Bioactive Glasses: Materials, Properties and Applications (Woodhead Publishing in Materials)

Bioactive glasses © Woodhead Publishing Limited, 2011 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1...

Author: H Ylanen

80 downloads 1504 Views 47MB Size Report

This content was uploaded by our users and we assume good faith they have the permission to share this book. If you own the copyright to this book and it is wrongfully on our website, we offer a simple DMCA procedure to remove your content from our site. Start by pressing the button below!
Report copyright / DMCA form

DOWNLOAD PDF

Bioactive glasses

© Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Related titles: Biomedical composites (ISBN 978-1-84569-436-4) Biomedical composites provides a thorough review of the current status, recent progress and future trends in composites for biomedical applications. Part I discusses the fundamentals with chapters on natural composites, design and fabrication, and general hard and soft tissue applications. Part II reviews specific applications. Chapters discuss bone repair, coatings for implants, spinal implants, injectable composites and composites for tissue engineered scaffolds. Part III covers the biocompatibility, mechanical behaviour and failure of biocomposites with such topics as cellular response, testing of biocomposites and tribology. Part IV analyses the future for biocomposites with chapters on nanostructured biocomposites, developing biocomposites as scaffolds and biocomposites in tissue engineering and regenerative medicine. Bioceramics and their clinical applications

(ISBN 978-1-84569-204-9) Bioceramics are potentially suitable for a wide range of important applications within the medical device industry. Bioceramics and their clinical applications provides an authoritative review of this highly active area of research, written by leading academics from around the world. Chapters in the first section of the book discuss issues of significance to a range of bioceramics, such as their structure, mechanical properties and biological interactions. The second part reviews the fabrication, microstructure and properties of specific bioceramics and glasses, concentrating on the most promising materials. The final group of chapters reviews the clinical applications of bioceramics. Bone repair biomaterials

(ISBN 978-1-84569-385-5) Bone repair is a fundamental part of the rapidly expanding medical care sector and has benefited from many recent technological developments. This unique book provides a comprehensive review of the materials science, engineering principles and recent advances in this important area. The first part of the book reviews the fundamentals of bone repair and regeneration. Further chapters discuss the science and properties of biomaterials used for bone repair such as metals and biocomposites. The final set of chapters analyses device considerations such as implant lifetime, failure, applications and ethics of bone repair biomaterials. Details of these and other Woodhead Publishing materials books can be obtained by: • •

visiting our web site at www.woodheadpublishing.com contacting Customer Services (e-mail: [email protected]; fax: +44 (0) 1223 832819; tel.: +44 (0) 1223 499140 ext. 130; address: Woodhead Publishing Limited, 80 High Street, Sawston, Cambridge CB22 3HJ, UK)

If you would like to receive information on forthcoming titles, please send your address details to: Francis Dodds (address, tel. and fax as above; e-mail: francis. [email protected]). Please confirm which subject areas you are interested in.


Bioactive glasses Materials, properties and applications Edited by Heimo O. Ylänen


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Published by Woodhead Publishing Limited, 80 High Street, Sawston, Cambridge CB22 3HJ, UK www.woodheadpublishing.com Woodhead Publishing, 1518 Walnut Street, Suite 1100, Philadelphia, PA 19102-3406, USA Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India www.woodheadpublishingindia.com First published 2011, Woodhead Publishing Limited © Woodhead Publishing Limited, 2011 The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Control Number: 2011932770 ISBN 978-1-84569-768-6 (print) ISBN 978-0-85709-331-8 (online) The publisher’s policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp that is processed using acid-free and elemental chlorine-free practices. Furthermore, the publisher ensures that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by RefineCatch Limited, Bungay, Suffolk Printed by TJI Digital, Padstow, Cornwall, UK


Contents

Contributor contact details Introduction Part I

1

ix xiii

Materials and mechanical properties of bioactive glass

1

Melt-derived bioactive glasses

3

L. HUPA, Åbo Akademi University, Finland

1.1 1.2 1.3 1.4 1.5

Introduction Manufacture and physical properties Chemical properties and bioactivity Future trends References

3 6 13 23 23

2

Surface modification of bioactive glasses

29

J. CHANG and Y. L. ZHOU, Chinese Academy of Sciences, China and Y. ZHOU, Shanghai Jiao Tong University, China

2.1 2.2

2.5 2.6

Introduction Surface modification of bioactive glasses to improve bioactivity Surface modification of bioactive glasses using organic molecules to improve dispersivity Surface modification of bioinert materials using bioactive glasses Conclusions and future trends References

39 43 44

3

Cell interaction with bioactive glasses and ceramics

53

2.3 2.4

29 30 35

R. P. K. PENTTINEN, University of Turku, Finland

3.1 3.2

Introduction Biology of bioactive glasses

53 54 v


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

vi

Contents

3.3 3.4 3.5 3.6 3.7

Reaction of cells with glasses and related ceramics Effect of silica on bone formation Future trends References Appendix: list of abbreviations

57 66 71 72 84

4

Regulatory aspects of bioactive glass

85

S. LINDGREN, T. PÄNKÄLÄINEN, J. LUCCHESI and F. OLLILA, BonAlive Biomaterials Ltd, Finland

4.1 4.2 4.3 4.4 4.5

Introduction General requirements Indication areas Market approval process in some geographical areas References

Part II Applications of bioactive glass 5

Bioactive glass and glass-ceramic scaffolds for bone tissue engineering

85 86 94 98 103 105

107

X. CHATZISTAVROU, University of Erlangen-Nuremberg, Germany, P. NEWBY, Imperial College London, UK and A. R. BOCCACCINI, University of Erlangen-Nuremberg, Germany and Imperial College London, UK

5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8

Introduction Requirements for bone tissue scaffolds Bioactive glasses and glass-ceramics in bone tissue engineering Bioactive glass-based scaffolds: fabrication technologies Scaffolds from boron-containing bioactive glass Polymer-coated composite scaffolds Conclusions References

107 108 109 110 114 119 123 124

6

Nanoscaled bioactive glass particles and nanofibres

129

M. EROL, Istanbul Technical University, Turkey and A. R. BOCCACCINI, University of Erlangen-Nuremberg and Imperial College London, UK

6.1 6.2 6.3 6.4 6.5 6.6 6.7

Introduction Characteristics of nanoscale bioactive glasses Fabrication of bioactive glass nanoparticles and nanofibres Applications of nanoscale bioactive glasses Conclusions Acknowledgement References


129 131 132 138 152 152 152

Contents

7

Bioactive glass containing composites for bone and musculoskeletal tissue engineering scaffolds

vii

162

S. VERRIER, AO Research Institute Davos, Switzerland, J. E. GOUGH, University of Manchester, UK and A. R. BOCCACCINI, University of Erlangen-Nuremberg, Germany and Imperial College London, UK

7.1 7.2 7.3 7.4 7.5 7.6 7.7

Introduction Composite materials approach to tissue engineering scaffolds In vitro and in vivo evaluation Discussion Conclusions and future trends References Appendix: list of abbreviations

162 165 169 179 180 181 188

8

Use of bioactive glasses as bone substitutes in orthopaedics and traumatology

189

J. HEIKKILÄ, Sports Clinic and Hospital Mehiläinen Turku, Finland

8.1 8.2 8.3 8.4 8.5 8.6 8.7 8.8 8.9 8.10 8.11 8.12 8.13 8.14 8.15 9

Introduction Glass surface reactions The bonding of bioactive glass and bone formation The biocompatibility of bioactive glasses The strength of bioactive glass Bone formation Clinical use for benign bone tumors Bioactive glass and infection Bioactive glass in cancellous bone and metaphyseal fractures Bioactive glass in diaphyseal bone fractures Bioactive glass in spinal surgery Arthroplasty Summary of applications in orthopaedics and traumatology Future trends References Bioactive glass S53P4 as a bone graft substitute in the treatment of osteomyelitis

189 192 194 195 196 197 200 201 201 202 202 203 203 204 204 209

N. C. LINDFORS, Helsinki University Central Hospital, Finland

9.1 9.2 9.3 9.4 9.5

Introduction Bone grafts in the treatment of osteomyelitis Antibacterial properties of bioactive glass S53P4 Vascularization-promoting properties of bioactive glasses Bioactive glass S53P4 in the treatment of osteomyelitis: a multicentre study


209 210 211 211 212

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

viii

Contents

9.6 9.7

Conclusions References

213 215

10

Bioactive glass for maxillofacial and dental repair

217

M. J. PELTOLA and K. M. J. AITASALO, Turku University Hospital, Finland

10.1 10.2 10.3 10.4 10.5 10.6 11

Introduction Current materials and requirements in maxillofacial reconstruction Properties of bioactive glass Clinical applications of bioactive glass in maxillofacial reconstruction Clinical applications of bioactive glass in dentistry References Bioactive glass and biodegradable polymer composites

217 217 219 220 223 224 227

T. NIEMELÄ and M. KELLOMÄKI, Tampere University of Technology, Finland

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8

Introduction Biodegradable polymers Manufacturing of the composites Bioactive glass particle composites Bioactive glass fiber composites Coatings Future trends References

227 228 229 231 235 238 241 241

12

Bioactive glasses for wound healing

246

M. SHAH MOHAMMADI, C. STÄHLI and S. N. NAZHAT, McGill University, Canada

12.1 12.2 12.3 12.4 12.5 12.6

Introduction Silicate-based versus phosphate-based bioactive glasses Antibacterial properties of bioactive glasses Stimulation of angiogenesis Conclusions References

246 246 250 254 259 260

Index

267


Contributor contact details

(* = main contact)

Editor and Introduction H. O. Ylänen Tampere University of Technology Department of Biomedical Engineering Hermiankatu 12A PL 692 33101 Tampere Finland E-mail: [email protected]

Chapter 1 L. Hupa Åbo Akademi University Biskopsgatan 8 20500 Turku Finland E-mail: [email protected]

Chapter 2 J. Chang* and Y. L. Zhou State Key Laboratory of High Performance Ceramics and Superfine Microstructure Shanghai Institute of Ceramics Chinese Academy of Sciences 1295 Dingxi Road Shanghai 200050 China

Y. Zhou Med-X Research Institute Shanghai Jiao Tong University 1954 Hua Shan Road Shanghai 200030 China

Chapter 3 R. P. K. Penttinen Department of Medical Biochemistry and Genetics University of Turku Kiinamyllynkatu 10 20520 Turku Finland E-mail: [email protected]

Chapter 4 S. Lindgren*, T. Pänkäläinen, J. Lucchesi and F. Ollila BonAlive Biomaterials Ltd Biolinja 12 20750 Turku Finland E-mail: [email protected]

E-mail: [email protected]

ix © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

x 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X


Chapter 5

Chapter 7

X. Chatzistavrou and A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany

S. Verrier Musculoskeletal Regeneration AO Research Institute Davos 7270 Davos Switzerland

E-mail: aldo.boccaccini@ww. uni-erlangen.de

P. Newby Department of Materials Imperial College London Prince Consort Road London SW7 2BP UK

Chapter 6 M. Erol Department of Chemical Engineering Istanbul Technical University Maslak 34469 Istanbul Turkey E-mail: [email protected]

A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany

J. E. Gough School of Materials University of Manchester Manchester M1 7HS UK A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany E-mail: aldo.boccaccini@ww. uni-erlangen.de

Chapter 8 J. Heikkilä Assistant Professor in Orthopaedics and Traumatology Sports Clinic and Hospital Mehiläinen Turku Kauppiaskatu 8 20100 Turku Finland E-mail: [email protected] [email protected]

E-mail: aldo.boccaccini@ww. uni-erlangen.de



Chapter 9

Chapter 11

N. C. Lindfors Department of Orthopaedic and Hand Surgery Helsinki University Central Hospital Töölö Hospital Topeliuksenkatu 5 00260 Helsinki Finland

