of hydroxyapatite and carbonated apatite by photo acoustic FTIR spectroscopy. J. Mater. Sci. Mater. Med. 8 (1), 1e4. Ren, F., et al., 2009. Characterization and structural analysis of zinc-substituted hydroxyapatites. Acta Biomater. 5 (8), 3141e3149. Resende, N.S., Nele, M., Salim, V.M.M., 2006. Effects of anion substitution on the acid properties of hydroxyapatite. Thermochim. Acta 451 (1), 16e21. Shepherd, J.H., Shepherd, D.V., Best, S.M., 2012. Substituted hydroxyapatites for bone repair. J. Mater. Sci. Mater. Med. 23 (10), 2335e2347. Shepherd, D.V., et al., 2014. An in vitro study into the effect of zinc substituted hydroxyapatite on osteoclast number and activity. J. Biomed. Mater. Res. A 102 (11), 4136e4141. Sogo, Y., et al., 2004. Zinc containing hydroxyapatite ceramics to promote osteoblastic cell activity. Mater. Sci. Technol. 20 (9), 1079e1083. Stojanovic, Z., et al., 2009. Hydrothermal synthesis of nanosized pure and cobalt-exchanged hydroxyapatite. Mater. Manuf. Process. 24 (10e11), 1096e1103. Sun, G., Ma, J., Zhang, S., 2014. Electrophoretic deposition of zinc-substituted hydroxyapatite coatings. Mater. Sci. Eng. C 39, 67e72. Suruagy, A.A.P.D.S., et al., 2016. Physico-chemical and histomorphometric evaluation of zinccontaining hydroxyapatite in rabbits calvaria. Braz. Dent. J. 27, 717e726. Sych, O., et al., 2015. Si-modified BHA bioceramics as a drug delivery system: effect of modification method on structure and Rifampicin release. Process. Appl. Ceram. 9 (3), 125e129. Szurkowska, K., Kolmas, J., 2017. Hydroxyapatites enriched in siliconeBioceramic materials for biomedical and pharmaceutical applications. Prog. Nat. Sci. Mater. Int. 27 (4), 401e409. Thian, E., et al., 2013. Zinc-substituted hydroxyapatite: a biomaterial with enhanced bioactivity and antibacterial properties. J. Mater. Sci. Mater. Med. 24 (2), 437e445.

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Thian, E.S., et al., 2013. Zinc-substituted hydroxyapatite: a biomaterial with enhanced bioactivity and antibacterial properties. J. Mater. Sci. Mater. Med. 24 (2), 437e445. Uysal, I., et al., 2014. Co-doping of hydroxyapatite with zinc and fluoride improves mechanical and biological properties of hydroxyapatite. Prog. Nat. Sci. Mater. Int. 24 (4), 340e349. Vladescu, A., et al., 2014. Enhancement of the mechanical properties of hydroxyapatite by SiC addition. J. Mech. Behav. Biomed. Mater. 40, 362e368. Wang, J., et al., 2010. Study of synthesis of nano-hydroxyapatite using a silk fibroin template. Biomed. Mater. 5 (4), 041002. Wang, Q., et al., 2018. Experimental and simulation studies of strontium/zinc-codoped hydroxyapatite porous scaffolds with excellent osteoinductivity and antibacterial activity. Appl. Surf. Sci. 462, 118e126. Xiao, X., et al., 2008. Structural characterization of zinc-substituted hydroxyapatite prepared by hydrothermal method. J. Mater. Sci. Mater. Med. 19 (2), 797e803. Yamaguchi, M., 1998. Role of zinc in bone formation and bone resorption. J. Trace Elem. Exp. Med. 11 (2-3), 119e135. Yamaguchi, M., Uchiyama, S., 2004. Receptor activator of NF-kB ligand-stimulated osteoclastogenesis in mouse marrow culture is suppressed by zinc in vitro. Int. J. Mol. Med. 14 (1), 81e85. Yamaguchi, M., Oishi, H., Suketa, Y., 1987. Stimulatory effect of zinc on bone formation in tissue culture. Biochem. Pharmacol. 36 (22), 4007e4012. Yu, W., et al., 2017. Evaluation of zinc-doped mesoporous hydroxyapatite microspheres for the construction of a novel biomimetic scaffold optimized for bone augmentation. Int. J. Nanomed. 12, 2293. Zhu, M., et al., 2014. Mesoporous silica nanoparticles/hydroxyapatite composite coated implants to locally inhibit osteoclastic activity. ACS Appl. Mater. Interfaces 6 (8), 5456e5466.

Silver-substituted hydroxyapatite 1

2

3

10

Zohaib Khurshid , Muhammad Sohail Zafar , Shehriar Hussain , Amber Fareed 4 , Safiyya Yousaf 5 , Farshid Sefat 6 1 Department of Prosthodontics and Dental Implantology, College of Dentistry, King Faisal University, Saudi Arabia; 2Department of Restorative Dentistry, College of Dentistry, Taibah University, Saudi Arabia; 3Department of Dental Materials, College of Dentistry, Jinnah Sindh Medical University, Karachi, Pakistan; 4Department of Preventive Dentistry, Oman Dental College, Muscat, Oman; 5Department of Biomedical and Electronics Engineering, School of Engineering, University of Bradford, Bradford, United Kingdom; 6Interdisciplinary Research Center in Polymer Science & Technology (IRC Polymer), University of Bradford, Bradford, United Kingdom

10.1

Introduction

This chapter describes the rationale of synthesis and characterization of sliversubstituted apatite in terms of their physical, chemical, and biological characteristic for bone ingrowth. There are three main constituents of bones and teeth, which are living cells, extracellular collagenous matrix, and the calcium phosphate mineral part also known as hydroxyapatite (HA) or Ca5(PO4)3OH. Therefore, the ideal replacement of bone and teeth would be any combination of the above components. Currently, the best materials for these replacements are titanium alloys and calcium phosphate ceramics (Bir et al., 2012; Khurshid et al., 2016). Titanium alloys are widely used due to their strength, biocompatibility, and corrosion resistance (Najeeb et al., 2016), whereas calcium phosphates have shown excellent compatibility with bone due to the favorable biological responses and improved boneeimplant adhesion, and they can also be a scaffold for bone growth. Nanosized silver particles contain more active surface sites and are chemically durable, which are prepared by the chemical reduction of silver salts by sodium citrate or sodium borohydride, and utilize silver nitrate as the source for Ag (Russell and Hugo, 1994). Previously, AgNPs were used as antibacterial agents in the food and food storage (Chowdhury et al., 2016; Kumar et al., 2018), health sector (Sodagar et al., 2016), cosmetics (Domeradzka-Gajda et al., 2017) textile industry, and in environmental applications (Syafiuddin et al., 2017).

10.2

The rationale of silver in apatite

Apatite is one of the main constituents of a teeth and bones; therefore, a significant amount of research on the synthetic preparations of apatite as a bone replacement biomaterial, implant coatings, and drug delivery is reported in the literature

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00010-0 Copyright © 2020 Elsevier Ltd. All rights reserved.

238

Handbook of Ionic Substituted Hydroxyapatites

e (Owens et al., 2019; Palazzo et al., 2007). Apatite is formed by Caþ2, PO3 4 , and OH ions, and these ions are required in the correct orientation and in enough numbers to form the stable apatite crystal. Synthetic biological apatite is characterized by the presence of several foreign ions, which play an important metabolic role. Moreover, HA has the ability of anionic substitution (carbonate, fluoride, and silicate, or cationic) and cationic substitution (magnesium, strontium, zinc, iron, and silver) to enhance the therapeutic bone formation. Therefore, apatite substituted with relevant ions is regarded as attractive biomaterials for hard tissue substitution/repair (Sefat et al., 2019; Sheikh et al., 2015, 2014; Zafar et al., 2019, 2015) (see Fig. 10.1). Silver (Ag), in this regard, is of particular interest as it has been long recognized for inhibitory effect toward several pathogenic bacteria and microorganisms (Alexander, 2009). For centuries, it is known that silver ions exhibit strong inhibitory effects toward a broad spectrum of bacterial strains (Clement and Jarrett, 1994). Ag in its many oxidation states (Ag0, Agþ, Ag2þ, and Ag3þ) had been incorporated in many biomaterials including HA. It is desirable for the implants to have antibacterial properties to reduce the treatment duration by providing localized antibacterial effect, thus decreasing the need and the side effects of systemic therapies thereby improving its efficacy. Silver, an inorganic antibacterial agent, is historically known for its broad-spectrum antimicrobial activities against bacteria, viruses, algae, and fungi (Gosheger et al., 2004;

Repair and reconstruction surgery

Conservative dentistry

Pharmacy

Biomedical applications of hydroxyapatite (HA) Bone defect filling materials

Orthopedics materials

Implant surface coatings

Figure 10.1 Application of hydroxyapatite (HA) in the human body.

Silver-substituted hydroxyapatite

239

Martínez-Gutierrez et al., 2012). Table 10.1 shows the several mechanisms of action by silver ions to interfere the cellular growth. Additionally, silver has the potential for the development of novel antimicrobial agents, formulations of drug delivery, biomaterial coatings, regeneration materials, and improved therapeutics. (Khurshid et al., 2017; Zafar et al., 2019, 2017; Zhao et al., 2009). The antibacterial mechanism of silver is related to the release of Agþ ions from Ag nanoparticles (AgNPs) to disrupt membrane permeability or by direct contact with bacterial result in microbial death (Agnihotri et al., 2013; Fu et al., 2016). The oxidation and release of Agþ in Ag-HA coating is essential for antibacterial activity; therefore, a mild heat treatment (170 C in air for 8 h) process of Ag-HA coating enhanced the consistent antimicrobial action of AgNP (Zhang et al., 2017). Recent studies suggested that the biological responses of silver-substituted hydroxyapatite (Ag-HA) may be improved with the addition of silicon (Si) because presence of Si in the HA structure is related to the rate of dissolution, precipitation, and biomineralization mechanisms (Zhang et al., 2017). The excellent bioactivity of sliver particles was experimentally evaluated against bacteria, fungi, viruses, and yeast. To achieve the therapeutic effects, incorporation of silver into the apatite is done by several methods that will be discussed later including coprecipitation process and ion exchange with the calcium ions in the apatite, wet precipitation reaction between calcium phosphate and orthophosphoric acid, and hydrothermal method (Kometani and Teranishi, 2010; Li et al., 2015). It is possible that the bioactivity and controlled release of ions from the HA is influenced by silver dispersion. Mostafa et al. (2010) reported the physical characteristics of Ag-HA composite using AgNPs and claimed that the release of Ca and P in Ag-HA composite can be controlled in presence of silver. More recently, Ruiz-Baltazar reported the green synthesis of AgNPs from biosynthesis on Melissa officinalis extract to develop Ag-HA nanocomposite. The results of this study showed that the impregnation process of AgNP was successful, and greater interaction of AgNPs with the HA matrix in Ag-HA composite was observed (see Fig. 10.2) (Ruíz-Baltazar et al., 2018). Lim et al. discussed the incorporation of silver ions in the apatite structure during the coprecipitation process (Ag-HA-CP) or underwent ion exchange with calcium ions in the apatite (Ag-HA-IE) and reported the chemical, physical, antibacterial properties, and biological responses of Ag-HA-CP and Ag-HA-IE (Lim et al., 2015). They claimed that the antibacterial action of Ag-HA-IE was related to the released silver ions, whereas in Ag-HA-CP, it was dependent on the surface-bound silver ions as shown in Fig. 10.3. Therefore, antibacterial efficacy of Ag-HA-CP is possibly achieved for a longer time compared with Ag-HA-IE, and a silver content between 0.5e2 wt.% in Ag-HA-CP may achieve effective antibacterial effect (Lim et al., 2015).

10.3

Substitution in hydroxyapatite

HA, (Ca10(PO4)6(OH)2), generally used for compensating bone defects and dental complications, is one of the basic material constituents of metallic

(a)

(b)

1.00kV-D 3.8mm x90.0k SE(T)

500 nm

(c)

1.00kV-D 3.8mm x350k SE(T)

100 nm

(d)

1.00 µm

SU8230 3.0kV 3.4mm x35.0k SE(U)

SU8230 3.0kV 3.4mm x35.0k HA(T)F0

1.00 µm

Figure 10.2 (a and b) Hydroxyapatite (HA) scanning electron microscopy (SEM) micrographs obtained by the secondary electrons detector, (c and d) SEM images of the Ag-HA nanocomposite.
Adapted from Ruíz-Baltazar, A.J., Reyes-L
opez, S.Y., Silva-Holguin, P.N., Larra~ naga, D., Estévez, M., Pérez, R., 2018a. Novel biosynthesis of Ag-hydroxyapatite: structural and spectroscopic characterization. Results Phys. 9, 593e597. https://doi.org/10.1016/J.RINP.2018. 03.016, with permission.

AgHA-CP

AgHA-IE Released Ag+ ions damage cell wall

Surface-bound Ag+ ions damage cell wall

Bacteria

Bacteria Crystal surface

Crystal surface Ag+ ions diffusion

Crystal structure of AgHA-CP

Crystal structure of AgHA-IE

Figure 10.3 Graphical representation of substitution of silver into nanosized silver-substituted hydroxyapatite (nAg-HA) is employed to create antibacterial properties. Silver ions were either incorporated during the coprecipitation process (Ag-HA-CP) or underwent ion exchange with calcium ions in the apatite (Ag-HA-IE). Adapted with permission from Lim, P.N., Chang, L., Thian, E.S., 2015. Development of nanosized silver-substituted apatite for biomedical applications: a review. Nanomed. Nanotechnol. Biol. Med. 11, 1331e1344. https://doi.org/10.1016/j.nano.2015.03.016.

Silver-substituted hydroxyapatite

241

SeO3 substituted HA

Ag+ substituted HA

Cu2+ substituted HA Substitution of HA with different biomaterials for antibacterial property

Zn2+ substituted HA

Sr2+ substituted HA

Ce3+ and Eu3+ substituted HA

Figure 10.4 Different materials substitution in the hydroxyapatite (HA) to enhance the antibacterial activity and strengthen it.

implants (Lyasnikova et al., 2016). Since the 1980s, the relative newness of HA layers remains a topic of substantial interest, especially with the ease of atomic doping or substitution. Thereby, when substituted with ions such as carbonate ion or magnesium, nanocrystalline HA develops similar characteristics to biological apatite, the dominating element of vertebrate tissues (Kolmas et al., 2017) (see Fig. 10.4). With wide-ranging biomedical applications, HA can be used as a microbial agent (Swetha et al., 2012), drug delivery system (Lin et al., 2011), and biomarker for potential nanomedical platforms (Zuo et al., 2012). Possessing excellent biocompatibility, HA exhibits a surface chemistry that supports bone ingrowth. However, when modified (in particular, with Ag-HA), high levels can be toxic, leading to damage of human biological tissues and argyrosis (Manshian et al., 2015).

10.4

Methods of Preparations

Special attention is paid to synthetic HA precursors by various ceramic processing routes including wet precipitation, coprecipitation, hydrothermal processing, solegel, etc. Production of these HA powders have stimulated academic and industrial research for several heterogeneous catalysis applications. Within each division, there are multiple variations that are dependent on the conditions of synthesis and reagents used (Fihri et al., 2017).

242

10.4.1

Handbook of Ionic Substituted Hydroxyapatites

Wet precipitation

The most extensively used technique for the formation of HA power is the wet chemical method by either simple precipitation or hydrolysis. Wet precipitation can be performed at ambient or elevated temperatures using water or inorganic solvents. Highly dependent on the nucleationeaggregationeagglomerationegrowth mechanism, the technique releases harmless by-products (water) and is established for its clinical simplicity (Yelten-Yilmaz and Yilmaz, 2018). Reactions are conducted under atmospheric or high pressures offering control over the morphology and texture, thereby leading to a high yield of HA. Beginning with precipitation, the procedure is followed by aging, filtration, drying, and lastly heat treatment. The potential to apply diverse reagents, auxiliary additives, and apparatus enables tailorable properties on the synthesis method. For instance, these are determined by the Ca/P ratio, pH, and ripening time. To achieve reproducible HA powders, process parameters such as reactant addition rate, concentration, drying, and heat treatment conditions are essential for producing similar features of particle shape, particle size, and stoichiometry (Gentile et al., 2015; Mostafa, 2005). A lower addition rate of acid would promote larger sized particles (Ramesh et al., 2015). Being easily modified, the wet precipitation technique synthesizes HA/Ag and is usually prepared from Ca(NO3)2$4H2O, AgNO3, NH4OH and (NH4)2HPO4 (Badrour et al., 1998; Chen et al., 2010; Singh et al., 2011) or Ca(OH)2, AgNO3, and H3PO4 (Kim et al., 1998; Lim et al., 2013; Rameshbabu et al., 2007). Although wet chemical precipitation serves an economical benefit for use on an industrial scale (Abidi and Murtaza, 2014), a major drawback to this approach includes the presence of potential impurities due to various ions manifested in aqueous solution; therefore, it is incorporated into the crystal structure. Additionally, this can give rise to structures that are not crystallographically pure (Fihri et al., 2017).

10.4.2

Coprecipitation

Coprecipitation is a chemical process employed to prepare HA powders using low operating temperatures to generate high yields of pure products (Fihri et al., 2017; Ikoma et al., 1999). This simplicity of experimental methods has drawn attention to the field of biomedical applications (Kong et al., 2002). First proposed by Hayek and Newesely in 1963, coprecipitation is generally conducted by pH values ranging from 3 to 12 with the ability to perform in the presence of templates (Hayek et al., 2007). Considering the technique is variable dependant, various reagent and additives can be used (Fihri et al., 2017).

10.4.3

Hydrothermal

The hydrothermal process is a rather mature technique for developing complex oxide powders with high crystallinity in a confined environment (Zhang et al., 2011). The procedure requires a high temperature and pressure superior to that of the ambient pressure inside a pressure vessel (or autoclave). These determine the medium used,

Silver-substituted hydroxyapatite

243

such as subcritical or supercritical. Chemical bonds are formed through the effect of condensation and reactivity, thus generating nuclei that ensure stoichiometric and highly crystalline synthesis of HA, respectively (Fihri et al., 2017). Morphology and porosity are controlled with high pressure creating microsized crystallites, thereby serving to modulate interactions between the solid and solvent. Commonly, this method is combined with conventional methods, coprecipitation protocols, etc (Abdel-Aal et al., 2008; Fihri et al., 2017; Zhu et al., 2009).

10.4.4 Solegel The solegel synthesis method involves mineralization from precursors in a solution to create colloidal sol, which eventually forms a gel-like compound (Kara et al., 2005). This exemplary process requires no energy conditions (Ramesh et al., 2015) though it entails strict control parameters and is highly reliant on (1) temperature and pH; (2) the nature of the solvent; and (3) the chemical nature of reagents used (Fihri et al., 2017). A process comprising a nonalkoxide solegel for HA synthesis has been established only for the conventional sources of calcium and phosphate, without performing adjustments to the pH (Kim and Kumta, 2004; Rajabi-Zamani et al., 2008). Additionally, an intimate contact is required at molecular level to produce nanocrystalline powders with homogenous composition, high purity, and overall a sophisticated nanomaterial suitable for tailoring various applications (Kalaiselvi, 2017; Uskokovi
c and Uskokovi
c, 2011). The protocol involves generating micelles around templates in either an aqueous or organic phase to hydrolyze the precursors, followed by polycondensation via the formation of a 3D inorganic network (Chen et al., 2011; Velu and Gopal, 2009). The synthesis of these ceramic materials are generally conducted at room temperature, while the solegel films, according to Lim et al. (2001), are more homogenous and smooth at 300 C. Ultimately, the low-temperature nature of the solegel technique (including heat treatment for drying, calcining, etc.) offers major advantages to restrain effectively the formation of the amorphous phase. Being easily applicable, the process does not require a high vacuum or refined equipment (Wang et al., 2008). Despite gaining considerable popularity, there are potential limitations that hinder the technique’s expansion on an industrial scale including high cost and scarcity of solegel methods (alkoxide-based precursors) as well as time-consuming process control (Fihri et al., 2017).

10.4.5 Microwave Because of enormous progress in offering alternatives to heating techniques, a continued interest remains widespread among microwave-assisted synthesis. The method generates an increased field of perfectly crystalline powder, particularly homogenous in terms of size, morphology, and porosity (Fihri et al., 2017; Tang et al., 2009). Recent studies have presented surfactants such as cetyltrimethylammonium bromide (CTAB), ethylenediaminetetraacetic acid, and sodium dodecyl sulfate to control parameters of HA nanostructures (Arami et al., 2009; Luk and Abbott, 2002). Molecular agitation is initiated by a purely thermal origin that is caused

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by the inversion of dipole with alternations in the electric field, whereas in an electrostatic origin, interactions such as dipoleedipole between polar molecules and the electric field are performed. These are considered as contributing factors of microwave-assisted preparation that encounter a direct effect on the kinetics and activation energy (Fihri et al., 2017). Overall, this approach has evidently presented to be simple and economical to prepare nanosized materials with narrow particle size distribution (Guha et al., 2010).

10.4.6

Other methods

Among the different synthesizing techniques for HA, the study of its properties and ability to mimic biological systems remains on the economic front for adopting efficient functionality. Preparation, shape selectivity, and characterization are vital for a good understanding of solid catalysts, as well as morphology and kinetics. Other techniques include hydrolysis (Sturgeon and Brown, 2009) and alternate energy input methods to obtain HA powder (Han et al., 2004).

10.5

Use of surfactants during preparation

To obtain control over Ag-HA morphology, several macromolecules, monosaccharides, and related molecules are explored (Ruíz-Baltazar et al., 2018). For example, CTAB is extensively used in many aqueous synthetic methods (Wang et al., 2006). The X-ray diffraction (XRD) is a powerful nondestructive technique used to determine crystal and molecular structures. Characterization of crystal orientations (texture) and other structural parameters, such as average grain size, strain, and crystal defects, can be observed to identify various diffraction patterns (Poralan et al., 2015). Fouriertransform infrared spectroscopy (FTIR) is an analytical measurement technology that is routinely applied to the characterization of biomaterials (Rehman and Bonfield, 1997). FTIR analyzer functions by simultaneously collecting data from the entire infrared spectrum, which remains a current interest especially for examining tissue sections as an alternative to conventional histopathology. Preparing samples is often a tedious process and problematic for solid materials that are too opaque in their normal form; therefore, they require a reduced optical density using various sampling techniques (Gadaleta et al., 1996). Additionally, Raman spectroscopy is a valuable tool in the field of vibrational spectroscopy as it gathers important information on the nature of chemical bonding in material. Ionita (2009) used this method to study the changes in HA composition obtained from dental hard tissues in various pathological processes (Tavaf et al., 2017). At work with the complexity of HA, it is necessary to evaluate individual compositions for maintaining physicochemical properties and implementing appropriate methods accordingly.

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245

Structure of silver-substituted hydroxyapatite

Brittle in nature, HA belongs to a large family of isomorphic compounds and is substituted with metals (silver) to enhance its load-bearing capacity (Fihri et al., 2017; Shepherd et al., 2012). Substituting silver into HA has displayed considerable promise as numerous studies have observed and concluded a reduction in thermal stability while detecting an increase in the solubility of the generated apatite. Crystal lizing in the hexagonal system, the unit cell contains Ca2þ, PO2 4 , and OH groups  closely packed together, while OH groups serve as a backbone for HA (Poralan et al., 2015). HA can be defined as a compact assemblage of tetrahedral PO4 groups, where P5þ ions are present in the center of the tetrahedrons, thus occupied by four oxygen atoms. Moreover, each PO4 tetrahedron delimits two types of unconnected channels (Fihri et al., 2017). The substitution of AgNPs with HA through precipitation method with various Ca/P ratios resulted in the distribution of Ag in b-TCP homogeneously after sintering with higher solubility of Ag exhibited by b-TCP than HA and b-CPP (Gokcekaya et al., 2015).