T. Niemelä and M. Kellomäki* Department of Biomedical Engineering Tampere University of Technology PO Box 692 33101 Tampere Finland


Chapter 12

xi


M. J. Peltola* and K. M. J. Aitasalo Department of Otorhinolaryngology – Head and Neck Surgery Turku University Hospital 20521 Turku Finland

M. Shah Mohammadi, C. Stähli and S. N. Nazhat* Department of Mining and Materials Engineering McGill University 3610 University Street Montreal Quebec H3A 2B2 Canada



Chapter 10


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Introduction

Today, millions of prostheses, implants and bone grafts are needed to maintain the quality of life of the aging population. The increased human lifespan alone has created enough problems in this regard, even if we discount the need for repairing or replacing body parts damaged by an individual’s own actions. As a result, material scientists have for decades faced the challenge of developing advanced biomaterials to repair the defects of the human body or to replace damaged parts. A significant advance in the search for better biomaterials was provided by the introduction of two new synthetic biomaterials during the early 1970s. The materials were developed independently and almost simultaneously by several groups of materials scientists. The new synthetic biomaterials were able to bond to host tissue through chemical processes occurring on the materials’ surface, and the materials were termed bioactive ceramics. A comprehensive review of the state of the art in bioceramics from basic science to clinical applications is presented in The Handbook of Bioceramics and Their Applications, edited by Professor Tadashi Kokubo (Woodhead Publishing Limited, 2008). The current book focuses on a special subgroup of bioactive ceramics, namely bioactive glasses. Systematic research into bioactive glasses was started by Professor Larry Hench in 1969, when he introduced the concept of a strong bonding between bone and synthetic material brought about by chemical reactions occurring on a glass surface. The innovation concerned the chemical reactivity of the surface of a silica-based material that had the amorphous structure of silicate glass. Hench introduced the material, a bioactive glass, in the early 1970s. Bioactive glasses bond firmly to bone through chemical reactions and can ultimately be replaced by bone: these properties make them extremely promising as a material for medical applications. Most importantly, the constituents in bioactive glass are physiological chemicals found in the body, typically silicon, sodium, potassium, magnesium, oxygen, calcium and phosphorus. According to several studies, during the bonding and formation of bone, the concentration of the chemicals never rises to levels that could disturb the adjacent tissues. The use of bioactive glass as an implant material or in manufacturing medical devices is limited, however, by the mechanical properties of glass. Glass is brittle xiii © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

xiv 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Introduction

and cannot therefore be used in positions where load-bearing properties are required. Glass can be cast to plates, rods or simple devices; alternatively it can be formed by sawing or grinding cast rods to rigid medical devices. It can also be used as a filler material in the form of particulate. In order to widen the manufacturing of bioactive glass products to technically more demanding clinical applications, new bioactive glasses were developed. Today, new bioactive glasses can be tailor-made to a variety of clinical applications, to different shapes, fibers, microspheres and to show optimal bioactivity in different physical conditions found in the human body. Novel applications of bioactive glasses require not only tight control of the bioactivity but also a thorough knowledge of the influence of the composition of glass on its manufacture and its formation into different products. The tailoring of glasses to certain applications is thus based on understanding and mastering a wide range of properties important for both medical applications and for the manufacture of glass. The manufacturing of conventional melt-derived bioactive glasses demands extremely high temperatures and careful annealing procedures. To avoid this, new technologies were developed for manufacturing new types of bioactive glasses such as those derived from sol-gel, spun bioactive glass fibers and bioactive glass nanoparticles. Today, the use of bioactive glass as a component of biomaterial composites is one of several interesting options in the development of a variety of clinical applications of this material. In the current book, the development of different bioactive glasses is reviewed by globally distinguished experts and scientists. The book starts with an introduction to different types of bioactive glasses, the influence of surface modification on the properties of these special glasses, and the process of bonding to different types of host tissues. It also discusses cell interactions at the interface of bioactive glasses and tissues. When marketing any products containing bioactive glass, regulations imposed by authorities are of crucial importance. A chapter addressing this issue is written by a company specializing in products made of bioactive glass. Having been developed over several decades, bioactive glass products are today in clinical use across the world. This means that it is now possible to discuss experiences of the clinical applications of these products. The current book also covers this interesting area of bioactive glasses. Earlier works have been published that have covered the area of bioceramics and even their clinical applications. However, to my knowledge, the current book is the first one to discuss solely bioactive glasses and their application. The book should become a standard textbook in both the fields of materials sciences and medical sciences. I hope that by publishing this book we can encourage an interest in the development of bioactive glasses in students and researchers the world over. Heimo O. Ylänen


1 Melt-derived bioactive glasses L. HUPA, Åbo Akademi University, Finland

Abstract: This chapter discusses the properties of melt-derived bioactive glasses from the material technology point of view. The non-crystalline structure of bioactive glasses offers the possibility of adjusting their physical and chemical properties by altering their oxide composition within certain limits. Thus, understanding the relationships between the oxide composition and the relevant properties is essential when the glass composition is tailored for novel clinical applications. The focus is to summarize some published data on the in vitro and in vivo bioactivity. The restrictions put on glass composition by the manufacturing process are also discussed. The aim is to provide some fundamental tools for further studies and the development of melt-derived glasses to desired product forms for various clinical applications. Key words: melt-derived bioactive glasses, viscosity, crystallization, in vitro bioactivity, in vivo bioactivity.

1.1

Introduction

A new era in the development of materials for use in medicine began in the 1970s, when Professor Larry Hench discovered glasses capable of forming interfacial bonding with bone (Hench and Paschall, 1973). The compositions showing this special property were called bioactive glasses. In developing the glasses Professor Hench’s leading idea was to find a material that, rather than forming an interfacial layer of scar tissue, would instead form a living bond with the host tissues (Hench 2006). The hypothesis behind the glass development was simple but ingenious: as bone contains hydroxyapatite, HA, the implant material should be able to form an HA layer on the surface in biological solutions. Such a material would not be rejected by the body but be bonded directly with the tissue. Hench and co-workers tested whether phosphate containing soda-lime silicate glasses could fulfil the criterion of tissue bonding. Glasses within this system would then contain two important components of the hydroxyapatite (Ca5(PO4)3OH), namely Ca2+ and PO43− ions. The two other cations, Na2+ and Si4+, are also common components of the human body. Implants of Glass 45S5, one of the first compositions tested, were found to bond to rat femur. This glass, known also as Bioglass®, is still one of the most bioactive glasses known. The selection of the composition was ideal; the low silica content makes the glass easy to melt but also gives it much lower chemical durability than commercial soda-lime glasses in aqueous solutions. A low chemical durability and the ability of the composition to form a dual layer of silica and amorphous calcium phosphates on the surface are key features of 3 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

4 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

bioactive glasses (Hench, 1991, 1992). The subsequent reactions in the surface layers lead to bonding between the glass and surrounding tissue. Ideally, the glass will dissolve and be replaced by new tissue in time. The kinds of materials that were capable of bonding directly to living tissue were called bioactive materials. Bioactive materials are defined as (1) materials that have been designed to induce special biological activity; or (2) biomaterials that are designed to elicit or modulate biological activity (Williams 1999). Bioactive glasses are capable of forming a bond with both hard and soft tissue in vivo or in vitro environments by developing a surface layer of hydroxycarbonate apatite by release of ionic species from the bulk material (Williams 1999). Bioactive glass research deals in large part with developing a fundamental understanding of the dissolution and surface reactions of the glass and the tissue response to the dissolving material. Over the years, the interactions of several glass compositions with biological solutions have been studied. However, the research has been concentrated mainly on compositions close to that of the bioactive glass 45S5. Li et al. (1991) reported that sol-gel-derived glasses within the system Na2O-CaO-SiO2 show bioactivity within a much larger composition range than melt-derived glasses. Since then, sol-gel derived glasses have been studied intensively. Jones (2008, 2009) has summarized the use of sol-gel glasses as materials for nanostructured bioactive scaffolds. This chapter deals with the properties of melt-derived bioactive glasses. The standpoint is more in glass science and technology than in biological sciences. The goal is to summarize various criteria to be considered when developing glass compositions for new products to various clinical applications. Controlled bioactivity is the basis for development of glass compositions. However, the tissue-engineering approach to manufacture porous thin-walled scaffold structures of glasses or to use glasses as constituents in composites calls for a better understanding of the overall properties of the glasses. Generally, glasses have good chemical durability. However, one of the most important properties of bioactive glasses is their controlled reactivity in body solutions. The reactivity of glasses in aqueous solutions is strongly dependent on the glass composition and thus one of the key factors for choice of composition. Chemical durability as well as several other glass properties can be adjusted smoothly by the composition within a certain range. Further, when using bioactive glasses in clinical applications, other properties than chemical durability should be considered. Mechanical strength and especially the ability to sustain a certain mechanical impact and loading are important properties during surgery. Basically, the strength of glasses is high, but due to their brittle nature they cannot be used in load-bearing applications. As mechanical properties are basically more dependent on the surface condition of the glass than on the glass composition, they are not essential for the composition choice. Today the medical uses of bioactive glasses are based mainly on crushed fractions. The intense research on bioactive and biodegradable glasses as



5

components of composites is likely to give birth to novel applications. For example, when using glass fibres in composites together with organic polymers the inherent mechanical strength of glasses is utilized in order to reinforce the composite structure. The mutual reactions between bioactive glasses and various biodegradable polymers are not well established. When more information is available the reactivity of glasses with other materials used in medicine should also be taken into account when choosing the composition. The specific criteria of reactivity in body solutions limit the composition range for the glasses. Within this range, the glass composition should be chosen so that it can be melted and formed into specific shapes with available methods. One of the most important characteristics for glass manufacture is the viscosity– temperature relationship, as it defines the methods that can be utilized in glass forming. Also the liquidus temperature should be considered; all practical melt forming operations have to be carried out at higher temperatures than the liquidus. Both viscosity and liquidus depend on glass composition. For traditional glasses the crystallization characteristics are not usually critical. However, the specific composition range of glasses showing bioactivity brings a risk of rapid crystal growth during glass forming. Another example is the manufacture of specific porous structures via sintering of glass particulates. In the manufacture of glassy structures, crystallization during sintering is an undesired phenomenon. On the other hand, controlled crystallization can be utilized for achieving bioactive glassceramics with specific properties. In both cases, the processing parameters are mastered via a good knowledge of the crystallization characteristics. Although crystallization in thermal treating of glasses is partly a kinetic phenomenon, it can be controlled by the choice of the glass composition. Thermal expansion of glasses should be considered when the glass is applied as a bioactive coating, for example on a metal prostheses. The glass composition should be adjusted to give a compatible adhesion with the metal for achieving a good adherence without any chipping or crawling of the coating on the metal. The choice of the glass composition for a specific application should be based on a firm knowledge on the influence of all major components on the most relevant properties of the glass with regard to both the final use and the manufacture of the product. Despite extensive research during the past 40 years, only a few glass compositions have been accepted for clinical use. The two US Food and Drug Administration FDA approved melt-derived compositions 45S5 (Hench and Paschall, 1973) and S53P4 (Andersson et al., 1990) consist of four oxides, SiO2, Na2O, CaO and P2O5. In general, a great number of elements can be dissolved in glasses. The effect of Al2O3, B2O3, Fe2O3, MgO, SrO, BaO, ZnO, Li2O, K2O, CaF2 and TiO2 on the in vitro or in vivo properties of certain compositions of bioactive glasses has been reported (Andersson et al., 1990; Vrouwenvelder et al., 1994; Brink et al., 1997; Haimi et al., 2009; Lusvardi et al., 2009; Zhang et al., 2009; Gentleman et al., 2010; Watts et al., 2010). However, the effect of the composition on the properties of bioactive and biodegradable glasses is not fully understood.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

6

Bioactive glasses

1.2

Manufacture and physical properties

1.2.1 Melting and forming Melt-derived bioactive glasses are melted and formed with methods similar to traditional soda-lime glasses. However, the requirements at the processing must meet the standards for materials used in medical applications. The batches are mixed of high purity analytical- and reagent-grade raw chemicals and thus the content of trace impurities in the glasses is low. Bioactive glasses are produced by melting batch components at an elevated temperature, typically 1350 to 1450°C, in electrically heated furnaces. The glasses are melted in platinum crucibles to avoid any contamination from oxide crucibles. Usually, no fining agents are added to the batches. The low viscosity of typical bioactive glass compositions at the melting temperature aids in eliminating gaseous inclusions from the melt. Melting times of small batches for laboratory testing varying from 1 to 24 hours have been employed. The glasses are often melted twice in order to increase homogeneity. Volatilization of components with high vapor pressures at high temperatures should also be taken into account. In bioactive glasses alkalis, boron, phosphorus and fluorides may vaporize. The glasses can be melted in covered crucibles to minimize losses. The vaporization in a certain process can also been taken into account by adjusting the batch composition. Forming and shaping procedures vary depending on the product type; casting into monoliths and drawing into rods or fibres are the main forming processes for bioactive glasses. After forming, the glass is annealed at a temperature corresponding to the viscosity 1013 dPa·s (1013 Poise), to remove residual stresses caused by cooling after forming. Granulates and powdered glass are produced by crushing and sieving the annealed plates into desired particle fractions. Also quenching the melt between stainless steel plates or pouring the melt into deionized water are further steps in the process of granulate fabrication. However, bioactive glasses start to react easily in aqueous solutions, which might affect the composition of the particle surfaces. Crushing and sieving increase the risk of contamination from the equipment used in the particle manufacture. Thus, in all processing of bioactive glasses into specific shapes, care should be taken in order to minimize any contamination.

1.2.2 Viscosity Viscosity is important in determining the melting parameters for achieving a bubble-free and homogeneous melt. Glasses are usually melted at temperatures corresponding to the viscosity value 10 to 100 dPa·s. The low viscosity facilitates easy elimination of the gases by buoyancy from the melt. Viscosity and its change with temperature is the most crucial factor in determining the forming and shaping procedures that can be used for a particular composition. The approximate viscosity values of interest in forming bioactive glasses into various shapes are summarized in Table 1.1.



7

Table 1.1 Approximate viscosity values (dPa·s) for bioactive glass forming processes Processing

Viscosity (η) in dPa·s

Melting Pressing Drawing of continuous fibres Sinter glass powder to porous body Annealing

10–102 104–106 102.5–103.5 108–109 1012–1013

1.1 Viscosity–temperature points for 45S5, S53P4 and 13-93 (Vedel et al., 2008).

Figure 1.1 shows measured viscosity–temperature points for three bioactive glasses, 45S5, S53P4 and 13–93 at the low and high temperature ranges. The measured values are according to Vedel et al. (2008). The oxide composition of the glasses are given in Table 1.2. The dashed lines between the low and high temperature ranges give typical viscosity–temperature curves for glass forming melts. However, at the intermediate temperatures bioactive glasses crystallize and melt viscosity does not exist. The viscosity values at different temperatures are important criteria in glass forming. The high temperature values correlate with melt-forming processes, while the low-temperature values specify the suitability of the glass, for example for sintering into porous bodies or firing as a coating on metal implants. The high temperature values of glasses 45S5 and S53P4 in Fig. 1.1 could be measured by rotational viscometer only at values below 100 dPa·s (Vedel et al., 2008). Thus, it is likely that for these glasses the liquidus temperature, i.e. the temperature at


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

8 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Table 1.2 Nominal oxide compositions 45S5, S53P4 and 13–93 in wt% (mol%) Oxides in wt% (mol%) Glass

Na2O

45S5 S53P4 13–93

24.5 (24.4) 23 (22.7) 6 (6)

K2O

12 (7.9)

MgO

CaO

P2O5

SiO2

5 (7.7)

24.5 (26.9) 20 (21.8) 20 (22.1)

6 (2.6) 4 (1.7) 4 (1.7)

45 (46.1) 53 (53.8) 53 (54.6)

Sources: Hench et al., 1973 (45S5), Andersson et al., 1990 (S53P4) and Brink et al.,1997 (13–93).

which crystallization commences on cooling, is close to the lowest experimental high temperature values shown in Fig. 1.1. Further, the low viscosity values at their liquidus suggest that these glasses can only be formed by casting. At the low temperature range, 45S5 and S53P4 crystallize at around 109 dPa·s. This means that the compositions cannot be sintered through viscous flow into porous bodies without extensive crystallization. The strong tendency to crystallize within a large temperature range has been utilized in the manufacture of various glass-ceramics of the parent glass 45S5 (Chen and Boccaccini, 2006; Chen et al., 2006b; Boccaccini et al., 2007). The high-temperature viscosity values of glass 13–93 in Fig. 1.1 suggest that it can be formed without extensive crystallization to around 104 dPa·s. This composition has be pressed, blown and drawn into continuous fibres (Brink, 1997; Pirhonen et al., 2006). At the low temperature range crystallization starts below 108 dPa·s, which indicates that the glass can be sintered into porous glassy bodies (Ylänen et al., 2000; Fu et al., 2010). The bioactive glasses 45S5 and S53P4 that crystallize easily on thermal treatments contain only four oxides. In an attempt to decrease the crystallization tendency potassium oxide, magnesia and boron oxide have been added to the formulations. The viscosity of glasses within the system Na2O-K2O-MgO-CaOB2O5-P2O5 has been discussed in three extensive studies (Brink, 1997: Karlsson and Rönnlöf, 1998; Vedel et al., 2008). The studies had two goals: to get reliable data on the high-temperature properties of bioactive glasses and to meet the need to find compositions that can be formed by other methods than casting. Both Karlsson/Rönnlöf and Vedel et al. suggest for the viscosity–temperature relationship an Arrhenius-type expression ˚I nη = A + B/T, where the constants A and B are expressed by the oxide composition of the melt. Vedel et al. also give models for calculating the constants A, B and T0 in the Vogel–Fulcher–Tamman equation for the viscosity–temperature relationship, logη = –A + B/(T – T0). However, as no experimental points exist for the intermediate temperature range, a single relationship over the whole temperature range is questionable. Models for calculating temperatures for certain high- and low-temperature viscosity points



9

additively from the oxide composition are likely to be of reasonable accuracy (Vedel et al., 2008). Temperatures T at certain viscosities η were expressed as additive functions of the oxide content of the glass in wt%. The temperature values were based on certain experimental points when measuring the viscosity by hot-stage microscopy and rotational viscometer. The composition dependency was calculated for 30 statistically chosen glass compositions within the system Na2O-K2O-MgO-CaO-B2O3-P2O5-SiO2. The influence of composition on the temperature at a typical viscosity for sintering of porous bodies is given in equation 1.1. Equation 1.2 can be used to estimate the temperature at a viscosity value typical for glass melting. The rather low viscosity value was explained by the strong tendency of several glasses to crystallize at a relatively low viscosity (Vedel et al., 2008). Tη =1010 dPas (°C) = −148.036 + 3.566·xNa O + 7.071·xMgO 2

+ 9.740·xCaO + 7.770·xB O + 8.347·xP O 2 3

2 5

+ 9.287·xSiO

[1.1]

2

Tη =1015 dPas (°C) = 117.816 + 7.730·xK O + 6.078·xMgO 2

+ 18.469·xP O + 19.150·xSiO 2 5

2

[1.2]

The equations are valid for the composition range (wt%): Na2O (5 to 25), K2O (0 to 15), MgO (0 to 6), CaO (15 to 25), B2O3 (0 to 4), P2O5 (0 to 4), SiO2 (50 to 65).

1.2.3 Thermal expansion Bioactive glasses have been studied as coatings on metal prostheses in order to provide biological fixation to bone (Lacefield and Hench, 1986; Hench and Andersson, 1993b; Andersson et al., 1995; Bloyer et al., 1999; Moritz et al., 2004a, 2004b; Krause et al., 2006; Borrajo et al., 2007; Lopez-Esteban et al., 2009). The rapid surface reactions of bioactive glasses can limit their use as thin coatings. In such applications, glass compositions with a lower reactivity are preferred. If the coating is applied via traditional enamelling firing, the thermal expansion of the glass should be compatible with the metal. The coefficient of thermal expansion can be calculated additively from the composition with the factors suggested by Appen (1974). A compositional model for the linear thermal expansion α between 20 and 300°C for bioactive glasses was suggested by Karlsson and Rönnlöf (1998): see equation 1.3.

α (10−6 K−1) = 3.625 + 0.345·xNa O + 0.266·xK O 2

2

+ 0.098·xCaO + 0.064·xP O

2 5

[1.3]

The equation is valid for compositions with (wt%) 5 to 25 Na2O, 0 to 15 K2O, 0 to 5 MgO, 10 to 20 CaO, 0 to 3 B2O3, 0 to 6 P2O5 and 39 to 70 SiO2. © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

10 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

1.2.4 Crystallization characteristics The importance of the liquidus temperature for glass melts arises from the fact that all melt-forming should be performed at temperatures higher than the liquidus. Surprisingly little information is available on accurate liquidus temperatures of bioactive glasses. According to Hench, the choice of the first compositions of bioactive glasses was partially based on finding a composition that is close to a ternary eutecticum in the system Na2O-CaO-SiO2 (Hench, 2006). This ternary system is commonly used to interpret liquidus temperatures and crystallization paths for soda-lime-silica glasses, it is also used when they contain some additional components to the three oxides. Figure 1.2 shows the phase equilibria and some liquidus surfaces in the silica-rich corner of the Na2O-CaO-SiO2 system (Morey and Bowen, 1964). Compositions of bioactive glasses 45S5, S53P4 and 13–93 are superimposed in the ternary system by calculating all alkalis, alkaline earths and glass formers in wt% into Na2O, CaO and SiO2, respectively. Figure 1.2 shows that a simple superimposing of the four component glasses 45S5 and S53P4 as well as glass 13–93 containing also the oxides of potassium and magnesium (c.f. Table 1.2) would suggest liquidus temperatures around or higher than 1200°C for all compositions. According to thermal analysis the crystals formed in heating glass 45S5 above transition temperature start to melt between 1080 and 1160°C (Chatzistavrou et al., 2006; Hall, 2007; Arstila et al., 2008a; Bretcanu et al., 2009) and the offset of the melting is between 1250 and 1260°C (Chatzistavrou et al., 2006; Arstila et al., 2008a). The offset temperatures for melting of crystals in S53P4 and 13–93 are 1210 to 1230°C and 1020 to 1180°C, respectively (Arstila et al., 2008a). The offset temperatures for 45S5 and S53P4 are close to the values suggested by the liquidus surface in Fig. 1.2. Until more accurate data is available for systems containing some phosphorus oxide, the ternary Na2O-CaO-SiO2 system seems to give an acceptable first approximation of liquidus. The larger differences in the measured liquidus and the value approximated from the simple three-component phase diagram for glass 13–93 containing also potassium oxide and magnesia indicate that other approaches to estimate liquidus are needed for glasses consisting of several oxides. One way could be composing empirical models for calculating the liquidus from the oxide composition of the glass. In that case, separate models for different primary phase fields should be considered as suggested by Karlsson et al. (2002). Thermodynamic modelling is likely to provide the most appropriate method for estimating the liquidus in future. So far, reliable thermodynamic data of the liquid phase in the compositional area of bioactive glasses are not available. In the ternary Na2O-CaO-SiO2 (Fig. 1.2) the compositions of 45S5 and S53P4 fall into the primary phase field of Na2Ca2Si3O9, while 13–93 is within the CaSiO3 field. Interestingly, these phases seem to be the most commonly observed primary phases formed in the thermal treatment of bioactive glasses within the compositional range given for equations 1.1 to 1.2 (Arstila et al., 2008b). The