10.7

Effect on bioactivity of hydroxyapatite

Bone can be amicably visualized as a dynamic composite material, boasting a crossstriated pattern of collagen fibrils interspersed with recurring amorphous zones (Glimcher and Muir, 1984; Mahamid et al., 2010). These are apatite crystal nucleation points, a prelude to crystal growth. Collagen is deemed an important synergistic regulator of inhibitors of the apatite nucleation process and subsequently exerts an active control over mineralization (Landis and Silver, 2009; Nudelman et al., 2010). The use of surface functionalized implantable fixtures for enhancing biomimetic capabilities in vivo is predicated on achieving balance between optimal mechanical function and a localized mass transport delivery system facilitating biological nutrient payload delivery and subsequent tissue regeneration capabilities at the implantation site (Hollister et al., 2002). The coating of load-bearing implants with HA has been a mainstay mode of surface activation of metallic fixtures destined for bone tissue for a variety of applications [6]. A chemical composition similar to bone, osteoconductivity, and good biocompatibility (Ogilvie et al., 1987; Park et al., 2010) are some prominent features characteristic of HA. Implant coating features deployed for this purpose include dipping (Zhang et al., 2006), plasma spraying (Sun et al., 2001), precipitation, and, the most commonly used, electrochemical deposition (Yan et al., 2014). The electrochemical deposition technique delivers good control coupled with a simple performance but lends itself to drawbacks such as weak bonding at the coatingemetal interface. Plasma electrolyte processing techniques have also been investigated to develop silver-substituted HA for coating on the surface of commercially pure titanium (Cp Ti) (Venkateswarlu et al., 2012). The coprecipitation technique allows for adsorption from solution and subsequent incorporation at the calcium sites in HA. This ion-exchange process can be achieved through various

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techniques, which include microwave processing (Rameshbabu et al., 2007), electrostatic spraying (Hwang et al., 2008), and HA nucleation (coprecipitation) (Oh et al., 2004). A slow diffusion index of Ag from within the Ag-HA crystal lattice appears to provide the most significant effect in terms of bioactivity as determined by soaking studies of samples and XPS of media determining Ag release profile from the lattice. This inevitably led to enhanced osteoblast-mediated activity when utilizing Ag-HA. Enhanced cell spreading, proliferation, and mineralization coupled with increased enzymatic activity of alkaline phosphatase indicative of elevated bone matrix deposition were some of the observable changes. Analyses conducted using TEM, FTIR, and XRD confirmed the close association of chemical makeup, morphology, and dimensions of Ag-HA particles with mineralized bone tissue. Moreover, the bioactive properties of silver-substituted HA have also been promising in terms of conducive proliferation of human mesenchymal stem cells on sterilized Ag-HA discs (Lim et al., 2015).

10.8

Use of micro- and/or nanosilver particles in hydroxyapatite for implant coatings

The use of HA for developing bone grafts and as a coating layer on the surface of an implant is a commonly employed tactic. However, this does not lend HA any exemption from assaults meted out by the host immune response in vivo. Medical implanterelated infections accounted for approximately 50% of infectious patients acquired in the hospital setting (Stamm, 1978). With an increase in the use of medical implants particularly in the elderly and the immune-compromized segments of the population, an accompanying incidence of implant-related infections is also predicted to surge even higher, leading to pain and suffering coupled with a significant rise in medical expenditures (Hetrick and Schoenfisch, 2006). Therefore, active strategies need to be harnessed for improving the period of service of the implant by reducing instances of bacterial contamination. The biological acceptance of an implant is predicated on a complex interplay between bacterial contamination on the implant surface and integration into the surrounding tissues (Gottenbos et al., 2002; Lin et al., 2017). For the implantation process to be predictably successful, the competition must lean in the favor of tissue integration before any appreciable level of bacterial contamination of the implant surface (before tissue integration) ensues. This is especially significant in subjects who are immune-compromized, have underlying systemic conditions, and/or the elderly. Adhesion of bacteria followed by the formation of a biofilm lies at the heart of implant failure owing to infective process (Zimmerli et al., 1998). The bacterial agents residing within the biofilm are extremely resistant to attack by (i) antibacterial agents, thereby rendering them, for the most part, ineffective and (ii) host defences, as the biofilm assumes the properties of a very effective shield dampening assaults on these fronts (Hetrick and Schoenfisch, 2006; Zimmerli et al., 1984). Recently, Tian et al. reported

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the deposition of 10e30 nm sized AgNPs on the surface of HA scaffold coating on Ti6Al4V to enhance Ag load efficiency and antibacterial activity (Tan et al., 2019). Considering this, the incorporation of HA-coated bone grafts with antibacterial properties will help in lowering the incidence of implant-related infections. However, to mitigate issues pertaining to infection in HA, a functional ion substitution in the guise of silver has gained significant traction (Rameshbabu et al., 2007; Venkateswarlu et al., 2012). Silver ions have a propensity for binding to enzymes and proteins within bacterial cells and disrupting its integrity, which plays havoc with basic cellular processes such as respiration, oxidation, and enzyme signaling, ultimately leading to bacterial cell death.

10.9

Antibacterial effect of silver-substituted hydroxyapatite

Even though HA is an integral component of the chemical composition of bone, it does not display any prowess as an antibacterial agent. Hence, in spite of their glaring merits, neat HA coatings for metallic implants lack effective antibacterial properties. A yearning for a better sterile environment at this interface persists. The ray of opportunity to be capitalized here lies in the trait of convenient substitution of Ca2þ in HA substructure by more effective metallic ion species, yielding enhanced osseointegration. The incorporation of silver ions (Agþ) into the HA substructure via Ca2þ substitution has shown to provide enhanced protection against the colonization of pathogenic microbial species at the site of implantation in vivo (Balamurugan et al., 2008). Silversubstituted HA has been the subject of a great deal of attention in contrast to other metallic ions for displaying strong antibacterial effects with a statistically significant broad-spectrum inhibitory range. To that end, a number of approaches have been put forward, which focus on Agþ ion recruitment in the HA substructure. These deliver a Ag-HA composite elaborating a combination of osseointegration and antibacterial component against organisms such as Escherichia coli (Iqbal et al., 2012). The introduction of 2.5 wt.% Ag in HA showed a significantly high index of antibacterial activity (Gopi et al., 2014). Another promising technique was conducted on surface of stainless steel substrates coated with Agþ doped fluorohydroxyapatite (Bir et al., 2012). Other investigators proposed the use of nanosilver particle doped HA, which demonstrated significant antibacterial activity. The spectrum of activity covered important Gram-positive and Gram-negative species. These included Pseudomonas species and E. coli in the Gram-negative category and Gram-positive Staphylococcus aureus (Iqbal et al., 2012). An interesting utilization of the antibacterial properties of silver was investigated (Yan et al., 2014). They used silverdoped HA/TiO2 nanotube coating via electrochemical deposition technique. When applied on a Ti-based substrate, the resulting material showed an improved level of interaction with host cells along with enhanced antibacterial capabilities against E. coli.

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Hence, by combining the biocompatible nature of HA with the antibacterial properties of Ag, these approaches offer suitable substitutes and convenient ways to design HA-based biomaterials with superior antibacterial properties and biocompatibility.

10.10

Drug-loaded silver-substituted hydroxyapatite

Nanoparticles have received great attention regarding selective and specific drug actions to improve healthcare due to enhanced drug delivery systems. There have been significant scientific contribution and improvements in boneeimplant interface and in the antibacterial activity by utilizing Ag-HA combination. In this regard, incorporation of AgNPs as a drug deliver carrier for cancer therapy, as antiinflammatory agent in local anesthesia, and to control antioxidant and antimicrobial activates is well established. Similarly, Saravanan et al. studied the antibacterial effects after the addition of AgNPs (80-120 nm) in a scaffold of chitosan and HA. However, silversubstituted HA exhibits better osteoblast adherence, and antibacterial activity in lower silver (0.5%) is considered suitable for drug delivery in implant applications. Nonetheless, surface modification and ion substitution in apatite and HA systems are based on silver as specific, selective, and versatile candidates for potential predictable drug delivery applications. The long-term silver ion release from functionalized silverdoped HA as local drug delivery agent was reported by Dubnika et al. The results of this study suggested that the Ag-HA scaffolds possess the antibacterial activity up to 1 year (Dubnika et al., 2014).

10.11

Other biomedical applications

HA, defined as a major member of bioceramic materials, is widely functional in clinical applications to recompense for defects of various etiology and size, i.e., the musculoskeletal system. Serving excellent osteoconductive and osteointegrative properties, HA uses in dental implants and orthopedic components have appeared to gain much research attention, owing practicality to modern medicine, chemistry, and biology (Gibson et al., 1999). According to literature, the incorporation of silver can enhance the toughness and strength of doped materials with the ability to promote porous material development, therefore contributing to the fulfillment of desirable properties (Nath et al., 2010). The sintering temperatures, the selected material preparation, and its phase composition among other parameters can be altered to fit various applications (Dubnika et al., 2014). Silver has been generally used for medical equipment and wound dressings, as it possesses a broad spectrum of activity against Gram-positive and Gram-negative bacteria, fungi, viruses, and protozoa. By inhibiting the electron transport chain of microorganisms, binding of Ag ions with enzymes, nucleic acids, and membranes helps to promote bioimplant surface modification

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such as implant-associated infection (Funao et al., 2016). Other applications for Ag-HA include periodontal treatment, maxillofacial surgery, alveolar ridge augmentation, and otolaryngology (Hench, 1991).

10.12

Dental applications

Although silver has been used efficiently and effectively in dentistry, it is controversial for this specific area of research, due to its variable toxicity in biological systems. However, silver is a major component in dental amalgam restoratives, the materials for dental barrier membranes used for efficient alveolar bone reconstruction, caries prophylaxis in the form of silver diamine fluoride, dental prostheses, restorative and endodontics, and implantology. Antibacterial Ag-HA could be used in clinical dental applications, both in orthodontics and restorative dentistry (Kolmas et al., 2014; Sivolella et al., 2012). Akhavan et al. studied the effect of Ag-HA nanoparticles on shear bond strength of dental adhesive and reported that 5% increase of Ag-HA shear bond strength offers appropriate antimicrobial and mechanical properties for orthodontic adhesive (Akhavan et al., 2013). Similarly, Sodagar et al. tested antimicrobial properties of Ag-HA incorporation in adhesives used for orthodontic application and showed that the experimental adhesive containing 5% Ag-HA is the superior most antimicrobial action against tests bacteria, while increasing the amount Ag-HA particles by more than 5% had an undesirable effect when compared with the control group (Sodagar et al., 2016). The addition of AgNPs and HA nanoparticles to dental porcelain was an effective method to decreasing the colonizing bacterial growth activity. However, Mohsen et al. (2015) claimed that there was decrease in the fracture strength and the color of dental ceramic after incorporation of silver HA nanoparticles in dental ceramics. Several agents were added to dental composites to control the secondary caries around resin composite fillings such as fluoride-releasing fillers (strontium fluoride and ytterbium trifluoride), quaternary ammonium, and silver-containing fillers (Itota et al., 2004; Nejatian et al., 2017; Syed et al., 2019). AgNPs were loaded in polydopamine (PDA)-coated hydroxyapatite (HAePDA) nanowires to form AgNPladen HA (HAePDAeAg) nanowires to achieve reinforcement and antibacterial effect in dental resin composites (Ai et al., 2017).

10.13

Conclusion

To sum up the above discussion, silver is playing a key role in the biomedical sciences as an active ingredient or we can say antibacterial agent. Its substitution with HA brings drastic changes in combating the microbes. We highlighted the different synthesis methods and applications in biomedical field.

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Park, D.-S., Kim, I.-S., Kim, H., Chou, A.H.K., Hahn, B.-D., Li, L.-H., Hwang, S.-J., 2010. Improved biocompatibility of hydroxyapatite thin film prepared by aerosol deposition. J. Biomed. Mater. Res. Part B Appl. Biomater. 94B https://doi.org/10.1002/jbm.b.31658 n/a-n/a. Poralan, G.M., Gambe, J.E., Alcantara, E.M., Vequizo, R.M., 2015. X-ray diffraction and infrared spectroscopy analyses on the crystallinity of engineered biological hydroxyapatite for medical application. IOP Conf. Ser. Mater. Sci. Eng. 79, 012028. https://doi.org/10. 1088/1757-899X/79/1/012028. Rajabi-Zamani, A.H., Behnamghader, A., Kazemzadeh, A., 2008. Synthesis of nanocrystalline carbonated hydroxyapatite powder via nonalkoxide solegel method. Mater. Sci. Eng. C 28, 1326e1329. https://doi.org/10.1016/J.MSEC.2008.02.001. Ramesh, S., Natasha, A.N., Tan, C.Y., Bang, L.T., Niakan, A., Purbolaksono, J., Chandran, H., Ching, C.Y., Ramesh, S., Teng, W.D., 2015. Characteristics and properties of hydoxyapatite derived by solegel and wet chemical precipitation methods. Ceram. Int. 41, 10434e10441. https://doi.org/10.1016/J.CERAMINT.2015.04.105. Rameshbabu, N., Sampath Kumar, T.S., Prabhakar, T.G., Sastry, V.S., Murty, K.V.G.K., Prasad Rao, K., 2007. Antibacterial nanosized silver substituted hydroxyapatite: synthesis and characterization. J. Biomed. Mater. Res. Part A 80A, 581e591. https://doi.org/10.1002/ jbm.a.30958. Rehman, I., Bonfield, W., 1997. Characterization of hydroxyapatite and carbonated apatite by photo acoustic FTIR spectroscopy. J. Mater. Sci. Mater. Med. 8, 1e4. https://doi.org/10. 1023/A:1018570213546. Rodríguez-Lorenzo, L.M., Vallet-Regí, M., Ferreira, J.M.F., Ginebra, M.P., Aparicio, C., Planell, J.A., 2002. Hydroxyapatite ceramic bodies with tailored mechanical properties for different applications. J. Biomed. Mater. Res. 60, 159e166. https://doi.org/10.1002/jbm. 1286.
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Iron-substituted hydroxyapatite

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Christie Lung Ying Kei Dental Materials Science, Applied Oral Science, Faculty of Dentistry, The University of Hong Kong, Hong Kong, 5/F, Prince Philip Dental Hospital, Sai Ying Pun, Hong Kong SAR, P.R.China

11.1

Introduction

Hydroxyapatite (HA) is structurally and chemically similar to the inorganic mineral component of mammalian bone and teeth. It is bioactive and biocompatible and has been widely used in orthopedics and dentistry as bone scaffolds, fillers implant coating, and drug, cell, and gene delivery agent (Szcze
s et al., 2017). There are two crystal structures for HA: hexagonal and monoclinic, both with a molar Ca to P ratio of 1.67. The lattice parameters for hexagonal HA are a ¼ b ¼ 9.542 Å, c ¼ 6.881 Å and g ¼ 120 degrees and a ¼ 9.421 Å, b ¼ 2a, c ¼ 6.881 Å and g ¼ 120 degrees for monoclinic HA (Posner et al., 1958; Elliott et al., 1973). The phosphate, PO34, tetrahedral is connected to Ca2þ, which is occupied on two different sites: (i) Ca (I) site and (ii) Ca (II) site. The Ca (I) site is the columnar Ca, and Ca (II) site is the screw axis Ca. For Ca (I) site, six oxygen atoms of PO3 4 groups are coordinated to Ca through the oxygen atoms. For Ca (II) site, the Ca is coordinated to six oxygen atoms of PO3 4 and one oxygen atom of hydroxyl group (Fig. 11.1) (Jiang et al., 2002). The major limitation of HA is the poor mechanical properties that are not favorable for load-bearing applications. The mechanical strength of HA is lower than natural bone (Lin et al., 2015). HA is relatively difficult to resorb in the physiological environment because of the low degradation rate. This affects their performance as scaffolds, drug carriers, and bone grafts because good degradability is required. However, HA bioceramics have good biocompatibility and high osteoconductivity. They still generally lack sufficient osteoinductive ability to induce osteogenic stem cells differentiation and osteoblasts and to trigger new bone formation. These are important for regeneration of large bone defects and bone tissue engineering (Lin et al., 2015). HA materials have also been applied as drug delivery agents. However, comparatively low drug-loading capacity, uncontrolled rate of drug release, low degradation rate, and the lack of targeting efficiency and labeling capacity are the problems. Furthermore, the synthesized HA nanoparticles form aggregates. This means only limited application in intravenous injection (Lin et al., 2015). The other limitation of HA is the very low antibacterial activity against pathogenic bacteria (Shi et al., 2015). Different techniques have been used to improve the every aspect of performance of HA materials (Marinea et al., 2018; Okada and Matsumoto, 2015). One approach is a substitution with other ions, which can change the physical, chemical, and biological Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00011-2 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Figure 11.1 Hydroxyapatite lattice structure: (a) Ca (I) sites and (b) Ca (II) sites. Reproduced from Jiang, M., Terra, J., Rossi, A.M., Morales, M.A., Saitovitch, E.M.B., Eillis, D.E., 2002. Fe2þ/Fe3þ substitution in hydroxyapatite: theory and experiment. Phys. Rev. B 66, 224107 with permission of The American Physical Society.  properties of HA. The Ca2þ, PO3 4 , and OH sites in HA can be substituted by other  cations and anions within the lattice (Supov
a, 2015). The substitution of other cations  occurs in the Ca2þ sites, whereas anionic substitution occurs in PO3 4 or OH sites. The substitution of other cations and anions induces lattice structural changes owing  to different ionic size and charge (Supov
a, 2015). Substitution of divalent cations such as Mg2þ and Zn2þ would not cause a charge imbalance in the lattice structure. Substitution of monovalent cations such as Kþ and Naþ would cause a charge imbalance. This charge imbalance can be neutralized by forming vacancies (or “holes”). The formation of vacancies would cause the distortion  of lattice (Supov
a, 2015).

11.2

Biological importance of iron

Iron is an essential component in human body. It mainly forms complex hemoglobin that functions as an oxygen transporter. It is also important for the heme enzymes formation and takes part in other electron transfer and redox (reductioneoxidation) reactions (Lieu et al., 2001), whereby iron takes part in bone metabolism. Along with mineral components, bone tissue consists of type I collagen. In the process of collagen synthesis, a-ketoglutarate, oxygen molecules, Fe (II), and a reducing agent are required (Laura Toxqui and Vaquero, 2015). Another mechanism is through the process of vitamin D metabolism, activation and deactivation. There are two steps of hydroxylation for vitamin D activation. These two steps are catalyzed by a hemebound iron complex, cytochrome P450. This step regulates calcium and phosphate adsorption and release in the kidney and bone. The second process deactivation is the oxidation of the hydroxylated vitamin D. This process is critical in regulating bone and mineral homeostasis (Jones et al., 2014).

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Iron deficiency occurs when losses of iron in the body or requirements exceed its adsorption. This is especially significant for women of childbearing age. The result can cause anemia and fatigue and lower work performance. Low iron level inhibited the process of osteogenesis. This could result in bone loss and increase the risk of a broken bone, i.e., osteoporosis (Zhao et al., 2012). On the other hand, excess iron intake would frequently result in osteoporosis and fractures, such as the thalassemias and hereditary hemochromatosis. It is due to the direct action of iron on bone or is the concomitant endocrine deficiencies, such as hypogonadism, invalid erythropoiesis, or other comorbidities that may influence bone metabolism. Recent in vitro and in vivo studies show that excess iron can affect bone formation and remodeling (Tsay et al., 2010). Owing to the biological importance of iron in human, the substitution of HA by iron would affect the physiochemical and biological properties of HA, which depends largely on the synthetic methods and the amount of iron incorporated.

11.3

Synthesis methods

Several methods have been developed to prepare iron-substituted HA: solegel, hydrothermal, ultrasonic irradiation, mist process, microwave irradiation, ultrasonice microwave method, biomimetic method, dry milling, and ion-exchange method. Different methods can affect the phase purity, homogeneity, crystallinity, and particle size. They can be classified as wet-state methods (solution) and solid-state methods. The solid-state methods are less frequently used.

11.3.1 Wet-state methods 11.3.1.1 Solegel process Solegel process is a series of hydrolysis and condensation reactions of a precursor followed by aging reaction (Fig. 11.2). First, the precursor undergoes hydrolysis to form hydrated intermediate, i.e., sol. Then, the hydrated products react with each other to form a chemical bond through condensation reaction. A gel is formed. Finally, 3D gel networks are formed by polycondensation of sol particles (Attia et al., 2002). The precursors are usually metal alkoxides or metal salts, e.g., chloride. The solegel process can be affected by different processing conditions: 1. acid and base catalystdthe rate of solegel process increases at acidic or alkaline pH and 2. ratio of reactants and solvent. The advantages for the solegel process include homogeneous mixing and low processing temperature, which is more energy-saving (Wang and Bierwagen, 2009). In the solegel synthesis of iron-substituted HA, calcium nitrate, diammonium hydrogen phosphate, and iron (III) nitrate are used. In a study, HA samples with four different Fe contents of 0, 0.7, 1.4, and 2.1 wt% were synthesized. The reaction mixture was stirred and heated. The dried product was calcined at high temperature of 800 C for 1 h. The scanning electron microscope images and energy-dispersive X-ray analysis showed that the sample particle size was less than 1 mm, and the detected iron

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Condensation

Hydrolysis

Sol

Drying

Calcination

Gel

Figure 11.2 Mechanism of a solegel process.

content increased with the amount of iron added (Fig. 11.3). From the X-ray powder diffraction (XRD) analysis, the product prepared was composed of mainly iron-doped HA and minor b-tricalcium phosphate (b-TCP). The HA/b-TCP ratio was found to be dependent on the doped amount of iron. The HA/b-TCP ratio decreased with increasing amount of substituted iron (Fig. 11.4). This may be related to a partial transformation of HA into a Fe-substituted b-TCP (Kaygili et al., 2014). The dielectric constant measured was increased with the iron concentration. This was attributed to the higher b-TCP phases in high content of iron-substituted HA, which accounted for the change in the polarization of HA samples. In another study, a range of iron-substituted HA with iron content of 0, 0.01, 0.05, 0.1 0.5, 1, 2, 3, 4, 5, and 10 mol% were prepared. The dried powder was annealed for 5 h at 800 C in air with subsequent annealing for another 10 h at 1000 C in air. From the XRD patterns, the iron-substituted HA had the hexagonal structure with space group P63/m. Some other phases such as b-TCP, CaO, Fe2O3, and calcium phosphate were also detected (Trinkunaite-Felsen et al., 2015).

11.3.1.2 Hydrothermal method Hydrothermal method is a chemical reaction that iron-substituted HA is synthesized under high temperature and high pressure conditions in water in a Teflon-lined autoclave. For solvothermal method, nonaqueous solvents are used as the reaction medium. The method can accelerate the hydrolysis (or solvolysis) reaction, followed by crystal growth in the solution. The structure, morphology, and size of the iron-substituted HA can be modified by changing the reaction parameters such as temperature, reaction time, pressure, reactant concentration, and filled volume of the autoclave (Rao et al., 2017). In this method, calcium nitrate, ammonium phosphate, and iron (III) chloride solutions with molar ratio of Ca/P 1.67 were mixed at room temperature in a Teflon vessel. The ratio of iron was varied in the range from 0 to 10 mol% with respect to the calcium

Iron-substituted hydroxyapatite

263

Figure 11.3 Scanning electron microscopy micrographs and energy-dispersive X-ray analysis reports of the pure hydroxyapatite (HA) and Fe-substituted HA samples. Reproduced from Kaygili, O., Dorozhkin, S.V., Ates, T., Al-Ghamdi, A.A., Yakuphanoglu, F., 2014. Dielectric properties of Fe doped hydroxyapatite prepared by solegel method. Ceram. Int. 40, 9395e9402 with permission of Elsevier.

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Indexed peaks : Hydroxyapatite : β-TCP

1.4FeHAp

20

25

30

35 40 2 (degree)

(113) (203) (222) (312) (213) (321) (410) (402) (004)

(311)

(310)

(112) (211) (202) (300) (301)

(200) (111)

HAp

(210)

0.7FeHAp

(002)

Intensity (a.u)

2.1FeHAp

45

50

55

Figure 11.4 X-ray diffraction pattern of iron-substituted hydroxyapatite (HA) with various doped amount of Fe. Reproduced from Kaygili, O., Dorozhkin, S.V., Ates, T., Al-Ghamdi, A.A., Yakuphanoglu, F., 2014. Dielectric properties of Fe doped hydroxyapatite prepared by solegel method. Ceram. Int. 40, 9395e9402 with permission of Elsevier.

nitrate solution. The pH was adjusted to 10 and maintained at this range. They were hydrothermally treated at 423 K for 24 and 72 h. Iron-substituted HA precipitates formed were filtered and dried at 323 K for overnight. From the XRD patterns, HA and a-Fe2O3 phases were detected. The proportion of a-Fe2O3 increased with iron content. The magnetization measured increased with the ratio of iron from 0 to 10 mol% (Sato and Nakahira, 2013).

11.3.1.3 Ultrasonic irradiation The basic principle of ultrasound is based on a generation of compression and expansion phases imparted by mechanical vibration. When the pressure generated exceeds the surface tension of a liquid, small vapor-filled voids, cavitation bubbles, are formed. The sound energy is concentrated in the cavitation bubbles produced. When the bubble absorbs more sound energy, the size of the bubble increases and finally exceeds the equilibrium radius. The bubble will implode and the liquid molecules will rush in. This generates high energy for chemical reactions because of an enormous local temperature and pressure (Wu et al., 2012). Iron cosubstituted with zinc HA nanopowder was prepared by an ultrasonic irradiation. Solutions of calcium nitrate, ammonium hydrogen phosphate, zinc chloride, and iron (III) chloride were mixed. The pH of the solution was adjusted to 10 by aqueous ammonia solution. The reaction mixture was under ultrasonic irradiation at 750 W for 1 h. The mixture is washed and the powder is dried at 80 C. Sodium hydroxide was

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265

added to raise the pH to 8, and black colloids were formed (Ramya et al., 2014). The XRD analysis revealed the HA phases with peak broadening in the iron-cosubstituted HA as the ionic radii of iron and zinc ions were smaller than Ca2þ. The dielectric constant measured for the iron cosubstituted with zinc HA increased with both the iron and zinc content.