11

1.2 Liquidus surfaces and phase equilibria in the system Na2O-CaOSiO2. Composition points of glasses 45S5, S53P4 and 13–93 are superimposed into the diagram by assuming alkalis, alkaline earths and P2O5 by wt% in Na2O, CaO and SiO2, respectively (redrawn from Morey and Bowen, 1964).

sodium oxide rich compositions form sodium calcium silicates, while glasses with less sodium oxide form mainly CaSiO3 (Arstila et al., 2008b). Na2Ca2Si3O9, Na2CaSi3O8 and Na2CaSi2O6 are the most often suggested compositions of the primary phase observed in the thermal treatment of 45S5 (Rizkalla et al., 1996; Chen et al., 2006; Lefebvre et al., 2007, 2008; Arstila et al., 2008b; Bretcanu et al. 2009). The discrepancy observed in the primary phase can partially be explained by the solid solutions of Na2Ca2Si3O9 and Na2CaSi3O8. During crystallization of a stoichiometric glass Na2Ca2Si3O9 the nucleation did not start with the stoichiometric composition but approached to that during crystallization (Fokin and Zanotto, 2007). The final composition of the fully crystallized glass was close to that of the parent glass. The primary phase, PCT, as well as the crystallization temperature, Tx, in the thermal treatment of bioactive glasses was predicted by additive functions from the oxide composition, equations 1.4 and 1.5 (Arstila et al., 2008b). Numerical values of PCT lower than 1.5 indicate sodium calcium silicate crystals, while values higher than 1.5 suggest the formation of CaSiO3.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

12 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses PCT = 0.018−0.107·xNa O + 0.002·x2Na O + 0.052·xMgO 2

+

0.032·x2

P2O5

2

+ 0.038·xSiO

[1.4]

2

Tx (°C) = 191.70 − 9.51·xNa O + 5.91·xMgO + 5.93·xCaO 2

+ 15.29·xB O + 12.92·xP O + 9.77·xSiO 2 3

2 5

2

[1.5]

The exact crystallization temperature depends on the heating rate. However, the crystallization temperature suggested by equation 1.5 can be used to compare the crystallization characteristics of different compositions and to give approximate values for the onset of crystallization in the thermal treatment of glasses. Generally, glasses forming CaSiO3 crystallize at temperatures 100 to 200°C higher than the glasses showing sodium calcium silicate crystals (Arstila et al., 2008b). On cooling the properties of glass forming melts turn into solid like properties at the glass transition temperature, Tg. As the exact transition temperature for a certain composition depends on the cooling rate, changes in several properties can be observed over a temperature range known as the glass transformation region. The transition temperature thus determines the lower limit for thermal processing of glass melts. The temperature span for sintering through viscous flow ranges from glass transformation up to crystallization temperature. The stability of glasses against crystallization upon thermal treatment is often described with relative values calculated from the temperatures of glass transition, liquidus and crystallization (Hruby, 1972; Zanotto, 1987). The reduced glass transition temperature Tgr, given by the ratio of the glass transition Tg to liquidus Tl, correlates with the general trends of nucleation in glass forming melts. A good glass former has Tgr = Tg/Tl higher than or equal to two-thirds, while lower values suggest crystallization. The crystallization kinetics depend also on the numerical value of Tgr; higher values than 0.58 indicate surface crystallization, while for lower values volume nucleation dominates (Zanotto, 1987). The compositional dependence of glass transformation temperature Tg of bioactive glasses has been described as an additive function of the oxide composition (Andersson, 1992; Karlsson and Rönnlöf, 1998; Arstila et al., 2008b). Equation 1.6 gives the glass transformation temperature based on dilatometric data (Karlsson and Rönnlöf, 1998), while the factors in equation 1.7 are based on thermal analysis (Arstila et al., 2008b). Tg (°C) = 635.9 − 5.26·xNa O + 3.18·xK O − 3.16·xMgO 2

2

− 0.47·xNa O·xK O − 0.12·xK2O·xCaO 2

[1.6]

2

Tg (°C) = −122.85 + 2.59·xK O + 5.36·xMgO + 9.37·xCaO 2

+ 4.08·xB O + 6.45·xP O + 8.04·xSiO 2 3

2 5

2


[1.7]


13

Table 1.3 Calculated and measured values of some physical properties for glasses 45S5, S53P4 and 13–93 Property

45S5

S53P4

13–93

Reference

T (°C) at η = 1010 dPa·s T (°C) at η = 101.5 dPa·s Tliquidus(°C) PTC Tx (°C) Tg (°C)

559 (570) – (–) – (1210) 1.5 (NCS) 621 (647) 506 (–)

572 (587) 1233 (1235) – (1188) 1.1 (NCS) 661 (693) 514 (–)

651 (655) 1330 (>1330) – (–) 2.2 (CS) 852 (853) 564 (–)

Tg (°C) α (10−6K−1)

507 (530) 14.9 (–)

516 (541) 13.8 (–)

574 (600) 11.1 (–)

α (10−6K−1)

15.2 (–)

14.1 (–)

11.7 (–)

Vedel et al., 2008 Vedel et al., 2008 Arstila et al., 2008a Arstila et al., 2008b Arstila et al., 2008b Karlsson and Rönnlöf, 1998 Arstila et al., 2008b Karlsson and Rönnlöf, 1998 Appen, 1974

Notes: The measured values are given in parenthesis. – means not available using the model/not measured.

In both models the oxide composition is given in wt%. The validity range of equation 1.6 is the same as for equation 1.3. The validity of equation 1.7 is according to equation 1.1. Some measured and calculated physical property values of bioactive glasses 45S5, S53P4 and 13–93 are summarized in Table 1.3. The differences between the calculated and measured values indicate that the equations can be used to estimate the physical property values of the bioactive glasses.

1.3

Chemical properties and bioactivity

1.3.1 Surface reaction mechanisms Controlled surface reactivity leading to tissue bonding is the key characteristic of bioactive glasses. Compared to soda-lime-silica glasses they have poor chemical durability in aqueous solutions. The low silica and high sodium oxide content contribute to the low chemical durability of bioactive glasses. The reactions of glasses in aqueous solutions are generally described by two main mechanisms: the exchange of alkali ions in the surface with H+ and H3O+, and network dissolution through the attack of hydroxyl ions on the silica structure. Therefore, the reactions are controlled by the pH of the surrounding and interfacial solutions. The chemical durability is commonly measured with methods developed for characterizing soda-lime-silica glasses, e.g. durability in water (ISO 719) and alkaline solutions (ISO 695). Water resistance according to ISO 719 is measured as the volume of hydrochloric acid needed to neutralize the pH increase of pure water after 1 hour’s contact time with a certain fraction of glass particles at 98°C.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

14 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses Table 1.4 Consumption of HCl (ml, 0.01 M), and hydrolytic resistance class (HGB) according to ISO 719. pH values of solutions were measured before the acid titration at room temperature Sample

VHCl (ml)

HGB

pH

E-glass* Float glass 1–98** 13–93 S53P4 45S5

0.3 ± 0.01 1.0 ± 0.03 2.3 ± 0.1 2.8 ± 0.1 3.4 ± 0.1 5.3 ± 0.2

3 4 5 5 5 >5

9.18 ± 0.02 9.88 ± 0.10 10.24 ± 0.01 10.43 ± 0.05 10.71 ± 0.03 10.79 ± 0.01

Notes: * E-glass (wt%): 0.1 Na2O, 0.7 K2O, 0.7 MgO, 23.5 CaO, 6.4 B2O3, 14.1 Al2O3, 53.9 SiO2. **1–98 (wt%): 6 Na2O, 11 K2O, 5 MgO, 22 CaO, 1 B2O3, 2 P2O5, 53 SiO2 (Itälä et al., 2002). Source: Taipale et al., 2008; Fagerlund et al., 2010.

The acid consumed is related with the amount of Na+ ions leached from the glass expressed as mass of Na2O. The Na2O amount is compared with a relative scale from 1 for glasses with very good water durability to 5 for glasses with poor durability. Table 1.4 gives the hydrolytic resistance, HGB, of e-glass and bioactive glasses 45S5, S53P4 and 13–93 and 1–98. E-glass fibres have been used to develop porous fibre-reinforced composites for use as load-bearing orthopedic implants (Mattila et al., 2009). The bioactivity of glass 1–98 has been verified by in vivo experiments (Itälä et al., 2002). E-glass has medium resistance, while the bioactive glasses have very low resistance. The acid consumption of 45S5 is higher than the limit given for HGB = 5 in the ISO 719 standard. Table 1.4 also shows the pH of the solutions before the titration. The pH values of the bioactive glasses are very high, thus indicating network dissolution at particle surfaces. The large differences in the hydrochloric acid consumption between the bioactive glasses suggest that a procedure according to ISO 719 gives a rapid method of comparing their overall biodegradability. However, the relative 1 to 5 HGB values fail to give any relevant information on the bioactive glasses. Understanding the surface reactions of glasses is of utmost importance when selecting compositions for different clinical applications. Silicate glass surfaces have been divided into five different classes according to their reactivity in different environments (Fig. 1.3; Hench and Clark, 1978; Hench, 1992). The type I surface, typical for high silica surfaces in neutral solutions develops only a very thin hydrous layer. Type II surfaces are typical for commercial soda-lime glasses. The alkalis in the surface layer leach in solutions with pH 9 to 10, the network structure of silicate glasses is attacked and the glass surface undergoes a congruent dissolution as described by the type V surface. Bioactive glass surfaces might show type V reactions if surrounding solution pH increases to high values due to insufficient solution circulation. Thus, the surface area to volume (SA/V) affects the reaction type.

1.3.2 Bioactive glasses in vitro The tissue bonding of bioactive silicate glasses can be related with their ability to form a dual layer of silica gel and hydroxyapatite on the surface in body fluids. The time-dependent reaction stages of interfacial reactions leading to tissue bonding of bioactive glasses have been discussed in detail by Hench and co-workers (Hench, 1991; Hench and Andersson, 1993a; Hench and Best, 2004). The reactions stages in biological solutions before the interaction of the surface with proteins and cells are described in five stages at the interface between the glass and solution in Table 1.5. The special characteristics of the hydroxyapatite layer formed on the glass in the reaction stages 1 to 5 allow biochemical adsorption of growth factors and other biological moieties resulting in rapid formation of new bone as discussed by Hench (1998). Depending on their ability to lead to both osteoconduction and osteoproduction or only osteoconduction, the bioactive materials are further divided into Class A and Class B (Hench, 1998, 2006). In this chapter, however, only factors affecting the formation of the calcium phosphate rich layer and crystallized hydroxyapatite on the glass surface are discussed. These inorganic Table 1.5 Reaction stages of bioactive glass surfaces in biological solutions Stage

Reaction

1

Exchange of alkalis from the glass surface with H+ or H3O+ in the solution. The rapid reaction is diffusion controlled and proportional to the square root of the time (Douglas and El-Shamy, 1967; Hench and Clark, 1978). Breaking of siloxane bonds in the glass interface leading to loss of soluble silica in the form of Si(OH)4 to the solution. The loss of silica is directly proportional to time. Condensation and repolymerization of a SiO2 rich layer on the surface. In this reaction stage the thickness of the silica rich layer increases. Migration of Ca2+ and PO42− ions to the surface through the SiO2 rich layer to an amorphous calcium phosphate rich layer on the top of the SiO2 layer. The layer grows by incorporation of Ca2+ and PO42− ions from the solution. Crystallization of the amorphous calcium phosphate layer by incorporation of OH−, CO32−, or F− anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer.