11.3.1.4 Mist process Mist process is a method to prepare nanopowder and film under the atmospheric conditions. In the process, a precursor solution is atomized by an ultrasonic vibrator. The mist solution is transported by a carrier gas, e.g., argon, to a furnace. The atomized precursor is first dried and then heated to thermally decompose and finally underwent reactions to form the product (Sato and Nakahira, 2014). In the preparation process, solutions of calcium nitrate and iron (III) chloride were mixed, and pH was adjusted to 8 by tris solution. The pH of diammonium hydrogen phosphate was adjusted to 12 by sodium hydroxide solution. The solutions were mixed with the ratio of Ca to P from 1 to 2. The ratio of Fe to Ca was varied from 0 to 5 mol%. The mixed solution was atomized, where atomized precursors were transferred by argon gas to a heated glass tube to dry and react to form an iron-substituted HA. The temperature was varied in the range of 873e1073 K, and the processing time was 60 min (Sato and Nakahira, 2014). The XRD patterns for all samples have the HA phase as the main phase. Other minor phases of Ca(OH)2 and CaClOH phases were observed at the iron ratio of 2 and 5 mol%. The grain size distribution of ironsubstituted HA varied with the flow rate of the carrier gas at temperature of 973 K: 1.19 mm iron-substituted HA size at the 3 L/min flow rate and 1.05 mm ironsubstituted HA size at the 5 L/min flow rate.

11.3.1.5 Microwave irradiation A microwave is an electromagnetic energy with frequency from 300 MHz (1 m) to 300 GHz (1 mm). The heating effect of microwave irradiation is mainly due to dielectric polarization mechanism. When a polar molecule is irradiated with microwaves, it aligns itself with the applied field. The electric field is changing rapidly, and this affects the pole orientation of the molecule. Therefore, the molecule aligns itself continually with the changing electric field. Heat energy is generated as a result of frictional force. The thermal energy generated is dependent on the dielectric constant of the molecule. The larger the dielectric constant, i.e., more polar molecule, the greater is the coupling with microwaves. Microwave method has several advantages over conventional heating method: it is economic, environmental friendly, has fast and uniform heating, and has high efficiency in the heat consumption (Rao et al., 2017; Belwal, 2013). For microwave synthesis, calcium nitrate, ammonium phosphate, and iron (III) nitrate solutions were mixed and heated. The iron content is at 2.0 at% with a molar ratio of Fe3þ/Ca2þ ¼ 0.2. The pH of the reaction mixture was raised to 10e11 by adding ammonia hydroxide. The mixture was transferred to a PTFE thermal reactor and irradiated with microwave for 10 min at 1200 W (2450 MHz) and at an internal

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pressure of 2 MPa. The HA precipitate was separated from the suspension mixture by centrifugation, later washed with ethanol and dried. The dried gel was ground into fine powder and dried at 150 C for 2 h to obtain the iron-substituted HA. The XRD patterns showed the samples with the hexagonal structure. There was no monoclinic phase, but a minor phase of brushite (CaHPO4$2H2O) phase was detected
(Robles-Aguila et al., 2017).

11.3.1.6 Ultrasonicemicrowave method Recently, ultrasonic and microwave-assisted method is used for the synthesis of iron-substituted HA. This method utilizes the advantages of both techniques: smaller nanoparticle size, high surface area, short reaction time, narrow particle size distribution, and high purity. In this method, calcium nitrate, diammonium hydrogen phosphate, and iron (III) nitrate solutions were mixed with the molar ratio of (Ca þ Fe)/P at 1.67. A range of iron-substituted HA with the iron content (0, 1.5, 3, and 5 wt.%) was prepared. The pH was adjusted to 10 by ammonia, and the mixed solution was heated in an ultrasonicemicrowave synergistic extraction apparatus with the microwave power 300 KW for 800 s. The solegel was aged for 24 h at room temperature. The solegel was centrifuged with deionized water and ethanol. After centrifugation, the precipitate was dried for 48 h to obtain iron-substituted HA powder in a vacuum drying oven. The iron-substituted HA fine powder was pressed into cylindrical blocks which were then sintered using a muffle furnace at temperature of 1200 C for 2 h. The XRD pattern showed that the iron-substituted HA prepared was pure-phase HA. The crystallinity of the powder decreased with an increase in the iron content. The friction coefficient of the iron-substituted HA ceramics with the iron content 1.5 wt.% measured was lower than the HA ceramics. The friction coefficient of the iron-substituted HA ceramics was higher than HA with increase of iron content. The wear resistance of iron-substituted HA was improved with the iron substitution, but the wear resistance was not improved further when the iron content is 5 wt% (Han et al., 2017).

11.3.1.7 Biomimetic method Recently, a new method is developed to synthesize pure phase of iron-substituted HA with an optimum crystallinity and control over its crystal size. It is based on the sole gel technique with the addition of organic matrices such as synthetic or natural polymers. The organic matrices are added to establish a kinetic control on nucleation and growth of the crystals (Sheikh et al., 2018). An alkaline solution of calcium nitrate at pH 9.5 was mixed with an iron (III) nitrate and 0.5 wt% polyvinyl alcohol solutions. A nominal composition of (Ca þ Fe)/P ratio of 1.67 was fixed. The concentrations of iron were 0% (control), 2%, 4%, 6%, 8%, and 10%. The solution was equilibrated for 24 h, followed by mixing of diammonium hydrogen phosphate and ammonia solution. The iron-substituted HA precipitates were aged for a week, rinsed until the pH was 7, filtered, and dried in oven at 60 C. The XRD pattern confirmed the HA lattice structure with a decrease in lattice size by substitution of iron (Sheikh et al., 2018).

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For the thermal behavior of iron-substituted HA examined, there was a gradual decrease in weight loss when the iron content increases from 2% to 10%. This could be due to the smaller ionic radius of Fe3þ with respect to Ca2þ. The saturation magnetization increased with the iron concentration from 2% to 10%. All the concentrations of iron-substituted HA showed better performance than the HA in the biocompatibility test on human osteosarcoma (MG-63) cell lines. Antibacterial activity tests showed that iron-substituted HA exhibited good antibacterial performance against Grampositive and Gram-negative bacteria (Sheikh et al., 2018).

11.3.2 Solid-state methods 11.3.2.1 Dry milling Milling is a process in which mechanical energy is applied to milling balls to break down coarse particles to very fine particles and blend of particles to form new phases. The kinetic energy that transferred from the balls to the particles is affected by the nature and size of ball mills, ball to powder ratio, milling speed, milling temperature, and the processing time (Yadav et al., 2012). In the milling process for the synthesis of iron-substituted HA, calcium oxide, ammonium hydrogen phosphate, and iron (III) chloride powder were grounded in the sealed nylon vial using mm-sized zirconia balls. After milling, distilled water was added, and the milling process was continued. The powder was thermally treated at 400 C for 2 h to initiate the chemical reaction to form iron-substituted HA. The XRD patterns show the HA phase without any other calcium phosphate phases (Rau et al., 2014).

11.3.2.2 Ion-exchange method In this technique, the reversible interchange of ions between a solid and a liquid solution is occurred until equilibrium is established. The ions in the lattice are substituted and diffused into the solution. The ion-exchange rate depends on the ion solution concentration and the particle size of the solid powder (Kramer et al., 2013). The HA powder was immersed in iron (III) chloride solution at room temperature. The Ca2þ ions diffused into the solution and Fe3þ ions diffused into the HA lattice and substitute the Ca2þ ions site until an equilibrium was established: Ca(HA)

(s)

þ Fe3þ(aq) 5 Fe3þ(HA)

(s)

þ Ca2þ(aq)

The XPS results revealed there was minority of ironephosphate bonding in the iron-substituted HA powder material. Substitution was achieved using iron (III) chloride solution but not iron (II) chloride solution. Furthermore, the number of iron (III) ions substituted increased with the soaking time (Kramer et al., 2013).

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Figure 11.5 The lattice structure of iron (II)-substituted hydroxyapatite (HA) at the Ca (I) site. Reproduced from Jiang, M., Terra, J., Rossi, A.M., Morales, M.A., Saitovitch, E.M.B., Eillis, D.E., 2002. Fe2þ/Fe3þ substitution in hydroxyapatite: theory and experiment. Phys. Rev. B 66, 224107 with permission of The American Physical Society.

11.4

Lattice structure of iron-substituted hydroxyapatite

The Ca2þ sites can be substituted by cationic Fe2þ and Fe3þ. The difference in ionic charge density and cationic size of Fe2þ and Fe3þ would induce a deformation of the HA lattice structure. The electronic configuration of Fe is 1s2 2s2 2p6 3s2 3p6 3d6 4s2. For Fe2þ substitution, there is no change of stoichiometry of HA, but there is a lattice relaxation as the charge density of Fe2þ is higher than Ca2þ (larger nuclear charge of Fe2þ). The Fe2þeO bond length ranges from 1.94 to 2.27 Å, which is shorter than that of Ca2þeO bond length from 2.36 to 2.51 Å. The Fe2þ substitution at the Ca(II) sites are energetically favored over Ca(I) sites using a LennardeJones potential mathematical model (Jiang et al., 2002). The most stable geometry is the sixfold coordination Ca (II) site (Fig. 11.5). For Fe3þ substitution, the ionic radius of Fe3þ (0.64 Å) is smaller than Ca2þ (0.99 Å). The Fe3þeO bond length ranges from 1.80e2.14 Å. There is a decrease of unit cell volume. The most stable geometry is the substitution at the fourfold Ca (I) sites over the Ca (II) sites using the LennardeJones potential mathematical model (Fig. 11.6) (Jiang et al., 2002).

11.5

Physical properties of iron-substituted hydroxyapatite

The physical properties of HA can be changed by the iron substitution at the Ca2þ ion sites. Some of the physical properties will be discussed in this section.

Iron-substituted hydroxyapatite

269

Figure 11.6 The lattice structure of iron (III)-substituted hydroxyapatite (HA) at the Ca (II) site. Reproduced from Jiang, M., Terra, J., Rossi, A.M., Morales, M.A., Saitovitch, E.M.B., Eillis, D.E., 2002. Fe2þ/Fe3þ substitution in hydroxyapatite: theory and experiment. Phys. Rev. B 66, 224107 with permission of The American Physical Society.

11.5.1 Magnetization As Fe2þ and Fe3þ are paramagnetic, the iron-substituted HA possesses paramagnetic properties. The magnetization of iron-substituted HA is measured by a vibrating magnetometer. The magnetization increases with the external magnetic field. The value also increases with the concentration of Fe (III) incorporated (Li et al., 2009). The unique paramagnetic properties of iron-substituted HA have many potential applications in biomedical fields.

11.5.2 Mechanical properties One of the disadvantages of HA is the brittleness and low mechanical strength. Thus, it is unsuitable for the load-bearing application. One approach to improve the mechanical strength is by doping with different cations. Iron-substituted HA can improve the mechanical properties of HA. The surface hardness and wear resistance of iron-substituted HA were improved. The surface hardness increased with an increase in the doped iron concentration from a range of 0.01, 0.05, and 0.1 M. The iron-substituted HA was prepared by the solegel method. The XRD results revealed pure iron-substituted HA phase with decreased peak intensities and increased peak broadening compared with pure HA. There were no secondary phases (Chandra et al., 2012). For another study, the iron-substituted HA was prepared by microwave method. The XRD patterns showed the pure iron-substituted HA phase. The wear resistance was improved with iron content (1.5 and 3 wt%). However, there was no improvement

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with iron content of 5 wt%. The average friction coefficient was decreased with the increase in the content of iron, i.e., 0.612 to 0.560 for pure HA to 1.5 wt% for ironsubstituted HA; however, for 3 wt% and 5 wt% iron-substituted HA, the coefficient was 0.587 and 0.607, respectively. It is indicated that an appropriate content of iron substitution could lead to decrease in the coefficient friction of the HA (Han et al., 2017).

11.6

Biological properties of iron-substituted hydroxyapatite

A material is bioactive when it triggers a specific biological reaction at the materiale tissue interface, by the mechanism of the formation of biochemical linkages between the tissue cells and the material. The biological performance, i.e., the interactions between the materials and the living tissues, of the metal ionesubstituted HA depends on the nature of the substituted ions and doped ion concentration. It can be assessed by in vitro bioactivity and cell culture tests. The results give a good prediction of how the material is in response with the tissue interaction.

11.6.1

In vitro bioactivity

A bonelike apatite, amorphous calcium phosphate or HA, is formed on the surface of the bioactive material when immersed in simulated body fluid (SBF). The formation mechanism of apatite on the surface is due to the ions released from the material surface and ion exchange with SBF. The apatite particles deposited on the surface stimulates signaling proteins and cells. This initiates a sequence of processes that led to bone formation. Therefore, the material can bond directly to host bone tissue. For the in vitro bioactivity of iron (III)-substituted HA, the apatite formation was more favorable than HA after soaking in SBF (Ereiba et al., 2013). Fig. 11.7 shows the FTIR spectra of the apatite formation after 24d soaking in SBF. The PO3 4 group peak intensities of apatite formed for iron-substituted HA is stronger than that of HA. The amount of apatite formed also increases with high amount of Fe3þ doped. A second phase, hematite (Fe2O3), is also detected from the XRD analysis (Fig. 11.8). This is due to the dissolution of iron-substituted HA after soaking in SBF.

11.6.2

Biocompatibility

11.6.2.1 Cell cytotoxicity Assessment of the cell toxicity is very important as this is the first step to determine the potential toxicity of the tested materials. They may release toxic substances or interact with intracellular metabolism (nanosize materials) to generate reactive oxygen species (ROS). ROS formation is one of the mechanisms for nanosize materials toxicity (Kumar et al., 2017).

Iron-substituted hydroxyapatite

(O-H) stretching

271 (CO3)–2

H2O stretching

HAp 3568

H2O bending 1631 1458- 1426

3431

Transmittance (%)

FeHAp1 3569

3574

3570

FeHAp2

3423

FeHAp3

3435

FeHAp4

4000

1454- 1423

1453- 1423 1032 962 631 1454- 1422

3428

3500

3000

2500

1034

876

1640

1634 3566

470 875 961 631 566 603 –3 (O-H) bending (PO4)

1036

875 469 962 631 604 567 875 468 1641 1453- 1424 1033 631 961 602 564 1634

3435

(PO4)–3

2000

1500

472

876 603 564 466 961 631 604 566 1035

500

1000

Wavenumber (cm–1)

Figure 11.7 Fourier-transform infrared spectra of the in vitro apatite formation on ironsubstituted HA after 24 d immersion in simulated body fluid. Reproduced from Ereiba, K.M.T., Mostafa, A.G., Gamal, G.A., Said, A.H., 2013. In vitro study of iron doped hydroxyapatite. J. Biophys. Chem. 4, 122e130 with permission of Scientific Research Publisher. Hematite HAp 211 300

Pure HAp

004

Intensity

002

FeHAp1

FeHAp2

FeHAp3

FeHAp4 10

20

30

40

50

60

70

80

2q

Figure 11.8 X-ray diffraction of iron-substituted hydroxyapatite (HA) after immersion in simulated body fluid (SBF) for 24 days. Formation of hematite is detected. Reproduced from Ereiba, K.M.T., Mostafa, A.G., Gamal, G.A., Said, A.H., 2013. In vitro study of iron doped hydroxyapatite. J. Biophys. Chem. 4, 122e130 with permission of Scientific Research Publisher.

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80000 Total number of cells

*** ***

60000

HA 2000 mg/ml HA 1000 mg/ml

n.s.

HA 500 mg/ml HA 200 mg/ml

40000

FeHA 2000 mg/ml FeHA 1000 mg/ml

20000

FeHA 500 mg/ml FeHA 200 mg/ml

0 d1

d3

d7

Figure 11.9 The cell proliferation of iron-substituted hydroxyapatite (HA) with different amounts of doped iron. Reproduced from Panseri, S., Cunha, C., D’Alessandro, T., Sandri, M., Giavaresi, G., Marcacci, M., Hung, C.T., Tampieri, A., 2012. Intrinsically superparamagnetic Fe-hydroxyapatite nanoparticles positively influence osteoblast-like cell behavior. J. Nanobiotechnol. 10, 32. https://doi: 10.1186/ 1477-3155-10-32 with permission of BioMed Central Ltd.

Iron-substituted HA possesses very low cell cytotoxicity toward human osteoblast cells compared with HA. Furthermore, there is no significant change in cell cytotoxicity when the amount of doped iron increased from 1 to 20 wt% (Li et al., 2009).

11.6.2.2 Cell proliferation Iron-substituted HA enhances cell proliferation toward human osteoblast cells. A significant increase in the cell proliferation at 3d and 7d for iron-substituted HA when compared with HA was observed (Fig. 11.9) (Li et al., 2009). When ironsubstituted HA nanoparticles were exposed to a static magnetic field, there was a significant increase in cell proliferation compared with the groups without any magnetic field application (Fig. 11.10) (Panseri et al., 2012). With the exposure of magnetic field, the lowest concentration of iron-substituted HA caused more cell proliferation than the highest concentration of iron-substituted HA. This suggests that the application of the magnetic field induces synergy to iron-substituted HA to enhance cell proliferation.

11.6.2.3 Cell differentiation and alkaline phosphatase activity When a static magnetic field was applied, the alkaline phosphatase activity of ironsubstituted HA nanoparticles was increased with days of culture (Fig. 11.11) (Panseri et al., 2012). This showed that the application of static magnetic field also promoted osteoblast activity. For the composite scaffold of poly(lactic-co-glycolic acid) and iron-substituted HA, alkaline phosphatase activity and osteoblastic differentiation process was enhanced in rat bone mesenchymal stem cells with the application of a static magnetic field. In vivo study showed that (poly(lactic-co-glycolic acid)/iron-substituted HA (PLGA/FeHA) scaffold enhanced the osteogenesis process (Yijun et al., 2017).

Iron-substituted hydroxyapatite

273

Total number of cells

80000 ***

**

**

60000

FeHA 2000 mg/ml ***

FeHA 1000 mg/ml FeHA 500 mg/ml FeHA 200 mg/ml

**

40000

FeHA 2000 mg/ml SMF FeHA 1000 mg/ml SMF

20000

FeHA 500 mg/ml SMF FeHA 200 mg/ml SMF

0 d1

d3

d7

Figure 11.10 Cell proliferation of human osteoblast cells for hydroxyapatite and ironsubstituted hydroxyapatite (HA) in the presence of a static magnetic field. Reproduced from Panseri, S., Cunha, C., D’Alessandro, T., Sandri, M., Giavaresi, G., Marcacci, M., Hung, C.T., Tampieri, A., 2012. Intrinsically superparamagnetic Fe-hydroxyapatite nanoparticles positively influence osteoblast-like cell behavior. J. Nanobiotechnol. 10, 32. https://doi: 10.1186/ 1477-3155-10-32 with permission of BioMed Central Ltd.

ALP activity (U/I/cell)

0.8 FeHA 2000 mg/ml FeHA 1000 mg/ml

0.6

FeHA 500 mg/ml FeHA 200 mg/ml

*

0.4

FeHA 2000 mg/ml SMF FeHA 1000 mg/ml SMF

0.2

FeHA 500 mg/ml SMF FeHA 200 mg/ml SMF

0.0 d1

d3

d7

Figure 11.11 Alkaline phosphatase (ALP) activity assay for iron-substituted hydroxyapatite (HA) in the presence of a static magnetic field. Reproduced from Panseri, S., Cunha, C., D’Alessandro, T., Sandri, M., Giavaresi, G., Marcacci, M., Hung, C.T., Tampieri, A., 2012. Intrinsically superparamagnetic Fe-hydroxyapatite nanoparticles positively influence osteoblast-like cell behavior. J. Nanobiotechnol. 10, 32. https://doi: 10.1186/ 1477-3155-10-32 with permission of BioMed Central Ltd.

11.7

Biomedical applications

Iron-substituted HA has vast applications in medical and biological fields because of its paramagnetic, nontoxic, and biocompatible properties. Applications in drug delivery, biosensor, protein and gene delivery, magnetic resonance imaging (MRI), scaffolds, and hyperthermia therapy will be discussed in details in the following sections.

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Handbook of Ionic Substituted Hydroxyapatites

Drug delivery

Drug delivery is the technique of transporting a drug compound to attain a therapeutic effect in humans and animals. The delivery is achieved by means of drug carriers. Carriers are molecules or complexes that transport the drug molecules to the target site of interest by the process of encapsulation. Delivering drug at controlled rate, slow delivery, and to the target sites are very important. The drug carrier should be stable, nontoxic, and biodegradable (Tiwari et al., 2012). One of the important applications of iron-substituted HA nanoparticles is drug delivery. The particles size, i.e., nanosize, should be small enough to be put into the blood circulation system to pass through the capillary systems. The drug release profile of an anticancer drug, methotrexate, loaded iron-substituted HA nanoparticles was controllable and sustainable. The iron-substituted HA was synthesized by solegel method with iron concentrations from 0.005 M, 0.01 M, 0.015 M, 0.02 M, to 0.025 M. The XRD patterns revealed the iron-substituted HA peaks accompanied by an increase in the peak broadening and peak shift with respect to increase in iron concentration. The release rate was decreased significantly with increase in the iron content. This meant a control drug release over a long period of time and minimizes the drug accumulation and systemic side effects. The substitution of iron in HA would enhance the interaction between iron-substituted HA nanoparticles and the drug molecules, i.e., the strong bonding between the iron and the drug molecules. It was postulated that Fe3þ ions form a covalent interaction with the drug molecules, thus delaying the drug release, i.e., hindering the free movement of drug molecules. As the Fe3þ concentration increased, more bonding between Fe3þ and drug molecules was formed. Thus, it further delayed the drug release. This was more effective in sustained and controlled drug release (Sheikh et al., 2018).

11.7.2

Biosensor

A biosensor is a device that can transform a biological response to an electrical signal. It is very specific and stable toward pH and temperature (Mehrotra, 2016). The mechanisms are based on (i) input signal reception by the cells owing to an external stimuli; (ii) transmission of data toward the brain for clarification in the form of the neurological impulses; and (iii) receptors respond to the stimulus by the operation center. There are two classes of sensors: physical sensors and chemical sensors. Physical sensor responds on the external physical stimuli such as acoustic waves and pressure and electromagnetic radiations. Therefore, a sensing device that provides response to the physical property of the medium is termed as physical sensor. For chemical sensors, it can transform physical and chemical properties of the system into analytic signals. The magnitude is, in general, proportional to the concentration of the analyte (Ali et al., 2017). One of the applications of iron-substituted HA as the biosensor is the measurement of amino acid such as L-tyrosine in food. L-tyrosine can act as a precursor for dopa, dopamine, thyroxin, noradrenalin, and adrenalin as the hormone or nontuberculosis mycobacterial in mammalian central nervous systems. Sometimes, it is added to food products as the amount is limited in foods.

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275

The biosensor was fabricated by mixing tyrosinase with iron-substituted HA nanoparticles suspension and was cast (by drop-casting method) onto the glassy carbon electrode. The electrode had a wide range of detection and lower detection limit. It was sensitive and selective toward the detection of L-tyrosine in the existence of excess of boric acid, oxalic acid, urea, galactose, and sucrose (Kanchana et al., 2014).

11.7.3 Protein and gene delivery The design and development of efficient carriers to deliver protein and genes such as DNA and RNA into the target site is important to treat or prevent a genetic disease. The new DNA usually contains a functional gene to remove the effects of a gene mutation. Different techniques such as electrostatic forces, cleavable linkers, and degradable property are used after specific cells are targeted. One of the approaches is the binding of the protein or genetic materials to the surface of iron-substituted HA nanoparticles through the electrostatic forces. The coating of the iron-substituted HA can be tuned to change the surface charges, i.e., positive, negative, and neutral. Furthermore, many target molecules have good affinity and interactions to the HA. Nowadays, there are different new technologies such as recombinant DNA technology to treat human diseases. The DNA loading and transfection efficiencies of ironsubstituted HA were found to be higher than HA (Xiong et al., 2015). It was found that iron-substituted HA (MHA) nanoparticles had a binding affinity to plasmid DNA (pDNA) to form a transfection complex (pDNA-MHA). The pDNA-MHA complex nanoparticles showed an increase in gene delivery efficacy across the cell membrane and demonstrated specific localization under the control of a magnetic field. Iron-substituted HA nanoparticles have good potential for gene delivery (Wu and Lin, 2010).

11.7.4 Magnetic resonance imaging MRI is an imaging technique from which a noninvasive tomographic visualization of soft tissue structures with high spatial resolution and high contrast is obtained. The development of MRI is one of the most powerful techniques in clinical diagnosis as it is also accompanied by the advancement of deign of contrast agents (CAs), which enhances the image quality. MR images’ contract depends on the differences in tissue relaxation times, both longitudinal (T1) and transverse (T2), to produce image contrast. When a radiofrequency (RF) pulse is applied perpendicular to the magnetic field, the protons are excited and the nuclear spins realign themselves with the magnetic field, which is called as relaxation. MRI signals are modulated by the relaxation rates of protons to return to an equilibrium state after an RF pulse. The difference in T1 and T2 relaxation times allow to distinguish between soft tissues, bone, air, and liquids in the body. Detection of disease with MRI is usually difficult because the areas of tissue disease have similar signal intensity with the surrounding normal tissue cells. Therefore, a signal enhancement area is required using CAs. Contact agents interact with water molecules such that proton T1 or T2 relaxation times are altered. Contrast in MRI is often defined by “T1” (spinespin) or “T2” (spinelattice)

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5 mm

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Figure 11.12 In vivo magnetic resonance imaging (MRI) of the mouse liver after injection of iron-substituted hydroxyapatite (HA) nanoparticles. Reproduced from Adamiano, A., Iafisco, M., Sandri, M., Basini, M., Arosio, P., Canu, T., Sitia, G., Esposito, A., Iannotti, V., Ausanio, G., Fragogeorgi, E., Rouchota, M., Loudos, G., Lascialfari, A., Tampieri, A., 2018. On the use of superparamagnetic hydroxyapatite nanoparticles as an agent for magnetic and nuclear in vivo imaging. Acta Biomater. 73, 458e469 with permission of Elsevier.

relaxation times, which depend on the exact pulse sequence used to excite and then measure the relaxing spins. An enhanced tissue contract is produced by CAs that can alter local T1 or T2 relaxation times as compared with the initial background T1 or T2 signals (Yadollahpour et al., 2017). CAs of MRI, which are paramagnetic, superparamagnetic, and ferromagnetic materials, can improve the image contrast between normal and diseased tissue. It enhances the tissue visibility. In Fig. 11.12, the injection of iron-substituted HA nanoparticles showed higher image contract in the liver at 10 min compared with a commercial MRI CA, Endorem. The enhancement was more pronounced at 24 h. The iron-substituted HA nanoparticles have a longer endurance in the liver with respect to Endorem. That might be related to the high transversal relaxation rate (T2) and to the higher amount of iron-substituted HA nanoparticles that were injected with respect to Endorem to attain the same iron dose, as the two agents consisted of different amounts of iron (9.7 wt% vs. 71.0 wt%) (Adamiano et al., 2018).

11.7.5

Scaffolds

In bone tissue engineering, scaffolds are used for the repairing and substituting damaged or lost bone tissue. Some criteria must be met when developing scaffolds, e.g., porosity that allows cells to migrate through the pores and grow into the pores to prevent loosening and movement of the scaffold, good drug delivery, tissue infiltration, and vascularization. External stimulations such as magnetic pulse, ultrasound, and electrical induction are important to enhance new bone cell regeneration. Iron-substituted HA scaffolds are effective to improve bone cell growth after implantation by applying an external magnetic field (Panseri et al., 2013).

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Scaffold fixation is very crucial for success in osteogenic corrective surgery. It is especially important for large defects (>2 cm2). The scaffold is held efficiently in a fixed position at the boneescaffold interface by four small permanent magnet pins. This provides a sufficient scaffold fixation for efficient bone tissue regeneration (Russo et al., 2012). Superparamagnetic particles can act as a single magnetic domain to enhance micromotion at the cellescaffold interface. This might trigger the ion channels on the cell membrane and activate the mechanotransduction pathway, thus leading to increase in cell growth, proliferation, and differentiation (Mondal et al., 2017). A controlled temperature elevation can maintain therapeutic temperature, without using any local temperature control system (Shaterabadi et al., 2018). The bone repairing effect of iron-doped HA was investigated in an animal study (Zhao et al., 2019). An iron-doped HA composite scaffold was prepared from irondoped HA (mixed Fe2þ and Fe3þ doped), collagen, and chitosan by in situ crystallization technology (Chen et al., 2015). The scaffolds were implanted into the calvarial defects of the Sprague Dawley rats. Two control groups, HA composite scaffold and without scaffolds (blank group), were prepared. From the micro-CT (Fig. 11.13(aef)) and histological analysis (Fig. 11.13(gel)) after 12 weeks of implantation, the rat calvarial defects for the iron-doped HA composite scaffold were obviously filled with new bone tissues (Fig. 11.13(eef, k,l)) while only partially filled with new bone tissues for HA composite scaffold (Fig. 11.13(c, d, i, j)). There was no observable new bone tissues formed in the blank group (Fig. 11.13(a, b, g, h)). The bone mineral density and bone volume density (BV/TV) of iron-doped HA composite scaffold were significantly higher than HA composite scaffold and the blank group (Fig. 11.13(men)). From the results, the improvement of the new bone regeneration might be due to the magnetocaloric effect of the iron-doped HA composite scaffold arose from the earth’s magnetic field.

11.7.6 Hyperthermia therapy Hyperthermia has drawn much attention for cancer treatment because of its advantages over chemotherapy and radiotherapy. The main advantage of the method is that cancer cells are killed directly in a short time, and normal cells are uninfluenced. Magnetic hyperthermia is based on applying an external alternating magnetic field. This in turn oscillates the magnetic moment of nanoparticles. As a result, magnetic energy is converted into heat energy. Superparamagnetic materials are excellent candidates for antitumor therapy because they can kill tumors deep inside the body and are controllable by an external magnetic field. There are several advantages: first, the particle size can be controlled from a few nanometers to tens of nanometers by different preparation methods; second, they can be controlled by an external magnetic field; and third, heat energy is generated, which can be used as hyperthermia agents. As a result, a large amount of thermal energy is transferred to tumor cells and destroys them (Sneha and Sundaram, 2015). The iron-substituted HA can be used as a heating mediator in hyperthermia therapy for treatment of tumor cells. In an in vivo study, the iron-substituted HA nanoparticles

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Figure 11.13 Micro-CT (aef) and hematoxylin and eosin (H&E)estained histological sections (gel) of rat calvarial defect implanted with the blank control, chitosan/collagen/HA (CS/Col/ HA) scaffold, and chitosan/collagen/iron/HA (CS/Col/Fe/HA) after 12 weeks. (m) Local bone mineral density analysis (m) and (n) morphometric analysis (BV/TV) of new bone formation for three groups in the defect site (*P < 0.05 and **P < 0.01). Reproduced from Zhao, Y., Fan, T., Chen, J., Su, J., Zhi, X., Pan, P., Zou, L., Zhang, Q., 2019. Magnetic bioinspired micro/nanostructured composite scaffold for bone regeneration. Colloids and Surfaces B: Biointerfaces 174, 70e79 with permission of Elsevier.

suspended in phosphate buffer solution (PBS) were injected subcutaneously around the tumor of mice. It was treated inside an applied alternating magnetic field to achieve hyperthermia. A thermal energy was generated and transferred to the target site, and the cells were destroyed. The result showed a dramatic reduction in tumor volume after 15 days of irradiation in a mouse model. There was no reduction in tumor volume without the magnetic field application (Hou et al., 2009).

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279

Conclusions

HA is an important biomaterial in tissue engineering. Despite the success of HA for a broad range of biomedical applications, there is still great potential to improve it. Substitution by other ions can change the physical, chemical, and biological properties. Iron-substituted HA is used in drug delivery, biosensor, protein and gene delivery, MRI, scaffolds, and hyperthermia therapy. The bioactivity can be further improved by the surface modification of ironsubstituted HA. This could enhance the interaction with DNA, drug, and other tissue cells. However, this in turn affects the saturation magnetization. The future challenge will be the surface modification of iron-substituted HA that will enhance the bioactivity but not deteriorate the saturation magnetization.

References Adamiano, A., Iafisco, M., Sandri, M., Basini, M., Arosio, P., Canu, T., Sitia, G., Esposito, A., Iannotti, V., Ausanio, G., Fragogeorgi, E., Rouchota, M., Loudos, G., Lascialfari, A., Tampieri, A., 2018. On the use of superparamagnetic hydroxyapatite nanoparticles as an agent for magnetic and nuclear in vivo imaging. Acta Biomater. 73, 458e469. Ali, J., Najeeb, J., Asim Ali, M., Aslam, M.F., Raza, A., 2017. Biosensors: their fundamentals, designs, types and most recent impactful applications: a review. J. Biosens. Bioelectron. 8, 235. https://doi.org/10.4172/2155-6210.1000235. Attia, S.M., Wang, J., Wu, G., Shen, J., Ma, J., 2002. Review on sol-gel derived coatings: process, techniques and optical applications. J. Mater. Sci. Technol. 18, 211e218. Belwal, S., 2013. Green revolution in chemistry by microwave assisted synthesis: a review. Mod. Chem. 1, 22e25. Chandra, V.S., Baskar, G., Suganthi, R.V., Elayaraja, K., Ahymah Joshy, M.I., Beaula, W.S., Mythili, R., Venkatraman, G., Kalkura, S.N., 2012. Blood compatibility of iron-doped nanosize hydroxyapatite and its drug release. ACS Appl. Mater. Interfaces 4, 1200e1210. Chen, J., Pan, P., Zhang, Y., Zhong, S., Zhang, Q., 2015. Preparation of chitosan/nano hydroxyapatite organiceinorganic hybrid microspheres for bone repair. Colloids Surfaces B Biointerfaces 134, 401e407. Elliott, J.C., Mackie, P.E., Young, R.A., 1973. Monoclinic hydroxyapatite. Science 180, 1055e1057. Ereiba, K.M.T., Mostafa, A.G., Gamal, G.A., Said, A.H., 2013. In vitro study of iron doped hydroxyapatite. J. Biophys. Chem. 4, 122e130. Han, S.X., Ning, Z.W., Chen, K., Zheng, J., 2017. Preparation and tribological properties of Fe-hydroxyapatite bioceramics. Biosurface Biotribol. 3, 75e81. Hou, C.H., Hou, S.M., Hsueh, Y.S., Lin, J., Wu, H.C., Lin, F.H., 2009. The in vivo performance of biomagnetic hydroxyapatite nanoparticles in cancer hyperthermia therapy. Biomaterials 30, 3956e3960. Jiang, M., Terra, J., Rossi, A.M., Morales, M.A., Saitovitch, E.M.B., Eillis, D.E., 2002. Fe2þ/ Fe3þ substitution in hydroxyapatite: theory and experiment. Phys. Rev. B 66, 224107. Jones, G., Prosser, D.E., Kaufmann, M., 2014. Cytochrome P450-mediated metabolism of vitamin D. JLR (J. Lipid Res.) 55, 13e31.

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Kanchana, P., Lavanya, N., Sekar, C., 2014. Development of amperometric L-tyrosine sensor based on Fe-doped hydroxyapatite nanoparticles. Mater. Sci. Eng. C 35, 85e91. Kaygili, O., Dorozhkin, S.V., Ates, T., Al-Ghamdi, A.A., Yakuphanoglu, F., 2014. Dielectric properties of Fe doped hydroxyapatite prepared by solegel method. Ceram. Int. 40, 9395e9402. Kramer, E.R., Morey, A.M., Staruch, M., Suib, S.L., Jain, M., Budnick, J.I., Wei, M., 2013. Synthesis and characterization of iron-substituted hydroxyapatite via a simple ion-exchange procedure. J. Mater. Sci. 48, 665e673. Kumar, V., Sharma, N., Maitra, S.S., 2017. In vitro and in vivo toxicity assessment of nanoparticles. Int. Nano Lett. 7, 243e256. Laura Toxqui, L., Vaquero, M.P., 2015. Chronic iron deficiency as an emerging risk factor for osteoporosis: a hypothesis. Nutrients 7, 2324e2344. Li, Y., Nam, C.T., Ooi, C.P., 2009. Iron (III) and manganese (II) substituted hydroxyapatite nanoparticles: characterization and cytotoxicity analysis. J. Phys. Conf. Ser. 187, 012024. https://doi.org/10.1088/1742-6596/187/1/012024. Lieu, P.T., Heiskala, M., Peterson, P.A., Yang, Y., 2001. The roles of iron in health and disease. Mol. Asp. Med. 22, 1e87. Lin, K., Chang, J., 2015. Hydroxyapatite (HAp) for biomedical applications. In: Mucalo, M. (Ed.), Structure and Properties of Hydroxyapatite for Biomedical Applications. Woodhead Publishing, Waltham, USA, pp. 5e10. Marinea, J., Myers, C.P., Picquet, G.A., Zaidel, L.A., Wu, D., Uhrich, K.E., 2018. Reduction of bacterial attachment on hydroxyapatite surfaces: using hydrophobicity and chemical functionality to enhance surface retention and prevent attachment. Colloids Surfaces B Biointerfaces 167, 531e537. Mehrotra, P., 2016. Biosensors and their applications e a review. J. Oral Biol. Craniofac. Res. 6, 153e159. Mondal, S., Manivasagan, P., Bharathiraja, S., Santha, M., Kim, H.H., Seo, H., Lee, K.D., Oh, J., 2017. Magnetic hydroxyapatite: a promising multifunctional platform for nanomedicine application. Int. J. Nanomed. 12, 8389e8410. Okada, M., Matsumoto, T., 2015. Synthesis and modification of apatite nanoparticles for use in dental and medical applications. Jpn. Dental Sci. Rev. 51, 85e95. Panseri, S., Cunha, C., D’Alessandro, T., Sandri, M., Giavaresi, G., Marcacci, M., Hung, C.T., Tampieri, A., 2012. Intrinsically superparamagnetic Fe-hydroxyapatite nanoparticles positively influence osteoblast-like cell behavior. J. Nanobiotechnol. 10, 32. https://doi.org/ 10.1186/1477-3155-10-32. Panseri, S., Russo, A., Sartori, M., et al., 2013. Modifying bone scaffold architecture in vivo with permanent magnets to facilitate fixation of magnetic scaffolds. Bone 56, 432e439. Posner, A.S., Perloff, A., Diorio, A.F., 1958. Refinement of the hydroxyapatite structure. Acta Crystallogr. 11, 308e309. Ramya, J.R., Arul, K.T., Elayaraja, K., Kalkura, S.N., 2014. Physicochemical and biological properties of iron and zinc ions co-doped nanocrystalline hydroxyapatite, synthesized by ultrasonication. Ceram. Int. 40, 16707e16717. Rao, B.G., Mukherjee, D., Reddy, B.M., 2017. Nanostructures for Novel Therapy Synthesis, Characterization and Applications, vols. 10e11. Elsevier Inc, p. 13. Chapter 1 Novel approaches for preparation of nanoparticles. Rau, J.V., Cacciotti, I., De Bonis, A., Fosca, M., Komlev, V.S., Latini, A., Santagata, A., Teghil, R., 2014. Fe-doped hydroxyapatite coatings for orthopedic and dental implant applications. Appl. Surf. Sci. 307, 301e305.

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Robles-Aguila, M.J., Reyes-Avenda~no, J.A., Mendoza, M.E., 2017. Structural analysis of metaldoped (Mn, Fe, Co, Ni, Cu, Zn) calcium hydroxyapatite synthetized by a sol-gel microwave-assisted method. Ceram. Int. 43, 12706e12709. Russo, A., Shelyakova, T., Casino, D., Lopomo, N., Strazzari, A., Ortolani, A., Visani, A., Dediu, V., Marcacci, M., 2012. A new approach to scaffold fixation by magnetic forces: application to large osteochondral defects. Med. Eng. Phys. 34, 1287e1293. Sato, M., Nakahira, A., 2013. Influence of Fe addition to hydroxyapatite by hydrothermal process. J. Ceram. Soc. Jpn. 121, 559e562. Sato, M., Nakahira, A., 2014. Preparation of iron doped hydroxyapatite microsphere by mist process. Mater. Trans. 55, 1536e1539. Shaterabadi, Z., Nabiyouni, G., Soleymani, M., 2018. Physics responsible for heating efficiency and self-controlled temperature rise of magnetic nanoparticles in magnetic hyperthermia therapy. Prog. Biophys. Mol. Biol. 133, 9e19. Sheikh, L., Sinha, S., Singhababu, Y.N., Verma, V., Tripathy, S., Nayar, S., 2018. Traversing the profile of biomimetically nanoengineered iron substituted hydroxyapatite: synthesis, characterization, property evaluation, and drug release modeling. RSC Adv. 8, 19389e19401. Shi, C., Gao, J., Wang, M., Fu, J., Wang, D., Zhu, Y., 2015. Ultra-trace silver-doped hydroxyapatite with non-cytotoxicity and effective antibacterial activity. Mater. Sci. Eng. C 55, 497e505. Sneha, M., Sundaram, N.M., 2015. Preparation and characterization of an iron oxidehydroxyapatite nanocomposite for potential bone cancer therapy. Int. J. Nanomed. 10, 99e106.  Supov
a, M., 2015. Substituted hydroxyapatites for biomedical applications: a review. Ceram. Int. 41, 9203e9231. Szcze
s, A., Hołysz, L., Chibowski, E., 2017. Synthesis of hydroxyapatite for biomedical applications. Adv. Colloid Interface Sci. 249, 321e330. Tiwari, G., Tiwari, R., Sriwastawa, B., Bhati, L., Pandey, S., Pandey, P., Bannerjee, S.K., 2012. Drug delivery systems: an updated review. Int. J. Pharm. Investig. 2, 2e11. Trinkunaite-Felsen, J., Prichodko, A., Semasko, M., Skaudzius, R., Beganskiene, A., Kareiva, A., 2015. Synthesis and characterization of iron-doped/substituted calcium hydroxyapatite from seashells Macoma balthica (L.). Adv. Powder Technol. 26, 1287e1293. Tsay, J., Yang, Z., Ross, F.P., Cunningham-Rundles, S., Lin, H., Coleman, R., MayerKuckuk, P., Doty, S.B., Grady, R.W., Giardina, P.J., Boskey, A.L., Vogiatzi, M.G., 2010. Bone loss caused by iron overload in a murine model: importance of oxidative stress. Blood 116, 2582e2589. Wang, D., Bierwagen, G.P., 2009. Solegel coatings on metals for corrosion protection. Prog. Org. Coat. 64, 327e338. Wu, H.C., Lin, F.H., 2010. Evaluation of magnetic-hydroxyapatite nanoparticles for gene delivery carrier. Biomed. Eng. Appl. Basis Commun. 22, 33e39. Wu, T.Y., Guo, N., Teh, C.Y., Hay, J.X.W., 2012. Advances in Ultrasound Technology for Environmental Remediation. Spring Publishing Co., pp. 5e12. Chapter 2 Theory and Fundamentals of Ultrasound. Xiong, G., Wan, Y., Zuo, G., Ren, K., Luo, H., 2015. Self-assembled magnetic lamellar hydroxyapatite as an efficient nanovector for gene delivery. Curr. Appl. Phys. 15, 811e818. Yadav, T.P., Yadav, R.M., Singh, D.P., 2012. Mechanical milling: a top down approach for the synthesis of nanomaterials and nanocomposites. Nanosci. Nanotechnol. 2, 22e48.

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Yadollahpour, A., Asl, H.M., Rashidi, S., 2017. Applications of nanoparticles in magnetic resonance imaging: a comprehensive review. Asian J. Pharm. 11, S7eS13. Yijun, Y., Shuangshuang, R., Yingfang, Y., He, Z., Chao, L., Jie, Y., Weidong, Y., Leiying, M., 2017. Electrospun fibrous scaffolds with iron-doped hydroxyapatite exhibit osteogenic potential with static magnetic field exposure. J. Biomed. Nanotechnol. 13, 835e847. Zhao, G.Y., Zhao, L.P., He, Y.F., Li, G.F., Gao, C., Li, K., Xu, Y.J., 2012. A comparison of the biological activities of human osteoblast hFOB1.19 between iron excess and iron deficiency. Biol. Trace Elem. Res. 150, 487e495. Zhao, Y., Fan, T., Chen, J., Su, J., Zhi, X., Pan, P., Zou, L., Zhang, Q., 2019. Magnetic bioinspired micro/nanostructured composite scaffold for bone regeneration. Colloids Surfaces B Biointerfaces 174, 70e79.

Silicon-substituted hydroxyapatite

12

Aysha Arshad, Ather Farooq Khan Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

12.1

Introduction

Hydroxyapatite (HA; Ca10(PO4)6(OH)2) is a type of calcium phosphate (main inorganic mineral) found in hard bone tissues of human and animals (Pietak et al., 2007; Wang et al., 2012). HA has been tremendously used to address various orthopedics-related problems as skeletal substitute material (Pietak et al., 2007) due to its chemical similarity to the mineral component of hard tissue (Cacciotti, 2014), e.g., biocompatible coatings for metal implants, bone-filling materials, and organic/ inorganic composite scaffolds for tissue engineering applications (El Yacoubi et al., 2014). However, HA lacks the ability to stimulate the formation process of the new bone tissues because of its poor reactivity with the existing bone. It can stay long enough there on injury site as fixture because of having distinctively poor degradability property, which in turn makes it more liable to collapse (Pietak et al., 2007) without completing the bone repairing process. Therefore, lots of efforts have been made to improve the biological reactivity of HA, including the biomimetic approach, intended to make resemblance with the natural bone structure and chemical composition of mineral phase. Natural bone mineral HA crystals are certainly calcium-deficient nonstoichiometric apatite and possess varied amounts of either cationic substitutions (replacing calcium site), i.e., magnesium (Mg2þ), sodium (Naþ), Zinc (Zn2þ), strontium (Sr2þ), aluminum (Al3þ), potassium (Kþ), and silver (Agþ), or anionic substitu2 4  tions (replacing OH or PO3 4 ), i.e., carbonate (CO3 ), silicate (SiO4 ), fluoride (F ),  and chloride (Cl ) (Shepherd et al., 2012; Camaioni et al., 2015). These substitutional ions impart significantly improved biological characteristics to synthetic HA implant structures by modulating the surface charge, crystallinity, thermal stability, surface chemistry, solubility, and crystal morphology, all of which resultantly influence the biological activity and response by increasing the osteoblast cells proliferation and as a result enhancing osteointegration (Pietak et al., 2007; Shepherd et al., 2012; Camaioni et al., 2015). Therefore, the present trend is to mimic it to develop substituted HA by these ionic species such as magnesium-substituted HA, Si-HA, carbonated HA, etc. (Pietak et al., 2007). In fact, the ability to exchange ions in apatite structure allows to design, develop, and characterize new and better calcium phosphates for certain specific applications (Cacciotti, 2014). Si, in particular, has been broadly used as substitutional element in variable biomedical materials as it attributes

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00012-4 Copyright © 2020 Elsevier Ltd. All rights reserved.

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enhanced bioactivity to physiological processes involved in skeletal system (Pietak et al., 2007). Thus, Si-HA is an indispensable player of the promising research field of bone repair and regeneration (Khan et al., 2014).