2

3 4

5

Source: Hench, 1991.



17

reaction stages are also observed on bioactive glass surfaces in aqueous solutions such as Tris buffer (Hench, 2006). In solutions containing phosphate, the amorphous calcium phosphate-rich layer might form also on P2O5-free glasses (Ogino et al., 1980). The rate of calcium phosphate formation (stage 4) and the time of onset of crystallization (stage 5) vary greatly with composition. If these two stages take place slowly, the material is not bioactive (Hench, 1998). Surface reactions of the first bioactive glasses were studied in water or Tris buffer solution (Clark et al., 1976; Hench and Clark, 1978; Andersson and Kangasniemi, 1991). Since Kokubo and co-workers (Kokubo et al., 1990; Kokubo, 1991) developed the so-called simulated body fluid (SBF), it has been used widely to measure the bioactivity of glasses in vitro. Simulated body fluid is designed to contain the inorganic constituents of human blood and is thus assumed to provide similar conditions to those found in vivo. Over the years other similar types of solutions have also been suggested, and the composition of SBF has been adjusted (Oyane et al., 2003; Takadama et al., 2004; Kokubo and Takadama, 2006). When immersing bioactive glasses in SBF the five reaction stages described in Table 1.5 can be identified. The concentrations of different elements leached from the glasses into the dissolution medium at different immersion times have been analysed spectrophotometrically or with inductively coupled plasma analysis (ICP) (Jones et al., 2001; Clupper et al., 2003; Cerruti et al., 2005a; Zhang et al., 2009; Zhang et al., 2010). Surface layers formed on the glass at different immersion times have been analysed, e.g. by FTIR, SEM-EDXA, Raman spectroscopy, TF-XRD (Clark et al., 1976; Hench, 1991; Andersson and Kangasniemi, 1991; Ohtsuki et al., 1992; Kim et al., 1995; Rehman et al., 1998; Jones et al., 2001; Sepulveda et al., 2002; Notingher et al., 2003; Clupper et al., 2003; Cerruti et al., 2005a; Zhang et al., 2009). Formation of a bone-like apatite in SBF is often taken as an indication of in vivo bioactivity of the glass (Kokubo and Takadama, 2006). Glass 45S5 is a common reference for bioactive glass studies. Most studies on the influence of composition on the bioactivity deal with the effect of changing one or two components in the original 45S5 composition (Hench, 1991; Hench and Andersson, 1993a; O’Donnell et al., 2009; Gentleman et al., 2010). General approaches based on the mean number of non-bridging oxygen ions in the silica tetrahedron have been suggested to correlate with the bioactivity of glasses (Strnad, 1992; Strnad and Koga, 1999). Zhang et al. (2009) used changes in simulated body fluid and in glasses after immersion to characterize and give basis to models for the in vitro bioactivity as functions of the oxide composition. The coefficients in the additive models for pH of simulated body fluid as functions of the oxide composition at different immersion times are given in Table 1.6. At the shortest immersion times, the best prediction of the pH was given by models in which only the sums of alkalis and alkaline earths expressed as weight percentage were taken into account. At longer immersion times the contribution of the network-forming components B2O3, P2O5 and SiO2 affected the predicted pH of


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

18 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Table 1.6 Factors for calculating pH additively from glass composition at different immersion times in SBF Factor for pH at

Constant Na2O + K2O MgO + CaO B2O3 B2O32 P2O5 SiO2 SiO22

4h

8h

24 h

72 h

168 h

6.876 1.778·10−2 1.611·10−2

6.852 2.114·10−2 1.729·10−2

7.017 6.465·10−4 1.952·10−2

4.416 1.977·10−3 6.739·10−2 6.873·10−2

6.482 1.258·10−3 3.405·10−2 8.562·10−3

5.908·10−2

3.609·10−2

6.641·10−3 1.089·10−2

2.867·10−4

Notes: Validity range according to equation 1.1. The glass composition is given in wt%. Source: Zhang et al., 2009.

Table 1.7 Calculated and experimental pH of SBF at 4 and 72 h immersion of glasses 45S5, S53P4 and 13–93 in SBF at 37°C, SA/V = 0.4 cm−1 pH of SBF

At 4 h At 72 h

45S5

S53P4

13–93

calc.

exp.

calc.

exp.

calc.

exp.

7.77 8.19

7.62 8.17

7.65 7.85

7.60 7.85

7.64 7.78

7.64 7.78

Source: Zhang et al., 2009.

SBF. Thus, the models based on regression analysis of measured pH nicely correlate with the reaction stages of bioactive glasses given in Table 1.5. Calculated and experimental values for the pH of SBF at 4 and 72 h immersion of glasses 45S5, S53P4 and 13–93 are given in Table 1.7. Zhang et al. (2009) also suggested a model for the silica gel thickness formed on glasses at 72 h (TLSi) immersion in SBF. The gel thickness was expressed by using relative values from 1 to 5, equation 1.8. The thickness ranges are 1 for no silica layer at all, 2: 0.1 to 3.3 µm, 3: 3.3 to 6.5 µm, 4: 6.5 to 9.8 µm, and 5 for a silica layer thicker than 9.8 µm. The oxides in the equation are given in wt% (Zhang et al., 2009). TLSi = 6.5387 + 7.847·10−2·xNa O + 8.725·10−3·xK2 O 2

2

− 5.747·10−1·xB O + 1.529·10−1·xB2 O − 7.753·10−1·xP O 2 3

2 3

+ 2.207·10−1·xP2 O − 1.573·10−3·x2SiO 2 5

2


2 5

[1.8]


19

A model has been suggested also for formation of the hydroxyapatite layer on the glass surface as a function of the oxide composition of the glass at 72 h immersion is SBF (Zhang et al., 2009). The layer formation is described by a relative number, bioactivity number (BN): see equation 1.9. Glasses with a uniform hydroxyapatite layer with a measured thickness > 0.9 µm have BN = 3. Glasses with incompletely developed hydroxyapatite layers with a layer thickness 0.1 to 0.9 µm are described as BN = 2. Glasses showing only spot like hydroxyapatite formation have BN = 1. When using equation 1.9, BN > 2.5 indicates high bioactivity and BN < indicates 1.5 low bioactivity. Intermediate bioactivity is suggested by values 1.5 < BN < 2.5. BN = 124.680 − 1.545·xNa O − 4.500·10−3·x2Na O − 1.681·xK O 2

2

2

− 2.067·xMgO + 4.199·10−2·x2MgO − 1.661·xCaO − 2.238xB O

2 3

+

1.314·10−1·x2B O3 2

− 1.680·xP O − 2 5

1.551·10−2·x2SiO 2

[1.9]

The models of in vitro bioactivity can be utilized as a first estimation of the reactivity of the glass when developing new compositions. Layer development on glass surfaces in SBF is commonly used as an indication of bioactivity. As SBF is a supersaturated solution towards calcium phosphate precipitation, the glass surface condition might affect the layer formation (Karlsson et al., 2002; Bohner and Lemaitre, 2009). The sample form, surface condition, surface area to volume ratio and fluid circulation also affect the reactions (Greenspan et al., 1994; Jones et al., 2001; Cerruti et al., 2005b; Zhang et al., 2008a, 2008b).

1.3.3 Bioactive glasses in vivo The tissue bonding properties of glasses and glass-ceramics within the system Na2O-CaO-P2O5-SiO2 system have been studied by Hench and co-workers. The compositional dependence of bonding type is illustrated in the Na2O-CaO-SiO2 phase diagram, Fig. 1.4 (Hench, 2006). The bioactivity inside the region A is given for compositions containing 6 wt% P2O5. These glasses bond to bone. The surface reactions of these glasses are described by IIIA in Fig. 1.3. The level of bioactivity increases to the middle of the region A. The compositions inside the region S bond also to soft tissue. The composition of 45S5, E, is in the middle of this region. Glasses in the silica-rich corner, region B, are almost inert and elicit a fibrous tissue around the implant, while glasses C resorb and disappear within 10 to 30 days of implantation (Hench, 1991). The silica-rich compositions B correspond to Type I surface in Fig. 1.3. The dashed lines in the figure give the overall glass-forming tendency of different compositions in the Na2O-CaO-SiO2 system. The glasses in the silica-rich corner are difficult to melt with conventional methods, while compositions below the lowest dashed line do not form glasses. The most bioactive compositions crystallize easily, thus preventing manufacture of other than quenched or cast products as discussed above. Most commercial


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

20 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

1.4 Compositional dependence (in wt%) of bone bonding and softtissue bonding of bioactive glasses and glass-ceramics in the Na2OCaO-SiO2 system (redrawn from Hench, 2006). Notes: All compositions in region A contain 6 wt% P2O5. Tissue bonding of compositions within the different areas: A – bone bonding, B – nonbonding (reactivity too low), C – non-bonding (reactivity too high), S – soft tissue bonding, E – Bioglass® 45S5 composition, square – commercial soda-lime glasses. The dashed lines limit approximate composition ranges of glass-forming tendencies in the ternary system.

soda-lime glasses have compositions close to the region marked with the black square in the figure. Compositions within the soda-lime glass range can be manufactured into fibers, blown and pressed products, etc. It should be observed that the soda-lime glasses also contain other components but the three-component diagram gives a simplified overall illustration of the system. Two models have been developed for estimating in vivo bioactivity from the oxide composition of the glass (Andersson et al., 1990; Brink et al., 1997). Both models are based on observing the bone contact of glass cones implanted in rat tibia. The layer formation was studied with SEM-EDXA of the cross-sections at 8 weeks’ implantation. The model by Andersson et al. is valid within the compositional range (wt%) 15 to 30 Na2O, 10 to 25 CaO, 0 to 3 B2O3, 0 to 8 P2O5, 0 to 3 Al2O3 and 45 to 65.5 SiO2. An alumina addition was found to inhibit the bone bonding. Depending on the formation of SiO2, the calcium phosphate-rich layers and the bone bonding, the in vivo reactions were divided into five classes. These were given relative reaction numbers (RN) (Andersson et al., 1990): • •

nearly inert glasses showing only small changes in the surface and hardly any bone contact: RN = 1 fairly high solubility and bone contact but not bonding. The glasses show formation of a SiO2-rich layer but no calcium phosphate accumulation: RN = 2


Melt-derived bioactive glasses • • •

21

fairly high solubility and bone contact but not bonding. Formation of a SiO2rich layer and limited calcium phosphate accumulation: RN = 3 formation of a calcium phosphate-rich surface layer but no bone bonding: RN = 4 formation of a calcium phosphate-rich surface layer and bone bonding: RN = 6.

The relative reaction numbers, RN, correlate with the bone response so that glasses with BN >5 are bioactive. Equation 1.10 expresses the reaction number from the oxide composition given in wt% (Andersson et al., 1990): RN = 88.3875 − 0.980188·xNa O − 1.12306·xCaO 2

− 0.560527·x2B O − 1.20556·xP O − 2.808689·xAl O 2 3

2 5

2 3

− 0.0116272·x2SiO

[1.10]

2

The calculated RN numbers for 45S5 and S53P4 indicate bone bonding. The composition of 13–93 is outside the validity range of equation 1.10. Andersson and his co-workers suggested that calcium phosphate forms in vivo within the silica-rich layer, not on the top of it. Bone contact was good for glasses with a thick silica-rich outer layer. Such compositions were suggested to be biocompatible but do not bone with bone. Further, the formation of a calcium phosphate-rich layer on the surface was found insufficient as the only indication of bioactivity. Later, Andersson et al. verified the bone bonding of bioactive glasses with pushout tests in vivo (Andersson et al., 1992). Table 1.8 summarizes the results for glass surface reactions and the push-out test forces when using titanium cone as control. The inert glasses do not show any marked changed in the surface composition. These glasses are not chemically bonded to bone but encapsulated in connective tissue. Accordingly, the push-out strength of inert glasses is low.