12.2

Biological importance of silicon

Next to oxygen, the most prevalent element in the earth’s crust is silicon (Si) (Pietak et al., 2007). In human body, Si is present at the similar concentrations as of the other physiologically significant elements such as iron (Fe), copper (Cu), and zinc (Zn) and is excreted in urine in similar magnitude as of calcium (one of the main bone mineral and most vital cell signaling molecules), thus strengthening the notion that Si may be a significant if not essential biological element (Jugdaohsingh, 2007). Si is the third major ubiquitous trace element (El Yacoubi et al., 2014; Farooq and Dietz, 2015; Jugdaohsingh, 2007) and is considered to have an essential role in various biological systems, i.e., skeletal repair and development (El Yacoubi et al., 2014; Ruys, 1993; Zhang et al., 2011; Zou et al., 2009). In humans, a positive correlation exists between bone mineral density and Si intake and is also associated with a decreased risk of osteoporosis as well (Macdonald et al., 2012; Costa-Rodrigues et al., 2016). Si is also considered to have a role in human hormonal control besides that of metabolic and physiological control (Boguszewska-Czubara and Pasternak, 2011; Pontigo et al., 2015). It minimizes the alopecia risk and Alzheimer’s disease and is also recommended as a guardian against cardiac diseases (Boguszewska-Czubara and Pasternak, 2011, Pontigo et al., 2015). Si also appears to inhibit the activity of macrophage and osteoclastic cells (Casarrubios et al., 2016). Silicic acid (Si(OH4)) is the prime water dissolvable physiological form of Si in human beings and has crucial health beneficial structural and functional role in the skin, kidney, blood vessels, liver, bones, trachea, and tendons (Reffitt et al., 2003; Nielsen, 2014; Farooq and Dietz, 2015), where it has crucial controlling influence on local metabolism (Costa-Rodrigues et al., 2016). Recent research on orthosilicic acid reported that it is involved in increased synthesis of collagen (type I); cellular differentiation and amplified mRNA expression of these proteins could be the indicator of Si association in gene regulation (Jugdaohsingh, 2007; Khan et al., 2014). In mammalians, the incidence of Si is relatively inconsistent: about 1 ppm is present in serum, 200e600 ppm Si present in the liver, kidney, lung, and muscle, 2e10 ppm in the bone and ligaments, and 100 ppm is found in the other connective tissues (cartilage and the umbilical cord) (Pietak et al., 2007). From structural aspect, Si is present in connective tissues bound to ECM (extracellular matrix) components, i.e., mucopolysaccharides, polysaccharides, glycosaminoglycans (Jugdaohsingh, 2007), chondroitin sulfate, hyaluronic acid, and dermatan sulfate when examined by chemical techniques (Szurkowska and Kolmas, 2017; Pietak et al., 2007). Along with metabolic role, silicon’s high abundance in ECM compounds suggests that it also plays an important architectural role as a biological cross-linker and stabilizer of the glycosaminoglycans fibers and collagen network that contributes to the structure and resilience of

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connective tissue (Costa-Rodrigues et al., 2016; Pietak et al., 2007). At the cellular level, Si is considered as an affirmative regulator of osteoblast cells as well as an activity suppressor of osteoclast cells (Beck Jr et al., 2012, Wiens et al., 2010; M€uller et al., 2014; Schr€ oder et al., 2012; Friederichs et al., 2015; Lehmann et al., 2012; Costa-Rodrigues et al., 2016). Bioactive silica also involved in upregulated expression of vascular endothelial growth factor, which plays a role in blood vessel and bone formation (Casarrubios et al., 2016). Carlisle and Schwarz were the first scientists to rule out the role of Si at active calcification sites in embryonic chondrocytes/skull bones of chicks/rats and mice, respectively, during early bone biomineralization events using electron microprobe analysis (Pietak et al., 2007; Han et al., 2013; Jugdaohsingh, 2007; Wiens et al., 2010; Rodrigues et al., 2017; Coathup et al., 2011). They found that Si dependently amplifies prolyl hydroxylase (enzyme involved in collagen synthesis) expression (Jugdaohsingh, 2007). Their study led to the conclusion that scarcity of Si can be a causative aberration factor regarding formation, biomineralization, and development of connective tissues/articular cartilage in bone architecture (Costa-Rodrigues et al., 2016; Zhang et al., 2011). Hench et al. reported that wear and tear in the proliferation and the operation of bone forming cells due to osteoporosis and osteopenia are associated with the loss of biological availability of Si (Khan et al., 2014; Hench, 1998). Soluble Si was also found to enhance the osteocalcin and alkaline phosphatase (ALP) expression and to accelerate the proliferation of osteoblasts, which are osteogenic differentiation markers (Rodrigues et al., 2017). Thus, the Si seems to be essential for the bone formation and is linked with the inorganic as well as the organic osteogenic phase of bone development in human biology (El-Ghannam, 2005).

12.3

Silicon-substituted hydroxyapatite synthesis methods

Different synthesis methods of Si-HA can be largely divided into five classes. These methods are summarized in Table 12.1.

12.3.1 Precipitation method This is the most commonly used synthesis method for the pure HA and Si-HA. In this method, silicon tetra-acetate (Si(CH3CO2)4) or tetraethyl orthosilicate (Si(OC2H5)4, TEOS) is used as a main source of SiO4 4 ions, calcium nitrate (Ca(NO3)2.4H2O) supplies Ca2þ ions, and orthophosphoric acid or diammonium hydrogen phosphate ((NH4)2HPO4) provides phosphorus ions. The combination of aqueous solutions containing Si, Ca, and P ions results in the formation of a precipitate. Reaction is normally adjusted to basic pH (9e11) by adding ammonia (NH3). The basic steps involved in this method are precipitate aging following by the filtration step as well as washing step with distilled water to remove impurities. The obtained precipitate is dried and then sintered at >600 C or treated with microwave heating. To avoid the inclusion

Table 12.1 Silicon-substituted hydroxyapatite (Si-HA) Synthesis methods. Substituted Si wt%

References

Solid-state reaction Mechanochemical methods

Ca2P2O7, CaCO3, and SiO2

0.7 and 1.4

(Hahn et al., 2010; Leshkivich and Monroe, 1993)

Ca(OH)2, H3PO4, and Si(CH3COO)4

0.8

(Hayakawa et al., 2013)

Ca(OH)2, (NH4)2HPO4, and Si(OCH2CH3)4 (TEOS)

0.5, 1.0, 1.5, and 2.0

(Tian et al., 2008)

Ca(H2PO4)2$H2O, CaO, ans SiO2$0.7H2O

0.6, 0.8, 1.0, and 2.0

(Chaikina et al., 2014)

Ca(NO3)2.4H2O, (NH4)2HPO4, Si(CH3CH2O)4

0.8

(Casarrubios et al., 2016)

Ca(NO3)2.4H2O, (NH4)2HPO4, Si(OCOCH3)4

0.2, 0.4, 0.6, 0.8, 1.0, 1.5, 2.0, and 4.0

(Palard et al., 2008; Magnaudeix et al., 2016)

H3PO4, Ca(OH)2, Si(OCOCH3)4

1.6

(Bang et al., 2011)

i). Ca(NO3)2.4H2O, (NH4)2HPO4, and Si(CH3CO2)4 ii). Ca(OH)2, H3PO4, and Si(OC4H9)4 (TEOS)

1.40

(Bianco et al., 2009)

Ca(OH)2, H3PO4, and Si(OCH2CH3)4 (TEOS)

1.4, 1.6, and 1.8

(Palard et al., 2008; El Yacoubi et al., 2014)

(NH4)2HPO4, Ca(NO3)2.4H2O, and alkaline silicate solution (if applicable)

0.25, 0.50, 0.75, 1.00, and 1.25

(Marchat et al., 2013)

Ca(OH)2, Si(OCH2CH3)4 (TEOS), and H3PO4

0.8 and 1.5

(Jarcho et al., 1976; Friederichs et al., 2015)

Ca(NO3)2.4H2O, (NH4)2HPO4, Si(OCH2CH3)4 (TEOS), and NaHCO3

1.2

(Ibrahim et al., 2011)

CaCO3, Ca(OH)2, NaH2PO4, and Na4SiO

0.08, 0.75, 2.23, 5.15, and 8.1

(Mostafa et al., 2011)

Ca(NO3)2$H2O, (NH4)2HPO4, and Si(OCH2CH3)4 (TEOS)

0.8 and 1.6

(Yu et al., 2017)

Ca(OH)2, H3PO4, Si(OCH2CH3)4 (TEOS), and Si(OC4H9)4

1.4

(Lehmann et al., 2012)

Si(CH3COO)4, Ca(OH)2, and H3PO4

0.4, 1.4, 2.4, and 4.6

(Gasqueres et al., 2008)

Ca(NO3)2.4H2O, (NH4)2HPO4, 25% NH4OH, and Na2SiO3

1.5

(Bogya et al., 2015)

Ca(NO3)2.4H2O, (NH4)2HPO4, and Si(CH3CH2O)4

e

(Manzano et al., 2011)

Precipitation method

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Raw materials

286

Synthesis method

Solegel method

Ca(NO3)2.4H2O, (NH4)3PO4 or (NH4)2HPO4), and Si(OCH2CH3)4 (TEOS)

0.8 and 1.5

(Aminian et al., 2011)

Ca(NO3)2.4H2O, (NH4)3PO4, and Si(OCH2CH3)4 (TEOS)

0.8, 1.5, and 4.0

(Tang et al., 2005)

(Ca(NO3)2.4H2O, (NH4)2HPO4, CH3(CH2)15 N(CH3)3 Br (CTAB), NH4OH (25%), and Si(OCH2CH3)4 (TEOS)

0.2, 0.4, 0.6, 0.8, and 1.0

(Belmamouni et al., 2015)

P(OCH2CH3)3 (TEP), Si(OC2H5)4 (TEOS), and Ca(NO3)2$4H2O

1, 3, and 5 mol %

(Balamurugan et al., 2008)

CaCl2d2H2O, Na2HPO4d2H2O, CH3(CH2)15N(CH3)3Br (CTAB), 25% NH4OH, and Si(OC2H5)4 (TEOS)

2.4, 12, and 60

(Andersson et al., 2005)

P2O5, Ca(NO3)2.4H2O, C2H5OH, Si(OC2H5)4 (TEOS), C8H20O4Si), and C7H18O3Si (MTES)

10, 20, 30, and 40

(Latifi et al., 2011)

Silicon-substituted hydroxyapatite

Hydrothermal method

287

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of CO2 from the atmosphere, some scientists have used argon atmosphere for their experiment. This synthesis method can produce a finer single crystalline phase SiHA normally with a nominal composition of Ca:(PþSi) ratio of 1.67 (Palard et al., 2008; Magnaudeix et al., 2016; Casarrubios et al., 2016).

12.3.2

Solegel method

The solegel method is being used by various researchers for Si-HA powder preparation with 1, 3, and 5 wt.% Si content, using triethyl phosphate (P(OCH2CH3)3, TEP) (provides P), TEOS (supplies Si), and calcium nitrate (Ca(NO3)2, source for Ca). This method involves a transition from a colloidal solution (sol) to a gel form. The main steps of this method involve the hydrolysis of TEP followed by subsequent addition of TEOS and Ca(NO3)2 under stirring conditions. The produced sol is then processed at room temperature for 24 h, followed by aging and drying steps to remove gaseous byproducts. This processed gel can then further be subjected to heat treatment at approximately 800 C in air (Balamurugan et al., 2008; Camaioni et al., 2015; Szurkowska et al., 2017).

12.3.3

Hydrothermal method

Hydrothermal reactions are carried out in aqueous media at high pressure and high temperature, frequently in autoclave. This method typically produces HA with good dispersion and high degrees of crystallinity. These crystals can be modified into various morphologies by modulating the reaction conditions (Camaioni et al., 2015; Szurkowska et al., 2017). 1.8 wt% Si-HA was produced by Aminian et al. using hydrothermal method carried out at 200 C for 8 h. The precursors were Si(OCH2CH3)4 (TEOS), Ca(NO3)2, and (NH4)3PO4/(NH4)2HPO4, and polyethylene glycol was also used for better dispersive properties of material (Aminian et al., 2011). Kim et al. carried out a reaction combining both solvothermal and hydrothermal treatments (the former in silicon acetateesaturated acetone solution) to convert natural coral (provide Ca) into HA with up to 0.19 wt% Si (Kim et al., 2005). Si-HA with Si content of 0.8, 1.5, and 4.0 wt.% can be synthesized by following this method (Tang et al., 2005; Aminian et al., 2011; Szurkowska et al., 2017).

12.3.4

Solid-state reaction

No extensive literature is present on solid-state synthesis of solely Si-HA. The synthesis of apatites substituted with silicates, sulfates, lanthanum, and fluorine has been reported by Boyer et al. and Leshkivich et al. All reagents, i.e., Ca2P2O7, CaCO3, and SiO2, were ground in a ball mill for 30 h and were mixed and dried in rotary evaporator, compressed, and then followed by the sintering step for extending period of time at high temperature (1100 C for 20 h) under air atmospheric environment (Hahn et al., 2010; Leshkivich and Monroe, 1993). This novel method is the subject of a patent application.

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12.3.4.1 Mechanochemical method Mechanochemical method is also a solid-state method and has been conducted for producing Si-HA powder, and it is a very economical and simpler method, which can be carried out in aqueous phase or in dry state as well at ambient temperature. A detailed Si-HA procedure by mechanochemical method has been reported by Chaikina et al., where they implied amorphous silica, Ca(H2PO4)2$H2O or CaHPO4, and CaO. In only half an hour reaction in conventional ball mill, pure Si-HA was produced (Chaikina et al., 2014). Tian et al. also reported the mechanochemical synthesis method and obtained Si-HA with 0.5, 1.0, 1.5, and 2.0 wt.% Si, using Ca(OH)2, (NH4)2HPO4, and TEOS (Hayakawa et al., 2013; Chaikina et al., 2014).

12.4

Characterization of silicon-substituted hydroxyapatite

Synthetic HA has been tremendously used to address various orthopedics-related injuries as a skeletal substitute material. However, synthetic HA lacks the ability to stimulate the formation process of the new bone tissues (Pietak et al., 2007). It has been demonstrated in literature that HA bioactivity can be prudently enhanced by the inclusion of selected ions within the apatite lattice (Macdonald et al., 2012). Since 1990s, ion substitution in HA, particularly substitution of PO3 4 with SiO4 , has been extensively applied in biomedical material engineering, mainly 4 because of the simplicity of this process (Camaioni et al., 2015; Szurkowska and Kolmas, 2017; Shepherd et al., 2012). Synthetic stoichiometry HA with substituted small amounts of Si has shown superior biological functions in terms of improved bone apposition, bone ingrowth, and cell-mediated degradation (Bianco et al., 2009). Literature including both in vitro and in vivo studies has demonstrated that the substitution by SiO4 in HA results in increased solubility, topographical 4 changes, surface charge modification, grain size reduction, thermal stability, and 2þ ionic release of SiO4 4 and Ca , which modulates osteoclastic and osteoblastic cells response. Ionic substitution also efficiently increases the surface area, which in turn enhances the bioactivity of the material in vivo (Bang et al., 2015). The silanol groups in ceramics material have been considered as catalysts of the apatite (Ap) phase nucleation to form surface apatite layers. Additionally, it is important that substituted silicon does not result in thermal instability of the Si-HA as this effect on sintering would result in the decomposition of the Si-HA to undesirable second phases (Tian et al., 2008). Therefore, Si-HA shows comparatively better biological activity in vivo than HA, explaining its beneficial effects on the early bone developmental process. Therefore, these facts cumulatively make Si-HA an ideal candidate as bone substitute material usage, and Si-HA has presently been included to the biomaterials market as Actifuse ABXTM (Apatech Ltd, UK) for orthopedic, oral, spinal, periodontal, and craniomaxillofacial applications (Casarrubios et al., 2016; de la Concepci on Matesanz et al., 2012).

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Figure 12.1 Scheme of silicate substitution into the hydroxyapatite (HA) crystal (Szurkowska and Kolmas, 2017).

OH

Ca PO4

SiO4

Till date, the incorporated Si quantity in the HA structure has not been precisely determined. Studies have revealed that it may be dependent on the used synthesis method and its relative conditions. Ap modified with silicate ions can be obtained by several methods, i.e., the solegel method, hydrothermal synthesis, solid-state ceramic synthesis, and the wet precipitation method (Chaikina et al., 2014). Si-HA synthesis methods with their starting raw materials are enlisted in Table 12.1. There are numerous studies that explain the mechanism of Si-HA formation and the influence of Si incorporation on structural and biological characteristics of Si-HA (Bang et al., 2011). In literature, the most important characterization techniques, i.e., Fouriertransform infrared (FTIR) spectroscopy, X-ray diffraction (XRD), nuclear magnetic resonance (NMR) spectroscopy, etc., have been reported to study the structural and compositional changes after Si substitution in HA. Gibson et al. (1999a) reported the most trusted mechanism of silicate (SiO4 4 ) substitution into HA lattice structure with the formation of anionic vacancies at hydroxyl (OH) sites to fabricate singlephase bioceramics and proposed the generic chemical formula “Ca10(PO4)6x(SiO4)x(OH)2-x(VOH)x,” of the Si-HA, where x represents the molar number of SiO4 4 groups incorporated into the apatite lattice structure (0  x  2) and VOH depicts the vacancies maintaining the charge balance (Fig. 12.1) (El Yacoubi et al., 2014; Szurkowska and Kolmas, 2017). The Si amount that can be substituted into HA is considered to be limited up to  5 wt% (Gasqueres et al., 2008; Vallet-Regí and Arcos, 2005; Supov a, 2015), and almost 1 wt% Si amount has been recommended optimum to enhance main biological  activities (Supov a, 2015). The suggested equation in result of this substitution mechanism is as follows: 4  10Ca2 þ þ (6-x)PO3 4 þ xSiO4 þ (2-x)OH / Ca10(PO4)6-x (SiO4)x (OH)2-x

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291

Intensity (a.u)

Unmark peaks: HA and Si-HA • α- TCP

(d) (c) (b) (a)

20

25

30

35 40 2θ (degree)

45

50

55

Figure 12.2 X-ray powder diffraction patterns of hydroxyapatite (HA) and silicon-substituted (Si-HA) powder after heat treatment at 1250 C: (a) pure HA, (b) Si0.4HA, (c) Si0.8HA, and (d) Si1.6HA (Bang et al., 2011).

It is notable that such substitution mechanism and the ionic radii difference affect the substituted HA lattice parameters (Leventouri et al., 2003; Aminian et al., 2011). The crystals get smaller and display a lower degree of crystallinity (Camaioni et al., 2015; Shepherd et al., 2012; Palard et al., 2009). The process results in highly hydrated surface layer surrounding the HA core crystals.

12.4.1 X-ray diffraction Numerous XRD analyses have been carried out to determine the effect of Si substitution on HA phase composition, crystallinity, and crystal lattice. Alieh Aminian et al. reported that with the increased SiO4 4 ions incorporation in the HA, a decrease in particle crystal size and crystallinity has been observed, and this consequently increases the solubility of the Si-HA powder (Aminian et al., 2011; El Yacoubi et al., 2014). With increasing Si content, the diffraction peaks lose intensity, proving a progressive loss of crystallinity. The broadening of the X-ray peaks also shows the decrease of crystallinity and crystallite size (Aminian et al., 2011; El Yacoubi et al., 2014; Bang et al., 2011) (Fig. 12.2). El Yacoubi et al. reported a morphological particle change from spheroidal for pure HA to elongated ellipsoidal crystals as Si substitution increases in HA (El Yacoubi et al., 2014). XRD analysis of the Si-HA has also revealed a single crystalline phase, similar to stoichiometric HA but with slightly larger crystal lattice constants and larger unit cell volume, the extent of which increases with increasing contents of Si substituted for phosphorus (Bang et al., 2011; Balamurugan et al., 2008). This increment appears logical considering that the ionic radius of Si4þ (0.042 nm) is larger than that of P5þ (0.035 nm) (Bang et al., 2011; Balamurugan et al., 2008; Aminian et al., 2011; El Yacoubi et al., 2014). As a result, the increase in the lattice parameters of

Handbook of Ionic Substituted Hydroxyapatites

Intensity (a.u)

292

32

(d) (c) (b) (a) 32.1

32.2 32.3 2θ (degree)

32.4

32.5

Figure 12.3 X-ray diffraction (XRD) peak shift pattern for silicon-substituted hydroxyapatite (Si-HA) samples with increasing silicon content: (a) pure HA, (b) Si0.4HA, (c) Si0.8HA, and (d) Si1.6HA (Bang et al., 2011).

Si-HA resulted in a slight shift of the Si-HA peaks to a lower Bragg’s angle compared with pure HA (Fig. 12.3) (Bang et al., 2011). Structural refinement studies of HA and Si-HA using XRD by El Yacoubi et al. and Porter et al. have also exposed that incorporation of Si into the HA lattice causes a decrease in the a-axis and an increase in the c-axis (Porter et al., 2004a; El Yacoubi et al., 2014).

12.4.2

Fourier-transform infrared spectroscopy

FTIR spectroscopy has been used to investigate the effect of the Si substitution on the different functional groups, such as OH and PO3 4 groups of HA. These investigations reported that the most prominent effect is the decrease in OH group’s number 4 with the increase in the substitution of PO3 4 by SiO4 to maintain the charge bal ance. A change in OH bending and stretching bonds at 630 and 3570 cm1, respectively, has also been reported with increased incorporated SiO4 4 ions into HA hexagonal structure, and even at some point, these both bands nearly faded away, which can be considered as the charge compensation mechanism for the extra negative charge of the SiO4 4 groups (Aminian et al., 2011; El Yacoubi et al., 2014; Bang et al., 2015). Furthermore, increased substitution of the SiO4 4 ions also results in loss of strong intensity of the CO2 (carbonate)-related absorption bands. These results 3 3 also support the idea that SiO4 substitutes PO tetrahedral in the HA structure 4 4 3  as CO2 ions occupy both PO and OH sites (El Yacoubi et al., 2014) 3 4 (Table 12.2).

12.4.3

Other characterization studies

It is well recognized that Si inclusion within the HA structure can strappingly influence the material surface charge (Camaioni et al., 2015). Tetravalent SiO4 4 ions replacing surface trivalent PO3 4 ions, confirmed by X-ray photoelectron spectroscopy analysis,

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Table 12.2 Fourier-transform infrared results of silicon-substituted hydroxyapatite (Si-HA) (El Yacoubi et al., 2014). FTIR peaks

Wave number (cmL1)

PeO stretching vibration modes

1100, 1034, and 962

OePeO bending mode

603e567

OH stretching and bending modes

3572 and 631

PO3 4

bending

603e565

OH stretching and bending bands

3572 and 631

CO2 3 vibration mode

1550e1410

result in decrement in net surface charge as well as in isoelectric point by shifting zeta potential values of Si-HA toward more negative values at physiological pH determined through zeta potential measurements by Botelho et al. (2002). PO3 4 substitution by  SiO4 4 also results in the loss of OH ions for charge compensation evidenced by FTIR and 1H NMR analyses that can be related to demonstrate Si incorporation by 29 Si NMR results (Szurkowska and Kolmas, 2017; Hayakawa et al., 2013). Solidstate NMR studies on HA with 5.6% Si substitution has also reported that not all SiO4 4 ions get substituted into the HA structure, but majority of ions were found to be present as silica gel units on the outer side of the HA lattice (Gasqueres et al., 2008; Tomoaia et al., 2014).

12.5

Silicon-substituted hydroxyapatite in coatings

Si-HA can be fabricated in various forms, i.e., porous scaffolds, granules, and coatings (Ratnayake et al., 2017). The demand for metal prostheses in orthopedic, maxillofacial, and dental surgery has tremendously increased during the last few decades (Rau et al., 2013). The main crucial aspect behind the success or failure of any metallic implant lies in its efficiency to bond quickly with the existing surrounding bone. Hence, to accelerate the process of osteointegration, the ideal metallic implant material should display compositional and structural similarity to those of native biological infrastructure. To achieve this aim, one of the most commonly used approaches is to coat the implant surface with a coating similar to the main constituent of mineralized bone with improved biological activity (Graziani et al., 2017). One of the most commonly used metal implants is titanium (Ti) and its alloys, which display exceptional mechanical properties, good biocompatibility, good fatigue resistance, and high corrosion resistant properties in comparison with other metals (Kim et al., 2014; Gomes et al., 2010; Rau et al., 2015), which makes it potentially more ideal to employ as successful implant for the load-bearing bone repair and regeneration applications. However, the problem with Ti implants is that it has very poor

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osteogenic properties that consequently results in its premature failure at implantation site. One solution to address this problem is to modify the Ti implant surfaces with osteoconductive or osteoinductive molecules (Gomes et al., 2010; Ho and Ding, 2015), e.g., Si-HA, which facilitates to induce and accelerate the bone tissue repair and regeneration process due to its compositional and crystallographic structural similarity to bone mineral (Zhang and Zou, 2009; Zhang et al., 2009b; San Thian et al., 2011). Thian et al. developed 1 mm thick Si-HA coatings with a silicon content up to 4.9 wt.% using magnetron cosputtering, and in vitro testing revealed that Si-HA coatings exhibited enhanced bioactivity and biofunctionality (Macdonald et al., 2012). Huang et al. showed that Si-HA coatings fabricated using electrostatic spray deposition promoted cell attachment and growth. They suggested that the enhanced dissolution rate of Si-HA with increasing silicon content might exert a stimulatory effect on cell activity in vitro (Hahn et al., 2010). Another promising novel metallic implant being used for orthopedic applications is of magnesium (Mg) and its alloys, which own reported properties of outstanding bioactivity, biodegradability, and satisfactory mechanical strength in vivo. The problem with Mg is that it is highly prone to corrosion due to hydrogen gas production in vivo and hence can evoke inflammatory response that consequently leads to the rapid degradability and hence the failure of the Mg implant without accomplishing the bone formation process (Qiu et al., 2014). Numerous approaches have been applied to control or even stop the corrosion rate of Mg by alloying elements or protective coatings, and these approaches must lead to a bioactive and nontoxic material. The most effectual approach for improving the corrosion resistance of Mg alloys is possibly by depositing an additional coating layer on their surfaces. Si-HA deposited as coating on the surface of the Mg and its alloys is used to achieve the controlled corrosion rate and hence degradation of Mg alloys in vivo physiological environments as well as to provide better bioactive surface to match the bone healing rate (Bogya et al., 2015; Qiu et al., 2014). Carbon/carbon (C/C) composites, composed of carbon fiber and carbon matrix, have also been used as orthopedic prosthetic material because of their unwavering stability in the physiological environment, excellent mechanical properties, and good intrinsic biocompatibility. Additionally, C/C composites have Young’s moduli (5e30 GPa) close to that of human bones (1e30 GPa), which can ensure excellent distribution and transmission of stress between implants and human bone. Despite of their good biocompatibility, the intrinsic bioinertness of C/C composites is still a limiting factor for their extensive biomedical applications. To solve this problem, some researchers have reported that coating a bioactive layer on C/C composites is a successful method. Among the bioactive coating materials applied to C/C composites, Si-HA has attracted huge interest due to its exceptional biocompatibility and bioactivity (Kezhi et al., 2017). So considering it, many efforts have been carried out to develop and fabricate Si-HA coatings on metallic implants (Hahn et al., 2010). Previous studies from the last couple of years elaborate several deposition methods for Si-HA coating on Ti/Ti alloys, Mg/Mg alloys, and on C/C composites. Currently, the plasma

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spraying (Bogya et al., 2015; Graziani et al., 2017; Gomes et al., 2010) is the only FDA-approved commercially available method for the HA coated implants production for biomedical applications (Rau et al., 2015). Other reported coating techniques are pulsed laser deposition (Qiu et al., 2014; Rau et al., 2013; Wan et al., 2015; Graziani et al., 2017; Lopez-Alvarez et al., 2009; Rau et al., 2015), magnetron sputtering (San Thian et al., 2011; Thian et al., 2007), electrolytic deposition (Kim et al., 2014), sol-gel (Rojaee et al., 2013), electrochemical deposition (Kezhi et al., 2017; Li et al., 2011), electrophoretic deposition (Xiao et al., 2008, 2009), aerosol deposition (Hahn et al., 2010), and biomimetic coatings (Zhang et al., 2009a; Zhang and Zou, 2009; Lilja et al., 2013; Ballo et al., 2012), which are being used to deposit Si-HA on metallic implants (Hahn et al., 2010).