Table 1.8 Summary of layer formation and push-out strengths at 8 weeks in rabbit tibia. The glass codes give the silica and phosphorus pentoxide content in the glasses by wt%. The other constituents are Na2O, CaO, B2O4 and Al2O3 Sample

Silica layer

HCA layer

Bone response

Strength (MPa)

Glass reaction

S65.5P1 S52P3 S52P8 S45P7 S46P0 S55.5P4 Titanium

Thin Yes Yes Yes Yes Yes

No No Yes Yes Yes Yes

None Contact Contact Bonding Bonding Bonding Contact

0.5 ± 0.4 3.6 ± 0.9 3.0 ± 0.5 23.0 ± 2.9 16.4 ± 3.9 19.9 ± 4.0 2.2 ± 0.6

Almost inert Soluble, single layer Soluble, dual layer Soluble, dual layer Soluble, dual layer Soluble, dual layer

Source: Andersson et al., 1992.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

22 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Glasses with silica-rich layer formation but no hydroxyapatite precipitation show bone contact but no chemical bone bonding. The push-out force is comparable to titanium. Glasses with a silica-rich layer in the surface are biocompatible. Bioactive glasses forming a dual layer of silica and hydroxyapatite show bonding forces varying from one- to ten-fold values compared to titanium. In vivo bioactivity by layer formation on glasses within the system Na2O-K2OMgO-CaO-B2O3-P2O5-SiO2 has been expressed also by another relative number, the index of surface bioactivity (ISA) (Brink et al., 1997). The ISA number is related with the formation of surface layers on glasses in vivo: inert glass = 1, silica-rich layer = 2, layered structure = 3, bioactive = 4: equation 1.11. The compositional range of the equation is according to equation 1.1. The oxides are expressed in wt%:

[1.11] The calculated ISA values for 45S5 and S53P4 suggest that the glasses are bioactive, while 13–93 has an ISA value 3.2. This glass was, however, found to bond to bone and thus to be bioactive (Brink et al., 1997). The observations suggest that the models for in vivo bioactivity can be used to get a first estimation of the bioactivity of a specific glass composition. Also the different models describing the in vitro bioactivity correlate with the in vivo observations. The melt-derived bioactive glasses 45S5 and S53P4 are FDA-approved for certain clinical applications. When compared in vitro, the thickness of the dual layer was somewhat less in S53P4 than in 45S5 after 1 week in simulated body fluid (Hupa et al., 2010). Only small differences in the layer thickness on implants of the two glasses were observed after 8 weeks in the soft tissue of rats (Hupa et al., 2010). Both glasses bonded to bone but the layer thickness at 8 weeks was slightly less in S53P4. In clinical applications these glasses are used mainly in different bone-filling applications as glass particulates. The bioactive glasses have been found safe and promising as bone substitutes for the treatment of benign bone tumours (Lindfors, 2009; Lindfors et al., 2009; Lindfors et al., 2010). In a clinical follow-up study granules of S53P4 used to fill bone cavities were found to increase the cortical thickness (Lindfors et al., 2009). Even after 14 years some remnants of glass particles could be identified (Lindfors et al., 2010). The long-term clinical study demonstrated, however, the potential of bioactive glasses in benign bone tumour surgery both in children and adults. However, the observations suggest that the dissolution of bioactive glasses is slow. Although the in vitro, sin vivo and clinical observations give similar reaction trends, a physiochemical reaction mechanism and dissolution of bioactive glasses is not fully understood.



1.4

23

Future trends

Bioactive glass research during the past 30 years indicates that the bioactivity and physical as well as chemical properties can be adjusted by changes in the oxide composition of the glasses. Several properties describing the relationship between the oxide composition of the glass and in vitro, in vivo and other properties are available. These models are valuable when tailoring glass compositions for various clinical applications. Much bioactive glass research has dealt with the development of compositions that bond to tissue. Although detailed long-term physicochemical reactions are still obscure, an essential understanding of the mechanisms behind the interaction of the biological environment and the glass surface reactions has been established. As glasses are brittle materials, their increasing use in medical applications is restricted by their unpredictable mechanical behaviour in load bearing applications. Therefore, the use of glasses in composites together with biodegradable organic polymers has been one of the main focuses during the recent years. This often means the use of glasses as particulates, thin fibres or thin-walled sintered structures. Due to the large surface area the glasses are likely to react rapidly, and thus they lose their mechanical reinforcing capacity. This gives rise to the need of finding compositions that would be osteoconductive but not resorb too rapidly. Although much of the reactivity of glasses in biological solutions is understood, the mutual interaction of glasses, for example with bioresorbable polymers is not fully understood. Thus, research and development into bioactive glass compositions for novel medical applications will still be topical for some time to come.

1.5

References

Andersson, Ö. H., Liu, G., Karlsson, K. H., Niemi, L., Miettinen, J., Juhanoja, J., 1990. In vivo behavior of glasses in the SiO2-Na2O-CaO-P2O5-Al2O3-B2O3 system. Journal of Materials Science: Materials in Medicine, 1 (4), 219–27. Andersson, Ö. H, Kangasniemi, I., 1991. Calcium phosphate formation at the surface of bioactive glass in vitro. Journal of Biomedical Materials Research, 25 (8), pp. 1019–30. Andersson, Ö. H., 1992. Glass transition temperature of glasses in the SiO2-Na2O-CaOP2O5-Al2O3-B2O3 system. Journal of Materials Science: Materials in Medicine, 3 (5), 326–8. Andersson, Ö. H., Liu, G., Kangasniemi, K., Juhanoja, J., 1992. Evaluation of the acceptance of glass in bone. Journal of Materials Science: Materials in Medicine, 3 (2), 145–50. Andersson, Ö. H., Karlsson, K. H., Hero, H., Vedel, E., Yli-Urpo, A., Pajamäki, K. J. J., Lindholm, T. S., 1995. Bioactive double glass coatings for Co-Cr-Mo alloy. Journal of Materials Science: Materials in Medicine, 6, 242–7. Appen, A. A., 1974. Chemie des Glases. In: H. Scholze 1988. Glas. Natur, Struktur und Eigenschaften, 3rd Edition. Berlin, Heidelberg, New York: Springer Verlag, pp. 176–7. Arstila, H., Tukiainen, M., Taipale, S., Kellomäki, M., Hupa, L., 2008a. Liquidus temperatures of bioactive glasses. Advanced Materials Research (Zurich, Switzerland), 39–40 (Glass), pp. 287–92.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

24 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Arstila, H., Vedel, E., Hupa, L., Hupa, M., 2008b. Predicting physical and chemical properties of bioactive glasses from chemical composition. Part 2: devitrification characteristics. Glass Technology: European Journal of Glass Science and Technology, Part A, 49 (6), 260–65. Bloyer, D. R., Gomez-Vega, J. M., Saiz, E., McNaney, J. M., Cannon, R. M., Tomsia, A. P., 1999. Farbrication and characterization of a bioactive glass coating on titanium implant alloys. Acta Materialia, 47 (15), 4421–4. Boccaccini, A. R., Chen, Q., Lefebvre, L., Gremillard, L., Chevalier, J., 2007. Sintering, crystallization and biodegradation behaviour of Bioglass-derived glass-ceramics. Faraday Discussions, 136, 27–44. Bohner, M., Lemaitre, J., 2009. Can bioactivity be tested in vitro with SBF solution? Biomaterials, 30, 2175–9. Borrajo, J. P., Serra, J., Gonzalez, P., Leon, B., Munoz, F. M., Lopez, M., 2007. In vivo evaluation of titanium implants coated with bioactive glass by pulsed laser deposition. Journal of Materials Science: Materials in Medicine, 8 (2), 2371–6. Bretcanu, O., Chatzistavrou, X., Paraskevopoluos, K., Condradt, R., Thompson, I., Boccaccini, A. R., 2009. Sintering and crystallization of 45S5 Bioglass® powder. Journal of the European Ceramic Society, 29, 3299–306. Brink, M., 1997. The influence of alkali and alkaline earths on the working range for bioactive glasses. Journal of Biomedical Materials Research, 36 (1), 109–17. Brink, M., Turunen, T., Happonen, R. P., Yli-Urpo, A. (1997). Compositional dependence of bioactivity of glasses in the system Na2O-K2O-MgO-CaO-B2O3-P2O5-SiO2. Journal of Biomedical Materials Research, 37 (1), 114–21. Cerruti, M., Greenspan, D., Powers, K., 2005a. Effect of pH and ionic strength on the reactivity of Bioglass 45S5. Biomaterials, 26 (14), 1665–74. Cerruti, M. G., Greenspan, D., Powers, K., 2005b. An analytical model for the dissolution of different particle size samples of Bioglass in TRIS-buffered solution. Biomaterials, 26 (24), 4903–11. Chatzistavrou, X., Kontonasaki, E., Chrissafis, K., Zorba, T., Koidis, P., Paraskevopoulos, K. M., 2006. Surface and bulk contributions in the crystallization process of a bioactive glass. Key Engineering Materials, 309–311 (Pt. 1, Bioceramics), pp. 313–16. Chen, Q. Z., Boccaccini, A. R., 2006. Poly(D,L-lactic acid) coated 45S5 Bioglass-based scaffolds: processing and characterization. Journal of Biomedical Materials Research, Part A, 77A (3), 445–57. Chen, Q. Z., Thompson, I. D., Boccaccini, A. R., 2006. 45S5 Bioglass-derived glassceramic scaffolds for bone tissue engineering. Biomaterials, 27 (11), 2414–25. Clark, A. E., Pantano, C. G., Hench, L. L., 1976. Auger spectroscopic analysis of bioglass corrosion films. Journal of the American Ceramic Society, 59 (1–2), 37–39. Clupper, D. C., Gough, J. E., Hall, M. M., Clare, A. G., LaCourse, W. C., Hench, L. L., 2003. In vitro bioactivity of S520 glass fibers and initial assessment of osteoblast attachment. Journal of Biomedical Materials Research, Part A, 67A (1), 285–94. Douglas, R. W. and El-Shamy, T. M. M., 1967. Reactions of glasses with aqueous solutions. Journal of the American Ceramic Society, 50 (1), 1–8. Fagerlund, S., Ek., P., Hupa, M., Hupa, L., 2010. On determining chemical durability of glasses, submitted. Fokin, V. M., Zanotto, E. D., 2007. Continuous compositional changes of crystal and liquid during crystallization of a sodium calcium silicate glass. Journal of Non-Crystalline Solids, 353 (24–25), 2459–68.



25

Fu, Q., Rahaman, M. N., Bal, B. S., Brown, R. F., 2010. Preparation and in vitro evaluation of bioactive glass (13–93) scaffolds with oriented microstructures for repair and regeneration of load-bearing bones. Journal of Biomedical Materials Research, Part A, 93A (4), 1380–90. Gentleman, E., Fredholm, Y. C., Jell, G., Lotfibakhshaiesh, N., O’Donnell, M. D., Hill, R. G., Stevens, M. M., 2010. The effects of strontium-substituted bioactive glasses on osteoblasts and osteoclasts in vitro. Biomaterials, 31 (14), 3949–56. Greenspan, D. C., Zhong, J. P., LaTorre, G. P., 1994. Effect of surface area to volume ratio on in vitro surface reactions of bioactive glass particulates. Bioceramics, Proceedings of the International Symposium on Ceramics in Medicine, 7, 55–60. Haimi, S., Gorianc, G., Moimas, L., Lindroos, B., Huhtala, H., Räty, S., Kuokkanen, H., Sandor, G. K., Schmid, C., Miettinen, S., Suuronen, R., 2009. Characterization of zinc-releasing three-dimensional bioactive glass scaffolds and their effect on human adipose stem cell proliferation and osteogenic differentiation. Acta Biomaterialia, 5 (8), 3122–31. Hall, M. M., 2007. Influence of hydroxyl content on selected properties of 45S5 bioactive glass. Journal of Biomedical Materials Research, Part A, 83A (3), 720–24. Hench, L. L., Paschall, H. A., 1973. Direct chemical bond of bioactive glass-ceramic materials to bone and muscle. Journal of Biomedical Materials Research, 7 (3), 25–42. Hench, L. L., Clark, D. E., 1978. Physical chemistry of glass surfaces. Journal of NonCrystalline Solids, 28, 83–105. Hench, L. L., 1991. Bioceramics: From concept to clinic. Journal of the American Ceramic Society, 74 (7), 1487–510. Hench, L. L., 1992. Surface modification of bioactive glasses and ceramics. In: D. E. Clark and K. Z. Zoitos, eds. Corrosion of Glass, Ceramics and Ceramic Superconductors: Principles, Testing, Characterization and Applications. Park Ridge: Noyes Publications, pp. 298–314. Hench, L. L., Andersson, Ö., 1993a. Bioactive glasses. In: L. L. Hench and J. Wilson, eds. An introduction to bioceramics, Advanced Series in Ceramics, Vol.1. Singapore: World Scientific, pp. 41–56. Hench, L. L., Andersson, Ö. H., 1993b. Bioactive glass coatings. In: L. L. Hench and J. Wilson, eds. An Introduction to Bioceramics. Advanced series in Ceramics, Vol 1. Singapore: World Scientific, pp. 239–60. Hench, L. L., 1998. Bioceramics. Journal of the American Ceramic Society, 81 (7), 1705–28. Hench, L. L. and Best, S., 2004. Ceramics, glasses and glass-ceramics. In: B. D. Ratner, A. S. Hoffman, F. J. Shoen, J. E. Lemons, eds. Biomaterials Science, An Introduction to Materials in Medicine (2nd Edition). Elsevier, pp. 153–170. Hench, L. L., 2006. The story of Bioglass®, Journal of Materials Science: Materials in Medicine, 17, 967–78. Hruby, A., 1972. Evaluation of glass forming tendency by means of DTA, Czechoslovak Journal of Physics, 22 (11), 1187–93. Hupa, L., Karlsson, K. H., Hupa, M., Aro, H. T., 2010. Comparison of bioactive glasses in vitro and in vivo. Glass Technology: European Journal of Glass Science and Technology, Part A, 51 (2), 89–92. Itälä, A., Ylänen, H. O., Yrjans, J., Heino, T., Hentunen, T., Hupa, M., Aro, H. T., 2002. Characterization of microrough bioactive glass surface: surface reactions and osteoblast responses in vitro. Journal of Biomedical Materials Research, 62 (3), 404–11.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

26 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Jones, J. R., Sepulveda, P., Hench, L. L., 2001. Dose-dependent behavior of bioactive glass dissolution. Journal of Biomedical Materials Research, 58 (6), pp. 720–26. Jones, J. R., 2008. Bioactive glass. In: T. Kokubo, ed. Bioceramics and Their Clinical Applications. Cambridge, UK: Woodhead Publishing Ltd., pp. 266–86. Jones, J. R., 2009. New trends in bioactive scaffolds: The importance of nanostructure. Journal of the European Ceramic Society, 29 (7), 1275–81. Karlsson, K. H., Rönnlöf, M., 1998. Property-composition relationships for potentially bioactive glasses. Glass Science and Technology (Frankfurt/Main), 71 (5), 141–5. Karlsson, K. H., Backman, R., Hupa, L., 2002. Models for liquidus temperatures. In Vacher, R., ed. Proceedings: 2002 Glass Odyssey, 6th ESG Conference, July 2–6, 2002, Montpellier: Université Montpellier, CD-ROM. Karlsson, Kaj H., Backman, R., Hupa, M., 2002. An equilibrium study of phosphate precipitation on bioactive glass. Key Engineering Materials, 218–20 (Bioceramics-14), 103–8. Kim, H. M., Miyaji, F., Kokubo, T., Ohtsuki, C., Nakamura, T., 1995. Bioactivity of Na2O-CaO-SiO2 glasses. Journal of the American Ceramic Society, 78 (9), 2405–11. Kokubo, T., Kushitani, H., Sakka, S., Yamamuro, T. J., 1990. Solutions able to reproduce in vivo surface-structure changes in bioactive glass-ceramic A-W, Journal of Biomedical Materials Research, 24, 721–34. Kokubo, T., 1991. Bioactive glass ceramics: properties and applications, Biomaterials, 12, 155–63. Kokubo, T., Takadama, H., 2006. How useful is SBF in predicting in vivo bone bioactivity? Biomaterials, 27 (15), 2907–15. Krause, D., Thomas, B., Leinenbach, C., Eifler, D., Minay, E. J., Boccaccini, A. R. 2006. The electrophoretic deposition of Bioglass particles on stainless steel and Nitinol substrates. Surface and Coatings Technology, 200 (16–17), 4835–45. Lacefield, W. R. and Hench, L. L., 1986. The bonding of Bioglass® to a cobalt-chromium surgical implant alloy. Biomaterials, 7, 104–8. Lefebvre, L., Chevalier, J., Gremillard, L., Zenati, R., Thollet, G., Bernache-Assolant, D., Govin, A., 2007. Structural transformations of bioactive glass 45S5 with thermal treatments. Acta Materialia, 55, 3305–13. Lefebvre, L., Gremillard, L., Chevalier, J., Zenati, R., Bernache-Assolant, D., 2008. Sintering behaviour of 45S5 bioactive glass. Acta Biomaterialia, 4, 1894–903. Li, R., Clark, A. E., Hench, L. L., 1991. An investigation of bioactive glass powders by sol-gel processing. Journal of Applied Biomaterials: an official journal of the Society for Biomaterials, 2 (4), 231–9. Lindfors, N. C., 2009. Treatment of a recurrent aneurysmal bone cyst with bioactive glass in a child allows for good bone remodelling and growth. Bone (Amsterdam, Netherlands), 45 (2), 398–400. Lindfors, N. C., Heikkilä, J. T., Koski, I., Mattila, K., Aho, A. J., 2009. Bioactive glass and autogenous bone as bone graft substitutes in benign bone tumors. Journal of Biomedical Materials Research, Part B: Applied Biomaterials, 90B (1), 131–6. Lindfors, N. C., Koski, I., Heikkilä, J. T., Mattila, K., Aho, A. J., 2010. A prospective randomized 14-year follow-up study of bioactive glass and autogenous bone as bone graft substitutes in benign bone tumors. Journal of Biomedical Materials Research, Part B: Applied Biomaterials, 94B (1), 157–64. Lopez-Esteban, S., Gutierrez-Gonzalez, C. F., Gremillard, L., Saiz, E., Tomsia, A. P., 2009. Interfaces in graded coatings on titanium-based implants. Journal of Biomedical Materials Research, Part A, 88A (4), 1010–21.



27

Lusvardi, G., Malavasi, G., Menabue, L., Aina, V., Morterra, C., 2009. Fluoride-containing bioactive glasses: Surface reactivity in simulated body fluids solutions. Acta Biomaterialia 5, pp. 3548–62. Mattila, R. H., Laurila, P., Rekola, J., Gunn, J., Lassila, L. V. J., Mäntylä, T., Aho, A. J., Vallittu, P. K., 2009. Bone attachment to glass-fibre-reinforced composite implant with porous surface. Acta Biomaterialia, 5 (5), 1639–46. Morey, G. W. and Bowen, N. L., 1964. Na2O-CaO-SiO2 system. In E. M. Levin, C. R. Robbins, H. E. McMurdie, eds. Phase Diagrams for Ceramists. Ohio: American Ceramic Society, p. 175. Moritz, N., Vedel, E., Ylänen, H., Jokinen, M., Hupa, M., Yli-Urpo, A., 2004a. Characterization of bioactive glass coatings on titanium substrates produced using a CO2 laser. Journal of Materials Science: Materials in Medicine, 15 (7), 787–94. Moritz, N., Rossi, S., Vedel, E., Tirri, T., Ylänen, H., Aro, H., Närhi, T., 2004b. Implants coated with bioactive glass by CO2-laser, an in vivo study. Journal of Materials Science: Materials in Medicine, 15 (7), 795–802. Notingher, I., Jones, J. R., Verrier, S., Bisson, I., Embanga, P., Edwards, P., Polak, J. M., Hench, L. L., 2003. Application of FTIR and Raman spectroscopy to characterization of bioactive materials and living cells. Spectroscopy (Amsterdam, Netherlands), 17 (2, 3), 275–88. O’Donnell, M. D., Watts, S. J., Hill, R. G., Law, R. V., 2009. The effect of phosphate content on the bioactivity of soda-lime-phosphosilicate glasses. Journal of Materials Science: Materials in Medicine, 20 (8), 1611–18. Ogino, M., Ohuchi, F., Hench, L. L., 1980. Compositional dependence of the formation of calcium phosphate films on Bioglass. Journal of Biomedical Materials Research, 14, 55–64. Ohtsuki, C., Kokubo, T., Yamamuro, T., 1992. Compositional dependence of bioactivity of glasses in the system calcium oxide-silicon dioxide-aluminum oxide: its in vitro evaluation. Journal of Materials Science: Materials in Medicine, 3 (2), 119–25. Oyane, A., Kim, H. M., Furuya, T., Kokubo, T., Miyazaki, T., Nakamura, T., 2003. Preparation and assessment of revised simulated body fluids. Journal of Biomedical Materials Research A, 65, 188–95. Pirhonen, E., Niiranen, H., Niemelä, T., Brink, M., Tormälä, P., 2006. Manufacturing, mechanical characterization, and in vitro performance of bioactive glass 13–93 fibers. Journal of Biomedical Materials Research, Part B: Applied Biomaterials, 77B (2), 227–33. Rehman, I., Knowles, J. C., Bonfield, W., 1998. Analysis of in vitro reaction layers formed on bioglass using thin-film X-ray diffraction and ATR-FTIR microspectroscopy. Journal of Biomedical Materials Research, 41 (1), 162–6. Rizkalla, A. S., Jones, D. W., Clark, D. B., Hall, G. C., 1996. Crystallization of experimental bioactive glass compositions. Journal of Biomedical Materials Research, 32, 119–24. Sepulveda, P., Jones, J. R., Hench, L. L., 2002. In vitro dissolution of melt-derived 45S5 and sol-gel derived 58S bioactive glasses. Journal of Biomedical Materials Research, 61 (2), 301–11. Strnad, Z., 1992. Role of the glass phase in bioactive glass-ceramics. Biomaterials, 13 (5), 317–21. Strnad, Z. and Koga, N., 1999. Compositional dependence of concentration of nonbridging oxygen ions in silicate glasses – correlation to bioactivity. Advances in Science and Technology (Faenza, Italy), 28 (Materials in Clinical Applications), pp. 49–54.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