12.6

Silicon-substituted hydroxyapatite in biomedical applications

Till date, the Si-HA biomedical implications have been inadequate. Many medical studies report the utilization of synthetic porous SieCaP bone graft substitute named Actifuse (0.8 wt.% Si-HA granules), commercially supplied in two size ranges (2e5 mm and 5e10 mm) by Apatech Ltd. These granules have biocompatibility properties, interconnected microporous and macroporous structure, and osteostimulatory capability. They can be applied to osseous deformities as well as to open bone voids related to periodontal, craniomaxillofacial, orthopedic, spinal, and oral defects. They can favorably restrict the autologous as well as allogenic grafts use to avoid immune response and microbial infection. They have been used for spinal surgical procedure all over thoracic, cervical, and lumbar spines, with 90% successful fusion rate without arising any explicit complexities (Nagineni et al., 2012). Actifuse granules have also been employed by Jenis and Banco (2010) as alternative bone graft in place of autologous iliac crest bone graft to treat patients with level 1 and level 2 lumbar degenerative disorders. These granules have also been embedded with vertebral bone marrow (BM) aspirates to employ as bone graft extensor for posterior remedial surgery on patients with adolescent idiopathic scoliosis with no reported detrimental effect or implant failure (Lerner and Liljenqvist, 2013). Another clinically proposed application of actifuse granules is in total joint athroplasty, where they were loaded with sheep’s BM-MSCs (mesenchymal stem cells) and cultured in an osteogenic perfusion and static system. A comparatively remarkable raise in ALP activity and cell proliferation in the perfusion system than in static system has been reported (Camaioni et al., 2015). Si-HA has also been considerably used as therapeutically drug delivery carrier in pharmaceutical industry alongside its mechanical supportive and osseointegrative stimulatory role. Regarding to this, mesoporous silica, instead of SiO4 4 ions, may be introduced into the Ap crystal lattice because of its capability of increasing the surface area and hence resultantly have tendency to attach more drug molecules for

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delivery. Furthermore, peptides, i.e., osteostatin (pentapeptide with both in vitro and in vivo stimulative osteogenic potential) as well as insulin can also be attached through covalent bonding or adsorption to Si-HA scaffolds. Silica/HA-filled composite resins with reported excellent biocompatibility and no toxicity have also been fabricated by Chadda et al. for dental applications (Chadda et al., 2016). Sych et al. successfully used HA substituted with 2 and 5 wt% Si for delivery of antibiotic “rifampicin,” which has established activity against leprosy, tuberculosis, and bacterial osteomyelitis (Sych et al., 2015). HA and mesoporous silica have also been employed to deliver bisphosphonates, i.e., alendronate and zoledronic acid, both suppress the osteoclasts activity to avoid bone resorption and stimulate bone development by osteoblasts. Drug release studies as well as in vitro biological activity studies on BM-MSC and osteoclast cell line showed more gradual release from the composite in contrast to the pure silica (Huang et al., 2011; Zhu et al., 2014; Szurkowska and Kolmas, 2017).

12.6.1

In vitro studies

For prior assessment of materials, many in vitro studies have been carried out to evaluate their osteogenic potential, biological response, and cytotoxicity. Si-HA displays synergistic effect on surface Ap layer formation in comparison with HA alone. In vitro studies of human osteosarcoma cells showed enhanced metabolic activity when seeded with Si-HA reported by Gibson et al. (1999b). Moreover, Thian et al. reported that human osteoblast-like cells (HOBC) seeded on Ti implant coated with SiHA displayed a higher rate of biomineralization, cell proliferation, and differentiation (Fig. 12.4) (Thian et al., 2007). Botelho et al. in their study reported that 0.8 wt% Si-HA stimulate human osteoblasts and thus increase the gene expression of ALP and osteocalcin (osteoblast marker genes) (Botelho et al., 2006a), and these results have also been confirmed by Honda et al. In addition, he also reported that Si-HA also results in an increased expression of runt-related transcription factor 2 that is the main marker gene involved in the formation and differentiation of osteoblast cells and thus involves in early bone developmental process (Honda et al., 2012). In another in vitro study, biological activity of 0.8 wt.% Si-HA was evaluated by Aminian et al. by seeding osteoblast cells on

(a)

(b)

10µm

(c)

10µm

10µm

Figure 12.4 Scanning electron microscopy (SEM) images of HOBs on various samples. (a) Monolayer of cells attaching on 0.8 wt.% silicon-substituted hydroxyapatite (Si-HA) with the formation of extracellular matrices at culture day 2; sign of mineralization on (b) 4.9 wt.% SiHA and (c) HA at culture day 42 (Thian et al., 2007).

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Si-HA and soaking them in simulated body fluid (SBF). He reported that Si-HA compared with HA displayed the improved biological activity, increased proliferation of osteoblast cells, as well as higher nucleation and growth rate on the surface of SBFimmersed samples (Aminian et al., 2011). Tian et al. and Huaguang et al. in two different studies reported that Si-HA immersed in SBF showed significantly increased bioactivity in vitro (Tian et al., 2008; Yu et al., 2017). To evaluate the Si-HA impact on osteoclast cell’s differentiation, in vitro studies has also been carried out by Botelho et al., who reported that when CD14þ peripheral human blood monocytes (PBMC) were seeded on Si-HA material, an increase in osteoclasts differentiation was observed, which could be linked with higher dissolution rate of Si-HA (Botelho et al., 2006b). In contrast, delayed osteoclasts differentiation as well as a decrement in their resorption activity has been reported by Matesanz et al. (2014) as well as Friederichs et al., who also seeded PBMC with nanocrystalline Si-HA. Recently, Casarrubios et al. have reported that the nanocrystallinity of the Si-HA also influences the Si-HAebone tissue interface, which consequently leads to loss of osteoclast apoptosis and cell anchorage (Casarrubios et al., 2016).

12.6.2 In vivo studies To study the surveillance and the systemic effects of the treatment in a better way, implants have been used in living organisms, which helps out to examine the interacting behavior of the implant material with surrounding tissues more closely and accurately. Patel et al. were the very first scientist who conducted in vivo studies by implanting SiHA granules into the femoral condyle defect of female sheep and rabbits. He observed an increased bioactivity behavior of Si-HA granules than HA alone (Patel et al., 2002, 2005). In another study, Porter et al. introduced Si-HA precipitate into defects in the sheep’s femoral condyles, and he found an increased Si-HA biological activity due to its good solubility property, and a higher concentration of Ca, P, and Si ions was also observed, which stimulates the accumulation of Ap layer on the implant’s surface (Porter et al., 2003; Porter, 2006). In another study, experiments on sheep were conducted to evaluate the effects of 1.5% Si-HA on the collagen fiber production. Presence of well-organized collagen fibrils altogether with good osteoconductive properties and an elevated level of nodular aggregates of HA crystals was observed after 6 and 12 weeks at the boneeSiHA interface and pure HA surface, respectively (Fig. 12.5) (Porter et al., 2004b). Many different wt% of Si-HA and their correlation with the rate of bone healing process on New Zealand white rabbit femur defect model was investigated by Hing et al., and it was observed that the 0.8 wt.% Si-HA showed the best osteogenic potential as well as higher mineral apposition rate (Hing et al., 2006). Si-HA granules (prepared form cuttlefish bone) with bone morphogenetic protein 2erelated peptides were also found to successfully heal the in vivo rabbit calvarial defect by stimulating bone healing process more rapidly (Kim et al., 2017). In other in vivo studies, Zhang et al. implanted Si-HA coated titanium implant material in the rabbit femur defect model, which displayed a higher rate of osteointegrative properties than HA alone (Zhang and Zou, 2009) (Fig. 12.6).

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(a)

(b) 200 nm

(c)

(d)

Figure 12.5 Transmission electron microscopy (TEM) micrographs of the bone/1.5 wt% Si-HA interface at 12 weeks in vivo. (a) Fibrous structures with the appearance of calcified collagen aligned parallel to the implant. (b) Crystallites, predominantly with platelike appearance, aggregated into nodular aggregates (a) adjacent to and separated from the Si-HA grains. (c) Collagen fibrils aligned both parallel (ls) and perpendicular (ts) to the implant surface. (d) Collagen fibrils aligned perpendicular (ts) to the implant surface and overlaying the synthetic Si-HA grains. Black arrows indicate regions of interface where collagen fibrils overlay the synthetic Si-HA grains (Porter et al., 2004b).

Another in vivo study was carried out by Ballo et al. who implanted Si-HA-coated Ti implant in rat tibia defect model, and they noticed comparatively higher bone growth at Si-HAebone tissue interface and thus stimulate new bone development process at earlier stage (day 7) in case of Si-HAecoated implant than HA-coated material (Ballo et al., 2012). In another in vivo study, it was confirmed that Si-HA implanted into an ovine defect model showed better biological activity by enhancing cellular infiltration, bone penetration, and neovascularization processes (Ratnayake et al., 2017). All these studies confirmed the improved in vivo biological performance of the Si-HA materials.

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Figure 12.6 Scanning electron microscope (SEM) microstructure interface between the new bone tissue and Si-HA and the interface between the Si-HA coating and Ti substrate, as indicated by the black line. The new bone was in tight contact with the Si-HA coating, which was deposited on the titanium fibers (Zhang and Zou, 2009).

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Effects of strontium substitution in synthetic apatites for biomedical applications

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Nujood Ibrahim Alyousef 1 , Yara Khalid Almaimouni 1 , Mashael Abdullah Benrahed 1 , Abdul Samad Khan 2 , Saroash Shahid 3 1 College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia; 2Department of Restorative Dental Sciences, College of Dentistry, Imam Abdulrahman Bin Faisal University, Dammam, Eastern Province, Saudi Arabia; 3 Centre for Oral Bioengineering, Institute of Dentistry, Queen Mary University of London, London, United Kingdom

13.1

Introduction

Strontium (Sr) is a metallic element naturally found in the bone and teeth (Zahra et al., 2014). It is effective in treating bone-related conditions and is commonly used to treat osteoporosis (Giusti, 2015; Kim et al., 2004). The incorporation of Sr in hydroxyapatite (HA) can have a major rule in enhancing the physiochemical and bioactive properties of HA (Ahmed, 2008). The use of strontium-substituted HA (Sr-HA) as a bioactive material has been reported in both medical and dental fields. In vivo and in vitro studies were conducted, where the material was reported to have good mechanical properties, biocompatibility, bioactivity, and an ability to stimulate new bone formation (Chung and Long, 2011; Yang et al., 2011; Xue et al., 2007). The effect of Sr-HA coating was examined, and it showed a comparable effect to HA coating (Xue et al., 2007). It was shown that partial substitution of Ca by Sr showed a higher solubility when compared with pure HA, which is due to the differences in the ionic radii of the cations. Furthermore, alkaline phosphate activity, which is a biomarker for bone formation, was found to be higher in Sr-HA of 10% as compared with pure HA shown in Fig. 13.1 (Zhang et al., 2011). It also exhibited a good osteoprecursor cell proliferation and attachment without adversely affecting the extracellular matrix formation and mineralization (Xue et al., 2006).

13.2

Method of preparation of strontium-substituted hydroxyapatite

The technique and method of preparing Sr-HA can lead to having different morphology and stoichiometry with various levels of crystallinity (Zahra et al., 2014). Several routes of preparing Sr-HA have been used, such as biomimetic Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00013-6 Copyright © 2020 Elsevier Ltd. All rights reserved.

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Alkaline phosphatase activity (umol/pnp/hr/mg protein)

10

8

HA Sr10-HA Sr40-HA Sr100-HA

6

*

*

*

4

2

0 Day 1

Day 3

Day 7

Figure 13.1 Response of human osteoblast-like cells to hydroxyapatite (HA) and different concentrations of strontium-substituted HA (Sr-HA). (*) Significant difference (Zhang et al., 2011).

approach in simulated body fluid (SBF) (Oliveira et al., 2007), hydrothermal (Zhang et al., 2011), wet chemical synthesis (O’Donnell et al., 2008), pulse laser deposition (Capuccini et al., 2008), different precipitation methods (Lei et al., 2017), solegel route (Mardziah et al., 2009; Renaudin et al., 2008), accelerated microwave process (Ravi et al., 2012), and many more. The precipitation method is considered to be the most economical method used for HA synthesis for biomedical applications (Zahra et al., 2014).

13.2.1

Hydrothermal method

In facile hydrothermal method, Sr-HA can be synthesized without the use of any surfactants or organic additives. In this process, Na2HPO4.12H2O is dissolved in deionized water. At the same time, aqueous solution of Ca2þ and Sr2þ is prepared by dissolving analytical grade reagents of Ca(NO3)2.4H2O, Sr(NO3)2 in deionized water. Then both solutions are mixed slowly by a magnetic bar stirrer. Ammonia solution can be used to increase the pH value above 10. After the hydrothermal reaction, the obtained white milky suspension can be packed and sealed in a Teflon-lined stainless steel autoclave and maintained at 200 C for 8 h. Finally, after cooling the material to room temperature, the precipitate is filtered and washed multiple times with deionized water and ethanol. The obtained material is then left to dry at 100 C for 6 h or at 120 C for 24 h or even left to dry overnight in air at 90 C (Xu et al., 2018; Xu et al., 2014; Zhang et al., 2011). The obtained Sr-HA material via the hydrothermal method was analyzed using Fourier-transform infrared spectroscopy and X-ray diffraction (XRD) techniques. Results showed that the Sr-HA material had a very similar crystal structure as compared with pure HA. Furthermore, scanning electron microscope (SEM) showed that the material has a rodlike morphological structure, with a length of

Effects of strontium substitution in synthetic apatites for biomedical applications

309

Figure 13.2 Morphology of strontium-substituted hydroxyapatite (SreHA) powders: (a) HA; (b) Sr10eHA 10; (c) Sr40eHA; and (d) Sr100eHA (Zhang et al., 2011).

140e320 nm when compared with the uniform nanobelt morphological shape of HA with 1.0e1.4 mm in length (Fig. 13.2) (Xu et al., 2018). In another study, the synthesized Sr-HA was confirmed to be fully apatitic. The study by Zhang et al. (2011) also showed that crystallinity of Sr-HA was inversely related to the increasing ratio of Sr in HA. On the contrary, an increase in Sr ratio to a certain degree can increase the cell activity. It was reported that osteoblasts were spontaneously attached and proliferated on the HA, Sr-HA10%, and Sr-HA40% surfaces, and none were found on Sr-HA100% (Zhang et al., 2011).

13.2.2 Solegel method In the solegel method, calcium and phosphorus precursors are mixed at a molecular level to synthesize HA (Tsai et al., 2018). After obtaining pure HA, ammonia solution (NH4OH) and ethylenediaminetetraacetic acid (EDTA) are dissolved by stirring the solution at 60 C. Furthermore, calcium nitrate tetrahydrate (Ca(NO3)2.4H2O) and strontium nitrate (Sr(NO3)2) are stirred in distilled water in two separate beakers by using a magnetic bar stirrer until the mixture becomes homogeneous. After that, all solutions are mixed where urea and diammonium hydrogen phosphate ((NH4)2HPO4) are added. This solution is kept under reflux at 100 C for 24 h to obtain a consistent gel, which is then dried under ambient static air to produce a solid material (black gel). Crushing the black gel results in a fine powder that is heated at 900 C to obtain a white powder.

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These particles can be added in specific percentages to HA to have a Sr-HA of certain wt%. The preparation of the substituted Sr-HA was detailed previously (Kanazawa, 1989). Higher concentration of Sr in HA leads to stabilizing the HA phase, suppressing the enlargement of HA crystals and decreasing the agglomeration size of Sr particles due to the retardation of Sr crystal growth (Mardziah et al., 2009). Notably, the synthesis of Sr-HA using the solegel yielded Sr2þ ions in both crystalline and amorphous phases of HA, i.e., b-tricalcium phosphate (b-TCP) and CaO and SrO phases (Renaudin et al., 2008).

13.2.3

Coprecipitation method

In coprecipitation method, preparation of Sr-HA can be achieved with different Ca/Sr ratios. To do so, Ca(NO3)2, Sr(NO3)2, and (NH4)2HPO4 should be dissolved in deionized water, to produce Ca2þ, Sr2þ, and PO3 4 solutions. Ammonia solution can be used to adjust the pH of the mentioned solutions to achieve a pH value of 10. At first, Ca2þ and Sr2þ solution should be mixed by stirring, while the temperature is gradually increased to 90 C. At this point, the (NH4)2HPO4 (a source of phosphate ions) is added. The obtained solution should be stirred at the same temperature for an hour and left to fully precipitate for 12 h. Subsequently, the precipitate material is filtered and washed several times to remove the unreacted material with both deionized water and ethanol. Subsequently, the material is dried at 60 C and sintered at 900 C for 4 h; alternatively, it can also be sintered at 120 C for 12 h (Lei et al., 2017; Li et al., 2007). The Sr-HA nanocrystals synthesized using the coprecipitation method showed a smaller particle size, with increased cell parameters. The diffraction peak position of the synthesized Sr-HA with different Ca/Sr ratios was calculated and is shown in Table 13.1 (Lei et al., 2017). When selected-area electron diffraction (SAED) was used on a single Sr-HA crystal, the crystallinity of the synthesized Sr-HA showed lower percentages (0.3% and 1.5%) as compared with pure HA, unlike 15% Sr-HA, which also showed a reduced crystallinity and the presence of an amorphous phase (Fig. 13.3) (Li et al., 2007).

Table 13.1 Cell parameters and crystal size of strontium-substituted HA (Sr-HA) samples with different ratios (Lei et al., 2017). Cell parameters Sample

a (Å)

b (Å)

c (Å)

a (8)

b (8)

g (8)

Volume(Å3)

Crystal size (nm)

HA

9.416

9.416

6.872

90

90

120

528.80

70.4

Sr-HA1%

9.454

9.454

6.911

90

90

120

534.68

68.8

Sr-HA5%

9.577

9.577

7.072

90

90

120

548.08

58.5

Sr-HA10%

9.767

9.767

7.267

90

90

120

600.98

46.7

Effects of strontium substitution in synthetic apatites for biomedical applications

(a) HA

(b) 0.3% Sr-HA

(d) 15% Sr-HA

(c) 1.5% Sr-HA

311

Figure 13.3 Selected-area electron diffraction (SAED) patterns of strontium-substituted HA (Sr-HA) with different ratios (Li et al., 2007).

13.2.4 Wet chemical synthesis Both HA and Sr-HAs can be synthesized by the wet chemical synthesis route (Abert et al., 2014). When using this method, two reaction routes can be used. In the first route, calcium nitrate-4-hydrate (Ca(NO3)24H2O) and strontium nitrate (Sr(NO3)2) should be dissolved in distilled water. Meanwhile, in the second solution, diammonium hydrogen orthophosphate ((NH4)2 HPO4) is dissolved in distilled water. Similarly, ammonia solution can be added to adjust the pH of both solutions to achieve the desired pH value. Both solutions are constantly stirred, and the second solution is then filtered and added to the first one and mixed to obtain a homogenous solution. Finally, when the precipitate is formed, it is stirred for an hour. Thereafter, the produced material can be centrifuged and washed with ultrapure or distilled water. After the purification step, the material is left to dry at 85 C for 20 h. The obtained material is crushed and can be sieved to obtain a fine powder with a known particle size range (Tavares et al., 2011; Abert et al., 2014; O’Donnell et al., 2008). In this method, no hydroxides are precipitated when ammonia is added to the diammonium hydrogen orthophosphate solution, which has lower solubility than amide salts, TCP, and apatite. In the second synthesis route (O’Donnell et al., 2008), calcium hydroxide (Ca(OH)2) and strontium hydroxide (Sr(OH)2) are used instead of calcium nitrate-4-hydrate and strontium nitrate. And the second solution is made with orthophosphoric acid (H3PO4) solution diluted in distilled water. The different steps in the synthesis of the material are similar to the steps described by O’Donnell et al. (2008). During the wet chemical synthesis, only a small fraction of impurities were found in the final material. Lattice parameters were found to be similar when compared with the same material synthesized by the other techniques (O’Donnell et al., 2008). Most importantly, no cytotoxicity was reported with the Sr-HA samples of 0%e5% substitution (Tavares et al., 2011).

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13.3 13.3.1

Handbook of Ionic Substituted Hydroxyapatites

Structure of strontium-substituted hydroxyapatite Crystallographic analysis of strontium-substituted apatite

Few studies have been conducted on the structural modifications of HA as well as on the incorporation of the metals into the HA lattice. XRD analysis of the samples synthesized at high temperatures by diffusion in the solid state showed that the lattice structure of HA has P63/m hexagonal symmetry, and as such, lattice parameters a ¼ b s c, a ¼ b ¼ 90 degrees, and g ¼ 120 degrees (Terra et al., 2009). The unit cell of crystalline HA hosts 10 cations that are arranged in two nonequivalent positions: M(1) is at the fourfold symmetry 4(f) position aligned in the column, each cation is surrounded by nine oxygen atoms, M(2) is at the sixfold symmetry 6(h) position set at the apexes of “staggered” equilateral triangles, surrounded by seven oxygen atoms (Elliott, 2013; Bigi et al., 2007;Terra et al., 2009). Considering the different positions occupied by calcium and oxygen atoms in the space group P63/m, its formula per unit cell Ca10(PO4)6(OH)2 can also be rewritten as Ca14Ca26(PO1O2O32)6(OHH)2. Ca2 ions are thus coordinated to hydroxyls; they form triangles centered on and perpendicular to the c-axis. Replication of this pattern along the c-axis gives rise to the OH channel, where the building unit is formed by two monopyramidal polyhedra, each with a triangular base occupied by three Ca2 with one OH at the apex and twisted by 60 degrees relative to each other. The hydroxyls and Ca1 atoms lie along columns parallel to the c-axis in the ratio of 1:2 throughout the crystal. Such OH: Ca1 columns in the hexagonal symmetry lead to a honeycomb structure centered at the OH columns with vertices alternately composed of O1 and O2 triangles centered on the Ca1 columns (Terra et al., 2009). The interest is to determine the occupation site of different metal ions in the apatite structure for better understanding of the flexibility of the structure (Bigi et al., 2007). Calcium can be replaced by strontium in different Ca sites in the HA structure. It is known that strontium and other divalent cations with a similar charge-to-size ratio to calcium can readily substitute in the structure of HA (O’Donnell et al., 2008). In recent studies, Sr-HA Sr10(PO4)6(OH)2 has received the vast attention of researchers because of its crystallochemical similarity to calcium HA (Bigi et al., 2007). The Sr-HA solid solution, which can be obtained by hydrothermal methods or by treatment at high temperatures, shows a linear variation in the lattice parameters with the composition (Zhu et al., 2006). Increasing the amount of strontium that substitutes calcium in the HA structure leads to a linear variation in the infrared absorption bands and in the cell parameters, with the increase of mean dimensions of the cation. The effect of strontium on morphology and crystallinity changes with composition, low Sr replacement to calcium leads to a distortion in the shape of the crystals and decreases in the coherent length of the perfect crystalline, the mean dimensions of the crystals increase with a high strontium content. Similarly, although strontium distribution is similar in the two cationic sites, a slight Sr increase in the M(2) site triumphs in most of the range of composition, while a modest preference of Sr in the smaller M(1) site can be appreciated only with a very low strontium content (Bigi et al., 2007). As the Sr (1.12) ionic

Effects of strontium substitution in synthetic apatites for biomedical applications

313

radius is larger than that of Ca (0.99), a rigid ion model would suggest a corresponding enlargement of the SreO distances in the Sr1O13O23 and Sr2O1O2O34(OH) polyhedral (Terra et al., 2009).

13.3.2 Structural analysis of strontium-substituted apatite Line profile analysis has been used for the investigation of the line broadening increase and the peak shifts detected in the XRD patterns of the samples (Fig. 13.4).

(f)

(e)

(d)

(c)

(b)

(a)

10

20

30

40

50

60

2θ angle (degrees)

Figure 13.4 Powder X-ray diffraction patterns of (a) Sr0; (b) Sr10; (c) Sr30; (d) Sr50; (e) Sr70; and (f) Sr100 (Bigi et al., 2007).

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Table 13.2 Coherent lengths (shkl) of the perfect crystalline domains in the direction normal to 002 and to 310 planes calculated using the Scherrer method (Bigi et al., 2007). Samples

s002 (Å)

s310 (Å)

Sr0

469 (5)

224 (6)

Sr3

408 (1)

222 (4)

Sr5

396 (1)

217 (4)

Sr10

385 (5)

153 (6)

Sr20

212 (7)

92 (7)

Sr30

168 (4)

86 (8)

Sr50

113 (4)

93 (5)

Sr70

293 (3)

190 (3)

Sr90

391 (6)

230 (2)

Sr100

490 (5)

305 (3)

An estimation of the size of the crystallite coherently scattering domain as derived from the Scherrer equation, on the hypothesis of insignificant microstrain, is reported in Table 13.2. The s002 is correlated to the mean crystallite size in the c-axis, while s310 refers to the mean crystal size in a direction perpendicular to it. The crystallite sizes decrease as cSr increases with strontium content smaller than 50%. Yet, both s002 and s310 decrease, but the contraction seems to be somewhat greater along the c-axis direction. As a matter of fact, the observed maximum contraction is of about 75% along the 002 direction and of about 60% along the 310 direction. For higher Sr contents, crystallite sizes progressively increase to reach the maximum values in SreHA (Bigi et al., 2007). The XRD traces of the Sr-substituted apatites and Rietveld refinement show a diffraction peaks shift to lower 2q values when Sr is added, indicative of an increase in the d-spacings and hence lattice parameters. This would be expected as Sr is slightly larger than Ca (118 and 100 p.m.). With Sr addition, peak intensity also increases. This is to be predicted as Sr is heavier and contains more electrons than Ca and will, therefore, more effectively scatter X-rays, so there is a general increase in crystallinity with Sr addition (O’Donnell et al., 2008). Lattice parameters (a and c) increase linearly with increasing Sr content, as shown in Figs. 13.5 and 13.6, which is consistent with larger ion entering the apatite lattice. Because of the increase in lattice parameters, the unit cell volume also increases linearly. The crystal density also increases linearly with the addition of Sr due to the replacement of a heavier ion for Ca in the crystal structure (87.62 and 40.078 g mol1, respectively) (Greenwood and Earnshaw, 2012). Fig. 13.7 shows a linear increase in the density of fluorapatite with increasing fraction of Sr in the phase (O’Donnell et al., 2008). As fluoride and OH have roughly the same mass and ionic radii, the density of these materials should not differ greatly from Sr-HA.

9.80 9.75 y = 0.35x + 9.41 R2 = 0.99

9.70

a/Å

9.65 9.60 9.55 9.50 9.45 9.40 0.0

0.2

0.4

0.6

0.8

1.0

Fraction Sr

Figure 13.5 Variation in lattice parameter, a, for the series (SrxCa1x)5(PO4)3OH, where x ¼ 0.00, 0.25, 0.50, 0.75, and 1.00; error bars: þ2s from Table 13.3 (O’Donnell et al., 2008). 7.30 y = 0.42x + 6.86 R2 = 0.99

7.25 7.20

c/Å

7.15 7.10 7.05 7.00 6.95 6.90 6.85 0.0

0.2

0.4

0.6

0.8

1.0

Fraction Sr

Figure 13.6 Variation in lattice parameter, c, for the series (SrxCa1x)5(PO4)3OH, where x ¼ 0.00, 0.25, 0.50, 0.75, and 1.00; error bars: 2s from Table 13.3 (O’Donnell et al., 2008). 4.2 y = 0.93x + 3.18 R2 = 0.99

ρ / g.cm–3

4.0 3.8 3.6 3.4 3.2 3.0 0.0

0.2

0.4

0.6

0.8

1.0

Fraction Sr Rietveld

Literature

Figure 13.7 Variation in density for the series (SrxCa1x)5(PO4)3OH, where x ¼ 0.00, 0.25, 0.50, 0.75 and 1.00; error bars 1% (O’Donnell et al., 2008).

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Table 13.3 Parameters from Rietveld refinement of the series (SrxCa1x)5(PO4)3OH, where x ¼ 0.00, 0.25, 0.50, 0.75 and 1.00, with standard deviations (O’Donnell et al., 2008). Sr

Ca

a /Å

2sa / Å

c /Å

2sc / Å

V/ Å3

2sV / Å3

r/ g.cmL3

d/ nm

0.00

1.00

9.411

0.004

6.877

0.003

527.5

0.378

3.163

17.55

0.25

0.75

9.505

0.007

6.950

0.006

543.8

0.648

3.384

11.37

0.50

0.50

9.596

0.004

7.054

0.004

562.6

0.452

3.692

9.38

0.75

0.25

9.659

0.005

7.182

0.004

580.3

0.740

3.897

12.94

1.00

0.00

9.777

0.003

7.288

0.003

603.3

0.472

4.068

20.50

Table 13.4 Occupancies of M(I) and M(II) sites from Rietveld refinement (O’Donnell et al., 2008). x (Sr) Site

0.00

0.25

0.50

0.75

1.00

Ca(I)

1.00

0.81

0.52

0.31

0.00

Sr(I)

0.00

0.19

0.48

0.69

1.00

Ca(II)

1.00

0.75

0.45

0.20

0.00

Sr(II)

0.00

0.25

0.55

0.80

1.00

Mean Sr

0.00

0.22

0.52

0.74

1.00

The Rietveld refinement gives a good fit with a slight preference for strontium occupancy on the Ca(II) site. There is a 6%e12% preference of Sr for the M(II) site, with increasing preference as the Sr content increases. The occupancies of the sites can be seen in Table 13.4 (O’Donnell et al., 2008).

13.4 13.4.1

Biomedical applications of strontium-substituted hydroxyapatite Orthopedic and bone tissue regeneration

Strontium resembles calcium as 98% of it concentrated mostly in bone tissue exhibiting strong bone-seeking properties especially in new trabecular bone (Li et al., 2010; Tao et al., 2016; Landi et al., 2007). Strontium is known for its ability to stimulate bone formation and hinder bone resorption as it promotes preosteoblastic cell replication and osteoblast proliferation and prevents osteoclasts maturation and promotes their apoptosis in a dose-dependent effect (O’Donnell et al., 2008; Diepenhorst et al., 2018; Boanini et al., 2011; Raucci et al., 2015). Moreover, it has a mild antibacterial effect that aids in promoting osteoblast activity (Sriranganathan et al., 2016). When compared with calcium HA, Sr-HA revealed higher biocompatibility and bioactivity (Huang et al., 2019). Addition of Sr to HA provides an antibacterial effect and

Effects of strontium substitution in synthetic apatites for biomedical applications

317

enhances the radiopacity of the material (Panpisut et al., 2016). Strontium has an effect on the lattice parameter of Sr-HA in bone minerals, and it also stabilizes HA precursor € phases (Ozbek et al., 2016; Panpisut et al., 2016). Polyetheretherketone (PEEK) composites showed great mechanical properties along with being chemically inert, biocompatible, and nontoxic. Moreover, it is a radiolucent, sterilization-resistant, and easy-to-process material. Sr-HA can be added to PEEK to further promote their bioactivity and enhance mechanical properties including strength and modulus of elasticity. Compared with PEEK, 25 vol% and 30 vol% Sr-HAereinforced PEEK revealed a greater bending modulus with 113% and 136% increase, respectively, in an in vitro study. PEEK, HA/PEEK, and Sr-HA/PEEK were immersed in SBF for 14 days. PEEK showed no apatite layer formation while both HA/PEEK and Sr-HA/PEEK revealed apatite formation, whereas more apatite was formed on the Sr-HA/PEEK surface with greater thickness as it was evident in the images obtained from SEM in Fig. 13.8. The calcium deposit was treated with Alizarin Red stain. The increase of the stained area was significantly higher in the Sr-HA/PEEK composite compared with the other groups as shown in Fig. 13.9 (Wong et al., 2009).

(a)

Mag = 5.00 KX 10 µm

(b)

EHT = 5.00 kV Photo No. = 550

Signal A =SE2 WD = 11 mm Date : 7 Apr 2009 Time : 14:46:11

(c)

Mag = 12.00 KX 1 µm

Mag = 5.00 KX

10 µm

EHT = 5.00 kV Photo No. = 572

WD = 11 mm Signal A =SE2 Date : 7 Apr 2009 Time : 16:10:55

1 µm

EHT = 5.00 kV Photo No. = 528

Signal A =SE2 WD = 9 mm Date : 6 Apr 2009 Time : 14:23:46

(d)

EHT = 5.00 kV Photo No. = 513

Signal A =SE2 WD = 9 mm Date : 6 Apr 2009 Time : 11:61:34

Mag = 12.00 KX

Figure 13.8 Apatite formation after immersion of the samples in SBF for 1 day in the surface of (a) 25 vol% Sr-HA/PEEK composite and (b) 25 vol% HA/PEEK. The apatite layer thickness after immersion of the samples in SBF for 14 days observes for (c) 25 vol% Sr-HA/PEEK composite and (d) 25 vol% HA/PEEK composite (Wong et al., 2009).

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% increase of stained area

40

PEEK Sr-HA/PEEK HA/PEEK

30

20

10

0 PEEK

Sr-HA/PEEK

HA/PEEK

Figure 13.9 Polyetheretherketone (PEEK), strontium-substituted hydroxyapatite (Sr-HA)/ PEEK, and HA/PEEK percentage of stain area increase (Wong et al., 2009).

Treating large bony defects can be difficult due to the compromised vascularization. By activating calcium-sensing receptors, strontium-containing HA stimulates not only osteogenesis but angiogenesis as well. The in vitro study stated that regardless of the Sr rate in the Sr-HA, no cytotoxic effect on the mesenchymal stem cells (MSCs) was observed (Ehret et al., 2017). Osteogenesis genes in MSC can be activated by Sr-HA, and Sr2þ improves alkaline phosphatase (ALP) activity of MSC (Yuan et al., 2018; Raucci et al., 2015). Sr-HA with 5 mol% of Sr showed a higher MSC proliferation when compared with HA scaffold or when compared with Sr-HA with higher mol% of Sr (10, 15, and 20 mol%) (Raucci et al., 2015). In a study conducted on mice, the sample with 50% Sr substitution proved to generate a significantly higher amount of mineralized tissue when compared with the sample with 8% Sr (Sr-HA8%) substitution. After the 4 weeks, the sample with 50% Sr (Sr-HA50%) substitution (highest Sr substitution) revealed a very positive effect on angiogenesis because it showed a higher number of new blood vessels as compared with the low Sr sample (Sr-HA8%) and pure HA matrices (Ehret et al., 2017).

13.4.2

Implant coating

Sr-HA can be used as a coating on titanium (Ti) and titanium alloy implants using different techniques including solegel dip-coating method, pulsed-laser deposition, plasma spray, and biomimetic method (Li et al., 2010). Addition of Sr improves both the mechanical and biological properties of HA (Kaygili and Keser, 2015). In an in vivo study conducted on rats, it was reported that implants with Sr-HA10% coating were better than HA coating in terms of implant-to-bone contact and preimplant bone mass. Moreover, it showed lower trabecular spaces along with increased fixation, push-out strength by 107.2%, and shear strength by 132.9%, as the push-out force for HA-coated implant is 38.7  5.3 N, while the Sr-HA10% push-out force was 80.2  9.9 N. In addition, Sr-HA10% showed surface roughness, coating thickness, phase composition, and adhesion strength that are comparable to that of pure HA coating (Li et al., 2010). In a study conducted on rats with ovariectomy-induced osteoporosis,

Effects of strontium substitution in synthetic apatites for biomedical applications

(a)

100 μm

319

(c) 60 Bone area ratio

50

Implant

40 30 20 10 0

100 μm

Implant

10%SrHA

HA

10%SrHA

80 Bone-to-implant contact

(b)

HA

(d) 70 60 50 40 30 20 10 0

Figure 13.10 Histological evaluation obtained 12 weeks postimplantation shows implants and proximal tibia, whereas (a) HA-coated Ti implant, (b) 10% Sr-HAecoated Ti implants and (c) comparative bone area ratio, and (d) comparative bone-to-implant contact (Li et al., 2010).

Sr-HA coating was not only 1.84 folds better in push-out strength than HA, but it was also significantly higher than other coating materials including Zn-HA and Mg-HA. Moreover, micro-CT assessment of the former groups proved Sr-HA superiority in the following parameters: bone volume per total volume, trabecular number, connective density, and thickness (Tao et al., 2016). Histological analysis of HA-coated Ti implants using a 10% sample (Sr-HA10%) showed a significantly higher bone to area ratio (by 70.9%) and bone-to-implant contact (by 49.9%) among ovariectomized rats (Fig. 13.10) (Li et al., 2010). ALP expression was the highest for Sr-HA and Sr/AGHA coating material when compared with pure HA and Ag-HA coating, proving the positive effect of Sr-HA on osteoblast differentiation (Fielding et al., 2012).

13.4.3 Osteoporosis treatment Osteoporosis does not only increase the risk of fractures and bone pathologies but it may also negatively affect the success rate of bone reconstructing surgeries (Boanini et al., 2011). Strontium is used for osteoporosis treatment to prevent vertebral fracture among postmenopausal women (O’Donnell et al., 2008). Strontium is shown to be deposited in the newly formed bone and strontium ranelate (SrR) preserves mineralization among postmenopausal women treated with SrR for 3 years (Li et al., 2010). However, Sr effect is dose dependent as low concentration starting from 2 to 10 mg/ mL stimulate bone formation, but when the concentration is in the range of 20e100 mg, it leads to a defective mineralization (Suganthi et al., 2011; Ehret et al.,

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2017). Strontium can downgrade secretion of inflammatory cytokines including TNFa, IL-1b, and IL-6, and as a result of that, it downgrades osteoclastogenesis and, furthermore, prevents bone resorption. In an in vivo study conducted on rats, an injectable Sr-doped HA material with G3-K PS (15 mol% Sr) was tested to treat osteoporotic bone defects and compared this material with HA, Sr-HA15%, and HA/G3-K PS. Sr-containing samples showed the highest suppression of IL-1b and TNF-a with 15Sr-HA/G3-K PS revealing the lowest IL-1b and TNF-a gene expression of all samples. Thus, 15Sr-HA/G3-K PS can be used for treating osteoporotic bone defects as it revealed inflammatory cytokine suppressive effect, good injectability, and antiwashout property (Yuan et al., 2018). Sr-HA showed positive results for treating knee osteoarthritis as it showed clinical improvement as well as downgraded the radiographic progression when 2 g/day dosage was used (Han et al., 2017).

13.4.4 Enamel repair Among the low caries population, strontium content in enamel is found to be significantly higher than Sr content in high caries population (Landi et al., 2007; Surdacka et al., 2007). Sr stabilizes appetite structure by increasing its resistance to bacterial acids and eventually dental caries (Landi et al., 2007). White spot lesions may result from orthodontic treatment especially in maxillary anterior teeth increasing the risk of enamel caries. Loss of enamel may result due to debonding of orthodontic appliance, whereas approximately 55.6 mm of enamel is lost during this process. Nano-HA (nHA) is used to treat white spot lesions because it promotes mineralization from within the teeth and as it lacks downsides associated with fluoride such as hypermineralization of the surface layers and fluoride side effects such as gastrointestinal problems, chronic diseases, and skeletal fluorosis. Despite nHA advantages, it revealed several downsides such as weakness, brittleness, high degree of crystallinity, and low solubility at neutral pH. Sr-HA of 10% or more has a higher solubility, reactivity, bioactivity, and fluoride release when compared with HA. Moreover, the addition of Sr can influence the degree of crystallinity, phase stability, and reactivity. In an in vitro study conducted in extracted human teeth, different concentration of Sr in HA used to treat enamel white lesions. The Sr-HA with higher concentration of Sr (50 mol%) showed the highest hardness when compared with a pure HA or a Sr-HA with lower Sr concentration (Krishnan et al., 2016). Sr-HA was added to toothpaste to increase Sr content in dental enamel (Lippert and Hara, 2013; Surdacka et al., 2007). After 4 months of regular brushing with a Sr-HA containing toothpaste, Sr content was significantly increased and Sr reached deeper layers of enamel reaching as deep as 60 mm (Surdacka et al., 2007).

13.4.5 Guided bone regeneration Guided bone regeneration (GBR) is vital for preventing soft tissue cells from migrating to the bony defect, allowing the osteogenic and angiogenic cells to occupy the bony defect. There are resorbable and nonresorbable types of membranes and both types showed similar clinical outcomes, but the resorbable types eliminate the need for a

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second surgery to remove the membrane. Nevertheless, to prolong collagen membrane resorption, several cross-linking reagents were used. These cross-linking reagents include glutaraldehyde, diphenylphosphorylazide, and hexamethylene diisocyanate, which may reveal a cytotoxic effect, and moreover, it may compromise cell’s initial attachment and proliferation. Gelatin is a biocompatible material, but due to its poor mechanical properties, pure gelatin cannot be used as a membrane for GBR. In a study conducted on rats, strontium HA was added to gelatin to synthesize a resorbable membrane with an ability to stimulate bone formation. Sr-HA included gelatin membrane showed a significantly higher ultimate strength and tensile strength. Moreover, maximum load and tensile strength were increased when Sr-HA content increased. At 4 weeks, significantly greater bone formation and complete closure of the bony defect was observed in a radiographic evaluation for Sr-HA with 20 mg/mL Sr group compared with the collagen and 10 mg/mL Sr groups (Hao et al., 2013).

13.4.6 Drug carrier The chemical composition of HA results in a strong surface adsorption capability (Xu et al., 2018). Addition of strontium to HA increases surface area allowing drug release for a longer period of time (Suganthi et al., 2011). In an in vitro study, Sr-HA (1 wt.% Sr) adsorption capacity of bovine serum albumin was found to be 14.64% higher than HA. Lysosomes loading capacity was found to be three times higher in Sr-HA15% when compared with HA. Moreover, sustained protein release was observed in Sr-HA (Xu et al., 2018). In an in vitro study, mesoporous Sr-HA nanorods were used as ibuprofen carrier, and it revealed a loading capacity of 32.9 wt%. Impregnation process allows ibuprofen adsorption onto the mesoporous surface, while diffusioncontrolled mechanism allows its release. Almost 50.5% of ibuprofen was released after 3 h. Ibuprofen release reaches about 93.9% after 12 h, and at 24 h, there is no further release (Fig. 13.11) (Zhang et al., 2010).

Cummulative released IBU (%)

100 80 60 40 20 0 0

5

10

15 20 Time (h)

25

30

35

Figure 13.11 Strontium-substituted hydroxyapatite (Sr-HA) nanorods ibuprofen release (Zhang et al., 2010).

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In an in vitro study, Sr-HA-CaO-CO3 nanofibers exhibiting a mesoporous structure (mSr-HANFs) were fabricated as a potential carrier for tetracycline. Degradation ratio of mSr-HANFs was increased when Sr content increased. The 3mSr-HANFs’s tetracycline loading efficacy is about 97.21%. Tetracycline release was slow and steady with a rate of 2.36% per day coordinating with 3mSr-HANFs degradation rate. The 3mSr-HANFs sustained tetracycline release, and its antibacterial effect lasted for more than 3 weeks (Tsai et al., 2018).

References Abert, J., Bergmann, C., Fischer, H., 2014. Wet chemical synthesis of strontium-substituted hydroxyapatite and its influence on the mechanical and biological properties. Ceram. Int. 40, 9195e9203. Ahmed, S.A.A.,M., 2008. Synthesis of Ca-hydroxyapatite bioceramic from egg shell and its characterization. Bangladesh J. Sci. Ind. Res. 43, 501e512. Boanini, E., Torricelli, P., Fini, M., Bigi, A., 2011. Osteopenic bone cell response to strontiumsubstituted hydroxyapatite. J. Mater. Sci. 22, 10. Bigi, A., Boanini, E., Capuccini, C., Gazzano, M., 2007. Strontium-substituted hydroxyapatite nanocrystals. Inorg. Chim. Acta 360, 1009e1016. Capuccini, C., Torricelli, P., Sima, F., Boanini, E., Ristoscu, C., Bracci, B., Socol, G., Fini, M., Mihailescu, I.N., Bigi, A., 2008. Strontium-substituted hydroxyapatite coatings synthesized by pulsed-laser deposition: in vitro osteoblast and osteoclast response. Acta Biomater. 4, 9. Chung, C.-J., Long, H.-Y., 2011. Systematic strontium substitution in hydroxyapatite coatings on titanium viamicro-arc treatment and their osteoblast/osteoclast responses. Acta Biomater. 7, 7. Diepenhorst, N.A., Leach, K., Keller, A.N., Rueda, P., Cook, A.E., Pierce, T.L., Nowell, C., Pastoureau, P., Sabatini, M., Summers, R.J., Charman, W.N., Sexton, P.M., Christopoulos, A., Langmead, C.J., 2018. Divergent effects of strontium and calciumsensing receptor positive allosteric modulators (calcimimetics) on human osteoclast activity. Br. J. Pharmacol. 175, 14. Ehret, C., AID-Launais, R., Sagardoy, T., Siadous, R., Bareille, R., Rey, S.P.,S., Etienne, L., Kalisky, J., DE Mones, E., Letourneur, D., Amedee Vilamitjana, J., 2017. Strontium-doped hydroxyapatite polysaccharide materials effect on ectopic bone formation. PLoS One 12. Elliott, J.C., 2013. Structure and Chemistry of the Apatites and Other Calcium Orthophosphates. Elsevier. Fielding, G.A., Roy, M., Bandyopadhyay, A., Bose, S., 2012. Antibacterial and biological characteristics of silver containing and strontium doped plasma sprayed hydroxyapatite coatings. Acta Biomater. 8, 9. Giusti, A.A.B.,G., 2015. Treatment of primary osteoporosis in men. Clin. Interv. Aging 10, 105. Greenwood, N.N., Earnshaw, A., 2012. Chemistry of the Elements. Elsevier. Han, W., Fan, S., Bai, X., Ding, C., 2017. Strontium ranelate, a promising disease modifying osteoarthritis drug. Expert Opin. Investig. Drugs 26, 6.

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Hao, J., Acharya, A., Chen, K., Chou, J., Kasugai, S., Lang, N.P., 2013. Novel bioresorbable strontium hydroxyapatite membrane for guided bone regeneration. Clin. Oral Implant. Res. 26, 7. Huang, J., Huang, Z., Yao, D., Wu, C., Cheng, Y., Wu, F., 2019. Crystallization process and growth mechanism for hexagonal prism of strontium hydroxyapatite by urea hydrolysis. J. Cryst. Growth 512, 7. Kanazawa, T., 1989. Inorganic Phosphate Materials, vol. 52. Elsevier Science Ltd, p. 288. Kaygili, O., Keser, S., 2015. Solegel synthesis and characterization of Sr/Mg, Mg/Zn and Sr/Zn co-doped hydroxyapatites. Mater. Lett. 141, 4. Kim, H.W., Koh, Y.H., Kong, Y.M., Kang, J.G., Kim, H.E., 2004. Strontium substituted calcium phosphate biphasic ceramics obtained by a powder precipitation method. J. Mater. Sci. Mater. Med. 15, 1129e1134. Krishnan, V., Bhatia, A., Varma, H., 2016. Development, characterization and comparison of two strontium doped nano hydroxyapatite molecules for enamel repair/regeneration. Dental Mater. 32, 14. Landi, E., Tampieri, A., Celotti, G., Sprio, S., Sandri, M., Logroscino, G., 2007. Sr-substituted hydroxyapatites for osteoporotic bone replacement. Acta Biomater. 3, 9. Lei, Y., Xu, Z., Ke, Q., Yin, W., Chen, Y., Zhang, C., Guo, Y., 2017. Strontium hydroxyapatite/ chitosan nanohybrid scaffolds with enhanced osteoinductivity for bone tissue engineering. Mater. Sci. Eng. C 72, 9. Li, Z.Y., Lam, W.M., Yang, C., Xu, B., Ni, G.X., Abbah, S.A., Cheung, K.M.C., Luk, K.D.K., Lu, W.W., 2007. Chemical composition, crystal size and lattice structural changes after incorporation of strontium into biomimetic apatite. Biomaterials 28, 1452e1460. Li, Y., Li, Q., Zhu, S., Luo, E., Li, J., Feng, G., Liao, Y., Hu, J., 2010. The effect of strontiumsubstituted hydroxyapatite coating on implant fixation in ovariectomized rats. Biomaterials 31, 9. Lippert, F., Hara, A.T., 2013. Strontium and caries: a long and complicated relationship. Caries Res. 47, 17. Mardziah, C.M., Sopyan, I., Ramesh, S., 2009. Strontium-doped hydroxyapatite nanopowder via sol-gel method: effect of strontium concentration and calcination temperature on phase behavior. Artif. Organs 23, 105e113. Oliveira, A.L., Reis, R.L., Li, P., 2007. Strontium-substituted apatite coating grown on Ti6Al4V substrate through biomimetic synthesis. J. Biomed. Mater. Res. B Appl. Biomater. 83, 258e265. € € Ozbek, Y.Y., Bas¸tan, F.E., Ustel, F., 2016. Synthesis and characterization of strontium-doped hydroxyapatite for biomedical applications. J. Therm. Anal. Calorim. 125, 6. O’donnell, M.D., Fredholm, Y., Rouffignac, A.D., Hill, R.G., 2008. Structural analysis of a series of strontium-substituted apatites. Acta Biomater. 4, 10. Panpisut, P., Liaqat, S., Zacharaki, E., Xia, W., Petridis, H., Young, A.M., 2016. Dental composites with calcium/strontium phosphates and polylysine. PLoS One 11, 20. Raucci, M.G., Giugliano, D., Alvarez-Perez, M.A., Ambrosio, L., 2015. Effects on growth and osteogenic differentiation of mesenchymal stem cells by the strontium-added solegel hydroxyapatite gel materials. J. Mater. Sci. 26, 11. Ravi, N.D., Balu, R., Sampath Kumar, T., 2012. Strontium-substituted calcium deficient hydroxyapatite nanoparticles: synthesis, characterization, and antibacterial properties. J. Am. Ceram. Soc. 95, 2700e2708.

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Renaudin, G., Laquerriere, P., Filinchuk, Y., Jallot, E., Nedelec, J.M., 2008. Structural characterization of solegel derived Sr-substituted calcium phosphates with anti-osteoporotic and anti-inflammatory properties. J. Mater. Chem. 18, 3593e3600. Sriranganathan, D., Kanwal, N., Hing, K.A., Hill, R.G., 2016. Strontium substituted bioactive glasses for tissue engineered scaffolds: the importance of octacalcium phosphate. J. Mater. Sci. Mater. Med. 27. Suganthi, R.V., Elayaraja, K., Ahymahjoshy, M.I., Sarathchandra, V., Girija, E.K., Kalkura, S.N., 2011. Fibrous growth of strontium substituted hydroxyapatite and its drug release. Mater. Sci. Eng. C 31, 7. Surdacka, A., Stopa, J., Torlinski, L., 2007. In situ effect of strontium toothpaste on artificially decalcified human enamel. Biol. Trace Elem. Res. 116, 7. Tao, Z.-S., Zhou, W.-S., He, X.-W., Liu, W., Bing-Libai, Zhou, Q., Huang, Z.-L., Tu, K.-K., Li, H., Sun, T., Lv, Y.-X., Cui, W., Yang, L., 2016. A comparative study of zinc, magnesium, strontium-incorporated hydroxyapatite-coated titanium implants for osseointegration of osteopenic rats. Mater. Sci. Eng. C 62, 7. Tavares, D.D.S., Resende, C.X., Quitan, M.P., Castro, L.D.O., Granjeiro, J.M., Soares, G.D.A., 2011. Incorporation of strontium up to 5 Mol.(%) to hydroxyapatite did not affect its cytocompatibility. Mater. Res. 14, 456e460. Terra, J., Dourado, E.R., Eon, J.-G., Ellis, D.E., Gonzalezd, G., Rossi, A.M., 2009. The structure of strontium-doped hydroxyapatite: an experimental and theoretical study. Phys. Chem. Chem. Phys. 11, 568e577. Tsai, S.W., Yu, W.X., Hwang, P.A., Huang, S.S., Lin, H.M., Hsu, Y.W., Hsu, F.Y., 2018. Fabrication and characterization of strontium-substituted hydroxyapatite-CaO-CaCO3 nanofibers with a mesoporous structure as drug delivery carriers. Pharmaceutics 10, 179. Wong, K.L., Wong, C.T., Liu, W.C., Pan, H.B., Fong, M.K., Lam, W.M., Cheung, W.L., Tang, W.M., Chiu, K.Y., Luk, K.D.K., Lu, W.W., 2009. Mechanical properties and in vitro response of strontium-containing hydroxyapatite/polyetheretherketone composites. Biomaterials 30, 8. Xu, J., Yang, Y., Wan, R., Shen, Y., Zhang, W., 2014. Hydrothermal preparation and characterization of ultralong strontium-substituted hydroxyapatite whiskers using acetamide as homogeneous precipitation reagen. Sci. World J. 2014. Article ID 863137. Xu, Y., An, L., Chen, L., Xu, H., Zeng, D., Wang, G., 2018. Controlled hydrothermal synthesis of strontium-substituted hydroxyapatite nanorods and their application as a drug carrier for proteins. Adv. Powder Technol. 29, 1042e1048. Xue, W., Moore, J.L., Hosick, H.L., Bose, S., Bandyopadhyay, A., Lu, W.W., Cheung, K.M., Luk, K.D., 2006. Osteoprecursor cell response to strontium-containing hydroxyapatite ceramics. J. Biomed. Mater. Res. A 79, 804e814. Xue, W., Hosick, H.L., Bandyopadhyay, A., Bose, S., Ding, C., Luk, K.D.K., Cheung, K.M.C., Lu, W.W., 2007. Preparation and cellematerials interactions of plasma sprayed strontiumcontaining hydroxyapatite coating. Surf. Coat. Technol. 201, 4685e4693. Yang, F., Yang, D., Tu, J., Zheng, Q., Cai, L., Wang, L., 2011. Strontium enhances osteogenic differentiation of mesenchymal stem cells and in vivo bone formation by activating Wnt/ catenin signaling. Stem cells 29, 981e991. Yuan, B., Raucci, M.G., Fan, Y., Zhu, X., Yang, X., Zhang, X., Santin, M., Ambrosio, L., 2018. Injectable strontium-doped hydroxyapatiteintegrated with phosphoserinetetheredpoly(epsilon-lysine) dendrons for osteoporoticbone defect repair. J. Mater. Chem. B 6, 11.

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Zahra, N., Fayyaz, M., Iqbal, W., Irfan, M., Alam, S., 2014. A process for the development of strontium hydroxyapatite. In: IOP Conference Series: Materials Science and Engineering, p. 60. Zhang, C., Li, C., Huang, S., Hou, Z., Cheng, Z., Yang, P., Peng, C., Lin, J., 2010. Self-activated luminescent and mesoporous strontium hydroxyapatite nanorods for drug delivery. Biomaterials 31, 10. Zhang, W., Shen, Y., Pan, H., Lin, K., Liu, X., Darvell, B.W., Lu, W.W., Chang, J., Deng, L., Wang, D., Huang, W., 2011. Effects of strontium in modified biomaterials. Acta Biomater. 7, 800e808. Zhu, K., Yanagisawa, K., Shimanouchi, R., Onda, A., Kajiyoshi, K., 2006. Preferential occupancy of metal ions in the hydroxyapatite solid solutions synthesized by hydrothermal method. J. Eur. Ceram. Soc. 26, 509e513.

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Coating of hydroxyapatite and substituted apatite on dental and orthopedic implants

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Farasat Iqbal, Hira Fatima Interdisciplinary Research Centre in Biomedical Materials (IRCBM), COMSATS University Islamabad, Lahore Campus, Lahore, Pakistan

14.1

Introduction

Bioimplantation is the significant advancement of medical technology, which started developing as early as 1980s. Road accidents, trauma injuries, and cardiovascular, spinal, and ophthalmic disorders raised the demand for bioimplants. The bioimplant materials include mono- and polycrystalline materials, ceramics, polymerics, and glassy and composite materials. These biomaterials are assumed as the substitute of natural bone because of their comparable osteoinductive, osteoconductive, and mechanical reliability as that of the natural bone. The implanted tissue depends on the interaction type between implantation site and implanted material and is categorized as (1) biotoxic (2) nontoxic, biologically inert and inactive; (3) nontoxic and bioactive; and (4) nontoxic, biodegradable, which gets dissolved in body fluids and replaces the surrounding tissues (Prakasam et al., 2017). Implanted material is further classified as nondegradable and biodegradable, which includes stainless steel, cobalt-chromium, and titanium and its alloys as nondegradable while many ceramics and biopolymers, iron, magnesium, zinc, and their alloys fall under the category of biodegradables. Biocompatibility of implant material is considerably important as otherwise they cause postoperative inflammations, decline in mechanical strength, and various complications such as osteolysis. Biodegradable materials are expected to degrade gradually to assist in the healing process. Biodegradable alloys of magnesium (Chen et al., 2014) and iron (Schinhammer et al., 2010) are estimated to degrade with an appropriate metabolic degradation rate inside the human body. But the medical goal about these implants is to achieve and restore natural functions of the body as for mechanical load-bearing orthopedic implantation sites, high-strength metallic implants required having association with medical imaging artifacts and because of their high modulus, stress shielding, and resorption of unwanted bone occur (Nagels et al., 2003). Osseointegration and implant resorption are facilitated by bioresorbable polymers and composites, but these are limited only to soft tissue reforms and have quoted for incomplete degradation, prolonged inflammation, and implant breakage (Pereira et al., 2013). So, for endoprosthesis such as total replacement of hip and sockets for artificial

Handbook of Ionic Substituted Hydroxyapatites. https://doi.org/10.1016/B978-0-08-102834-6.00014-8 Copyright © 2020 Elsevier Ltd. All rights reserved.

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implantation of teeth, metallic implants are encountered because these implants have enough mechanical stability, but to undergo rebonding of bone, metals require some biocompatible coverings. Polyether ether ketone, comparatively new implant material, considered as a high strength polymer, is used primarily in spinal fusions and soft tissue reconstructions, with closely resembled stiffness and compatibility to bone (Kurtz and Devine, 2007). Because of its inert and hydrophobic surface, it restricts to attach with local bone as it suffers poor property of osteointegration (Arima and Iwata, 2007), which can be enhanced by surface modifications as by addition of bioactive coverings of titanium and hydroxyapatite (HA) to enhance cellular response (Han et al., 2010; Evans et al., 2015). Dental implantation requires high load-bearing metallic periodontal ligament (PDL) that usually consists of a titanium screw coated with HA, which provides bioactive PDL connection to titanium implants (Lee et al., 2017). In general, to make the surface of implant material osteoconductive and biocompatible, surface deposition required great chemical similarity to the inorganic constituents of bone and teeth, i.e., calcium orthophosphate (CaPO4), and HA is required. These surface depositions are manufactured on industrial scale by plasma spraying technique, with thickness ranging from w30 to 200 mm (Dorozhkin, 2012a). Medical results obtained for these coatings are not as auspicious as estimated because of quite thick coating layer with high tendency to get cracks and delaminate out (Fielding et al., 2012). To overcome this, various modifications in coating methods are employed including plasma-assisted technique, which deposit thin, more adhesive, and crack-free coating on substrate material (Dorozhkin, 2015; Surmenev, 2012; Surmenev et al., 2014). Magnetron sputtering technique is used to deposit highly adhesive layer on various substrates, while use of cosputtering enhances the adhesiveness further, resulting in the formation of quite fine films (Lu et al., 2008). Films deposited by pulsed laser technique depend greatly on annealing temperature and are crystalline nanostructures. Their thickness, microhardness, and adhesion also depend on temperature (Rau et al., 2015; Graziani et al., 2018). Deposition methods employed in lab include vapor deposition techniques, which are further categorized into two main groups: physical vapor deposition and chemical vapor deposition (as PVD and CVD, respectively). The advantage of all sputtering technique is deposition of thin coatings having compact structure and strong adhesive forces with itself as well as with substrate surface. Investigations revealed that the deposition taking place by ion beam assisted deposition has gradually decreased size of grains and crystalline nature toward the surface, leading toward the nanosized grains and deposition of amorphous layer on the surface (Wang et al., 2008a). Another type of pulsed vapor deposition that uses plasma sputtering technique is electron cyclotron resonance, which has very low productivity rate and is practicable for some definite applications, including deposition of thin epitaxial HA coatings and layers and films with specific crystal orientation (Akazawa and Ueno, 2014). Certain wet deposition techniques are employed at moderate temperatures (Nijhuis et al., 2010), from aqueous and nonaqueous suspensions or solutions, which also depend on the pH of solution to deposit various CaPO4 precipitates. These deposition

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techniques are generally based on the heterogeneous nucleation and kinetics of various parameters. Electrophoretic deposition is a broad term that includes electrocoating, ecoating, cathodic electrodeposition, and electrophoretic coating or electrophoretic painting. Electrodeposition technique, considered as the earliest known technique proposed in the United States, is used to deposit CaPO4 coatings. While by biomimetic deposition, coating of bioactive, bonelike apatite material is formed by immersing the substrate in various simulating solutions, i.e., HBSS, PBS, or simulated body fluid (SBF). Dip coating is a simple and common technique, which is usually used to deposit low adherent complicated mixture of many compounds (Dorozhkin, 2015). For in vivo applications, calcium orthophosphate (CaPO4) seems very friendly because it has prodigious similarity with inorganic constituents of mammalian’s teeth and bones (Dorozhkin, 2011, 2012b) also another phase of calcium phosphate, Monetite (CaHPO4) shows full viability when exposed to host cells (Ali et al., 2019). Calcium phosphate and some salts such as carbonated and calcium-deficient apatite with Mg, Na, K, chlorides, and fluorides make a biological apatite. Synthetic HA is chemically and crystallographically similar with the natural one (De Groot et al., 1998). HA surface shows selective chemical reactivity with adjacent tissues so it is also considered osteoconductive, and as a result of this nature, strong interaction of bone and implant is formed. Furthermore, studies also revealed that HA induced osteogenic differentiation of osteoblasts (Wang, 2004), which is improved by coating nanosized crystals of HA. Moreover, synthetic HA tolerates the ionic substituents in its lattice, i.e., divalent and trivalent cations, which greatly affect the size, physicochemical properties, and crystalline nature of HA. It is demonstrated by previous researches that synthetic HA can be incorporated by a variety of ionic substituents to produce bonelike mineral composition. These lattice substituents employ prominent effects on the physical characteristics including lattice parameters, morphology, solubility, and crystal structure, etc., in relevance to the nonsubstituted HA (Ratnayake et al., 2017). Orthopedic implant coatings of HA with Zn, Mg, Sr, Si, carbonate, and fluoride substitutions are mentioned in research, and some of these are also available commercially for augmentation and repair of bone (Shepherd et al., 2012; Worth et al., 2005). These coatings improve the contact and fixation of induced implant material with bone and enable the association of implant material and surrounding bone. Therefore, the clinical performance of orthopedic hip system and tooth implants is increased. Aim of these coatings is their close resemblance with the natural bone. Doping of HA by various appropriate ions and the effect of their concentrations with respect to biological relevance have been investigated in various investigations. Another important feature of the coatings is their antibacterial character, which can be obtained by doping of HA with silver or copper by magnetron sputtering (Ciuca et al., 2016; Ivanova et al., 2016), pulsed laser deposition (PLD) (Kotoka et al., 2016), and matrix-assisted pulsed laser evaporation (Jankovi
c et al., 2015). Degradation rate, surface smoothness, and adherent nature of coating increase the effective potential, while in vivo studies of these biocompatible coatings on various bones, as they have a little variation in their composition, improve the clinical significance.

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Hydroxyapatite coatings for dental and orthopedic implants

Regeneration and healing itself is the remarkable property of bone tissues. Large and complex bone defects are still challenging in medical science and now trying to overcome by implanting metallic (cobalt-chromium alloy, titaniume6Ale4V, stainless steel 316L, magnesium, and zirconium) plates and screws, which mechanically support the long-defected bone. HA coating is used to increase the mechanical strength (Hadidi et al., 2017; Wen et al., 2017) such as coating adhesion and load-bearing ability (Liu et al., 2011; Roy et al., 2008) of implant material. Requirement of an implant material depends on the fracture site and application. After healing, the removal of plates and screws from body does not require a lot of osteointegration, while rods and screws used in arthroplasty of hip and knee and in surgeries of spinal fusions require high osteointegration, so they stay and support the body throughout the life as they bind strongly to the bones allowing minimum or no movement while providing maximum load-bearing function (Ramazanoglu and Oshida, 2011). Depending on patient’s age and fitness, implant must develop strong adhesion with host tissues. The adhesion of coating layer with implant is very critical as during or after implantation, if the coated layer scratched out from metal, then it affects adversely to the surrounding (Mohseni et al., 2014). These implants may face complications such as infections, implant loosening, and improper loading, leading us toward the modifications of implant to increase bone growth by osteointegration and cell adhesion. Bioadhesive coatings such as calcium phosphate (CaP), interestingly, HA coatings, are used to increase rapid and fast osteointegration (Bryington et al., 2013) as they increase the bone contact and growth. They also decrease the exposure of implant metal to corrosive environment. These coatings must also be compatible with packing techniques, sterilization techniques, size, and type of implantation site. Generating biocompatibility and chemical similarity of mechanically load-bearing material with host tissue and bone is the main reason of coating. Titanium itself as well as its alloys is the most commonly used material in orthopedics and dental because of its nontoxicity, great tensile strength, low density, and excellent resistivity against corrosion. HA, a bioactive coating material bonded with implant material as well as with the host tissue, enhances the differentiation and attachment of cell and bone formation. On implantation, it forms a layer of carbonated apatite having structure and composition alike to the mineral composition of osseous tissue by the process of precipitation and dissolution, while fibrils of collagen incorporated in this carbonated apatite enhance the biocompatibility of implant. In total hip replacement, HA acts as a fixer for components of femoral bone. Partially coated Omnifit stem also shows excellent results; therefore, cemented fixation is considered unreliable and less efficient. In fact, CaP actively contributes in the remodeling of bone, and HA coating is substituted by new bone formation without generating any layer of fibrous tissues. HA particles do not cause any inflammatory reaction and do not act as foreign body particles (Epinette and Manley, 2008) (Fig. 14.1).

Coating of hydroxyapatite and substituted apatite on dental and orthopedic implants

331

Acetabulum (hip socket)

Osteointegration

Biocompatibility Femur Coating adhesion

Figure 14.1 Functions of hydroxyapatite (HA) coatings on total hip replacement implants.

As HA is chemically similar to the inorganic part of bone, its coating enhances bone growth at the interface of implant and host body tissues, while interfacial stability is achieved by HA bioactive layer on the surface of implant material, which provides strong interface compared with the bone strength and eliminates the requirement of cemented fixation. In the presence of aqueous environment as present in body fluids, implanted metal releases ions and generates hydrogen bubble, which separates the tissues and may block the blood circulation, also hindering the binding of cells to implant material and inducing loss of periprosthetic bone leading toward the implant failure. In this situation, HA layer acts as buffer and stops releasing such harmful metal ions (Grandfield et al., 2011). Hip prosthesis coated with HA is very suitable for long-lasting implant stability as it improves the bone and implant contact (Geesink, 2011) and induces fast osteointegration of HA-coated metallic implants (Chambers et al., 2007). According to literature, because of chemical dissolution of HA coating or osteoclastic action, 78% of it is replaced by the regenerated bone, which proves the stable fixation of HA-coated implants and prevents osteolytic reaction (Herrera et al., 2015). Dental implants do not have same biomechanical strength as orthopedic implants because implanted teeth left unloaded until it anchored, while orthopedic implants allow early weight bearing (Martinez-Carranza et al., 2014). Therefore, dental implant material is required to be bioinert with fast osteointegration to achieve long-term functionality. However, they are designed to promote strong interaction of implanted metal to host bone and to get primary mechanical stability through osteointegration (Javed et al., 2013). For osteointegration and bone formation, they should be compatible with hard and with soft tissue to inhibit formation of biofilm and for epithelium gingival adhesion. HA coatings by plasma spray method onto metallic core unite the implant teeth with the host bone (Triplett et al., 2003). According to research, rough surface is more adhering to osteoblastic cells than smooth ones (Elias, 2014),

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and they recognize specific molecules including amino acid, collagen, antimicrobial agents, pharmaceutical materials such as biophosphonates (Krishna Kumar et al., 2011; Palmquist et al., 2010), and some functional groups. Fluoride ions are also coated on dental implants as they enhance the calcification of bones (Elias, 2014) and deposition of minerals and proteins. HA as inorganic phase used to enhance osteoconductive nature of implant and impart direct contact of implant to host bone. HA coating on titanium implants may be etched out and may release coating particles and become the reason of implant failure (Le Guéhennec et al., 2007; Gaviria et al., 2014). So, mineralization and cell adhesion are increased by chemical treatment or molecular grafting of coating surface, and for adhesion of protein, hydrophilicity of surface has been modified by various treatments. Several studies revealed that implant with hydrophilic surface show fast implantehost integration. Dental implants are more susceptible to infection and have short life than natural ones, and to overcome this issue living PDL, connections are engineered for bioimplants that are made up of titanium screw coated with HA and wrapped in the sheets of human immortalized periodontal cells (Lee et al., 2017). To enhance osteointegration, the surface of implant is modified by incorporation of some antiresorptive, osteogenic medicines such as bisphosphonate, which are significant to medical cases. HA coating with gentamycin on implant may act as antibiotic and prophylactic agent in dental implantation, while treatment of implant with tetracyclineHCl causes detoxification and decontamination of implanted surface.

14.3

Processing of hydroxyapatite and substituted apatite coatings

Calcium phosphate, a class of bioceramics that includes HA, an inorganic compound also present in bones and teeth, has close structural and phase resemblance with the natural bone (Dorozhkin, 2009; Meurice et al., 2012). As its coatings apply on implanted material, it provides enough calcium and phosphate ions at initial implantation stage and makes the implant material biocompatible (Abdel-Aal et al., 2008). A number of techniques are used to prepare HA coatings (Dorozhkin, 2015), which include plasma spray technique (Singh et al., 2014), thermal spray technique using high-velocity oxygen fuel (Ros¸u et al., 2011), radio frequency magnetron sputtering, sputtered coating (Davison et al., 2015), biomimetic coating (Wang et al., 2008b), PLD (Arias et al., 2003), ion beam deposition (Luo et al., 1999), dip coating, frit enameling (Jansen and Leon, 2009), hot isostatic pressing, organometallic CVD, solegel derivation, electrophoretic deposition (Wei et al., 2005; Eliaz et al., 2005), chemical deposition (Tas, 2000), and electrodeposition (Lakstein et al., 2009; Eliaz and Sridhar, 2008; Eliaz et al., 2009; Metoki et al., 2014, 2016; Geuli et al., 2016; Thomas et al., 2017; Eliaz and Eliyahu, 2007). From these techniques, some are described in Table 14.1.

Coating of hydroxyapatite and substituted apatite on dental and orthopedic implants

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Table 14.1 Calcium phosphate deposition techniques (Harun et al., 2018a; Mohseni et al., 2014; Wang et al., 2014; Rojaee et al., 2014; Bolelli et al., 2015; Dorozhkin, 2015; Bose et al., 2015). Technique

Thickness

Advantages

Disadvantages

Thermal spray

30e200 mm

High rate of deposition, low cost

Line of sight technique, high temperatures induce decomposition and amorphous nature, rapid cooling causes cracks to appear, low porosity, coating spalling and interface separation between the coating, and the substrate may occurs

Plasma spray

30e300 mm

Fast deposition, relatively low cost, rapid bone healing, minimum coating degradation risk, uniform coating, simple and flexible technique

High temperature up to 1200 C, phase transformation due to high temperature, less crystalline coatings, increased residual stress, line of sight technique, impossible to incorporate biological agents, rapid cooling produce cracks

Electrostatic spray deposition

10 nme30 mm

Relatively low cost, easy setup, ambient conditions, extensive choice of substrate and precursors

Line of sight technique, problems to coat large surfaces, low flow rates, hightemperature requirements to decompose the precursor solvents and salts

High-velocity oxygen fuel spraying

30e200 mm

High deposition rates, enhanced biocompatibility, wear and corrosion resistance

Line of sight technique, partial decomposition and less crystallinity due to high temperature, difficult to incorporate biological agents, rapid cooling produces cracks Continued

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Table 14.1 Continued Technique

Thickness

Advantages

Disadvantages

High-velocity suspension flame spraying (HVSFS)