28 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Taipale, S., Ek, P., Hupa, M., Hupa, L., 2008. Continuous measurement of the dissolution rate of ions from glasses. Advanced Materials Research, Vols. 39–40, pp. 341–6. Takadama, H., Hashimoto, M., Mizuno, M., Kokubo, T., 2004. Round-robin test of SBF for in vitro measurement of apatite-forming ability of synthetic materials, Phosphorus Research Bulletin, 17, 119–25. Vedel, E., Arstila, H., Ylänen, H., Hupa, L., Hupa, M., 2008. Predicting physical and chemical properties of bioactive glasses from chemical composition. Part 1: viscosity characteristics. Glass Technology: European Journal of Glass Science and Technology, Part A, 49 (6), 251–9. Vrouwenvelder, W. C., Groot, C. G., de Groot, K., 1994. Better histology and biochemistry for osteoblasts cultured on titanium-doped bioactive glass: Bioglass 45S5 compared with iron-, titanium-, fluorine- and boron-containing bioactive glasses. Biomaterials, 15 (2), 97–106. Watts, S. J., Hill, R. G., O' Donnell, M. D., Law, R. V., 2010. Influence of magnesia on the structure and properties of bioactive glasses. Journal of Non-Crystalline Solids, 356, 517–24. Williams, D. F., 1999. The Williams Dictionary of Biomaterials. Liverpool, UK: Liverpool University Press. Ylänen, H., Karlsson, K. H., Itälä, A., Aro, H. T., 2000. Effect of immersion in SBF on porous bioactive bodies made by sintering bioactive glass microspheres. Journal of Non-Crystalline Solids, 275 (1, 2), 107–15. Zanotto, E. D., 1987. Isothermal and adiabatic nucleation in glass. Journal of NonCrystalline Solids, 89, 361–70. Zhang, D., Hupa, M., Aro, H. T., Hupa, L., 2008a. Influence of fluid circulation on in vitro reactivity of bioactive glass particles. Materials Chemistry and Physics, 111 (2–3), 497–502. Zhang, D., Hupa, M., Hupa, L., 2008b. In situ pH within particle beds of bioactive glasses. Acta Biomaterialia, 4 (5), 1498–505. Zhang, D., Vedel, E., Hupa, L., Aro, H. T., Hupa, M., 2009. Predicting physical and chemical properties of bioactive glasses from chemical composition. Part 3: in vitro reactivity. Glass Technology: European Journal of Glass Science and Technology, Part A, 50 (1), 1–8. Zhang, D., Leppäranta, O., Munukka, E., Ylänen, H., Viljanen, M. K., Eerola, E., Hupa, M., Hupa, L., 2010. Antibacterial effects and dissolution behavior of six bioactive glasses. Journal of Biomedical Materials Research, Part A, 93A (2), 475–83.


2 Surface modification of bioactive glasses J. CHANG and Y. L. ZHOU , Chinese Academy of Sciences, China and Y. ZHOU , Shanghai Jiao Tong University, China

Abstract: The surface modification of biomaterials has exhibited great potential in biological applications by transforming the existing surface into more appropriate compositions and/or topographies. The surface of bioactive glasses plays a critical role in their performance, and many studies have been conducted on the surface modification of bioactive glasses or modification of other materials with bioactive glasses, which can combine the properties of the two materials together to obtain desirable functions. In this chapter, the surface modification of bioactive glasses and using bioactive glasses are broadly reviewed in the following three aspects: (1) surface modification of bioactive glasses to improve bioactivity; (2) surface modification of bioactive glasses surface with organic molecules to improve the dispersivity; (3) surface modification of bioinert materials with bioactive glasses. Key words: bioactive glasses, surface modification, biomineralization, coating techniques.

2.1

Introduction

Bioactive glasses (BGs), which were first synthesized 40 years ago, have been extensively studied as an artificial bone grafting material for bone repairs, and have gained great acceptance in clinical applications (Hench, 1998, 2006). As an implant for hard tissue repair, the surface of BGs plays an important role in their properties after implantation (Kenny and Buggy, 2003; Hench and Polak, 2002; Gabbi et al., 1995). In order to control the surface properties precisely, modification of the surface of a certain bioactive glass (BG) or the surface of a substrate with BG are usually employed which can combine the properties of both the coatings and the substrates to obtain desirable functions (Verne et al., 2009a; Wang, 2009). Surface modification of BGs has promising potential for converting the existing surface into more desirable compositions and/or topographies for biological application (Duan and Wang, 2006). Surface modification of BGs can be broadly classified into three categories: 1) treating the surface of BGs to improve the surface bioactivity through silanization, biomineralization and microroughening; 2) modifying BG particles to render their compatibility with another phase (Neouze and Schubert, 2008) through condensation or a grafting reaction (Slowing et al., 2008; Vallet-Regi, 2006; Vallet-Regi et al., 2007); 3) coating bio-inert materials (such as titanium alloy) with BGs by physical or chemical techniques such as plasma spray coating, physical vapor deposition, sol-gel process etc (Liu et al., 2008b). 29 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

30 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

In this chapter, these three categories of the surface modification will all be introduced.

2.2

Surface modification of bioactive glasses to improve bioactivity

Surface characteristics are critical in developing a biomaterial implant, since the interaction between the biomaterial and the ambient physiological medium takes place at the surface of the implant (Iucci et al., 2004). After implantation of BGs into the body, numerous physiological reactions occur simultaneously in the intermediate layer between the targeted tissue and the implant. Taking the first synthesized bioactive glass, 45S5 Bioglass® (Table 2.1) as an example, the surface reactions that occurred at the Bioglass® surface are summarized in Table 2.2 (Stroganova et al., 2003; Hench, 1991; Lobel and Hench, 1996). These surface reactions occur within the first 12 to 24 hours after implantation. The first step is a rapid reaction of the release of sodium ions (Na+) from the surface of the glass via ion exchange with a hydrogen ion from the environment Table 2.1 The composition of 45S5 Bioglass® Compound

Percentage (wt%)

SiO2 CaO Na2O P2O5

45.0 24.5 24.5 6.0

Table 2.2 45S5 Bioglass® reaction stages with increasing time Increasing time Stage 11 10 9 8 7 6 5 4 2–3 1 0

Reaction event Crystallization of matrix Cellular attachment Differentiation of stem cells Attachment of stem cells Action of macrophages Adsorption of biological moieties (proteins, etc) Nucleation and crystallization of calcium phosphate to HCA Precipitation of amorphous calcium phosphate Dissolution and repolymerization of surface silica Sodium hydrogen ion exchange Initial glass surface



31

(H+ or H3O+), which results in a negatively charged surface. Then a silica-rich layer is formed on the surface of Bioglass® as the loss of Na+ causes the breakdown of the silica network, with the resultant formation of Si(OH)4 groups. After that, an amorphous calcium phosphate (Ca-P) layer starts to form on the surface of the silica-rich layer and eventually incorporates the biological moieties, such as blood proteins, growth factors and collagens. The adsorption of the organic species from the body fluid occurs concurrently with the first three reaction stages, which is believed to contribute at least partially to the biological nature of the hydroxyapatite (HA) layer. Within about three to six hours in vitro, the Ca-P layer will crystallize into the HA layer, which has been described as the bonding layer. Since the composition of the HA layer on the surface of the Bioglass® is chemically and structurally similar to the mineral component of natural bone, it allows the regenerated tissue to attach directly onto the surface. As the reaction continues, this surface HA layer grows in thickness to form a bonding zone of up to 100 microns, which is essential and mechanically compliant to maintain the binding of the implant with the natural tissue. The HA layer is also able to adsorb biological moieties and support cell adhesion and growth. Furthermore, the releasing of ionic components from the glass surface has been shown to continue for a long period of time, and enhances the development of the surface reactive layers. In this way, BG participates in the whole repair process, leading to the creation of a direct bond of the material to the tissue. Therefore, the term ‘bioactive’ means that, through a series of interfacial ion exchange reactions, a silica-rich gel layer forms followed by the formation of the Ca-P layer on the BG surface, which promotes the interfacial bonding with tissues after implantation and enhance new tissue regeneration (Montanaro et al., 2002). Based on the description above, the surface characteristics, the silica-rich layer, biomineralization of the Ca-P layer, and immobilizing proteins and growth factors are essential for the bioactivity of BGs. Therefore, many works have been undertaken to modify the surface of BGs through silanization, deposition of Ca-P layer and adsorption with proteins, to improve the bioactivity of BGs. It is believed that a biomaterial’s surface characteristics can alter cell behavior at many levels (Lipski et al., 2008). Surface morphology can directly influence the cellular response, for example, by inducing the release of growth factors and cytokines from the adhered osteoblasts (Montanaro et al., 2002). A smaller-scale surface roughness (in the range from 10 nm to 10 mm) can increase the surface areas of BGs and may influence the biological interaction with cells and large biomolecules. Gough et al. compared smooth and roughened 45S5 Bioglass® in vitro and found that the roughened samples enhanced mineralized nodule formation in human primary osteoblast culture (Gough et al., 2004). Therefore, creating microroughness on the surface of BGs is an important approach to further enhance the bioactivity of the glasses. In this section, surface modification of BGs were discussed in three parts: silanization, biomineralization of Ca-P and the surface structure.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

32 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

2.2.1 Silanization The silica gel layer forming on the surface of BGs is regarded as the first step for bone–BGs bonding. On the silica layer, an amorphous Ca-P phase is precipitated, which evolves with time into HA (Gao et al., 2001; Lusvardi et al., 2009). Thereby the concept of silanization is important for the surface modification of BGs. One of the common non-toxic protein-coupling agents is 3-aminopropyltriethoxysilane (APTS), and it is used to silanize the surface of BGs to promote protein adhesion and cell growth on biological implants (Howarter and Youngblood, 2006). Since there are plenty of silanol on the surface of BGs (Andrade et al., 2004), APTS can be surface-modified on BG particles through a wet-chemical method in a dynamic inert nitrogen atmosphere (Chen et al., 2008c). The reaction process of Si-OH on the surface of BGs with APTS is shown in Fig. 2.1. Highly porous 45S5 Bioglass®derived glass-ceramic scaffolds were surface silane functionalized with APTS, which can favor the formation of HA and improve cell attachment and growth. The osteoblasts proliferated better on the functionalized glass-ceramic surface (Chen et al., 2006, 2008b). After surface modification with APTS, the introduced amino-groups can be used for protein grafting. One example is the immobilization of a model protein, carnosine, on the surface of BGs to bind human bone morphogenetic protein (BMP) (Verne et al., 2009b). BMPs are growth factors that are known to induce bone and bone marrow regeneration, among which BMP-2 and BMP-7 have already been applied clinically for bone regeneration due to high osteoinductive activity (Takahashi et al., 2005). Alkaline phosphatase (ALP) participates in the bone formation and mineralization process, and is widely used as a marker of osteoblast differentiation (Groeneveld et al., 1995). Verne et al. have demonstrated that ALP covalently grafted to APTS-modified BGs can enhance the ability of the material to induce HA precipitation in simulated body fluid (SBF) while maintaining the enzyme activity (Verne et al., 2010). Furthermore, protein-loaded BG scaffolds can also be used as a delivery system. For example, porous 45S5 Bioglass®-derived scaffolds were surface-functionalized with APTS and loaded with collagen. After functionalization, the stability of the collagen attachment and the release stability against the pH change in the biological environment were increased, to create an environment suitable for enhancing cell attachment (Chen et al., 2008a). Laminin can also be adsorbed on the APTS surface-modified

2.1 Modification process of bioactive glasses (BGs) with 3-aminopropyltriethoxysilane (APTS).



33

binary 70S30 BGs (70mol%SiO2-30mol%CaO) and 58S BG (60mol%SiO236mol%CaO-4mol%P2O5) foams. Sustained and controlled release of laminin from the modified scaffolds, which has a beneficial effect on tissue formation (Lenza et al., 2003), was achieved over a 30-day period. It is known that nitric oxide (NO) plays an important role in regulating osteoblasts and osteoclasts in bone metabolism. APTS has also been used for preparation of NO-releasing BG materials (Pryce and Hench, 2004).

2.2.2 Biomineralization of calcium phosphate on the surface of bioactive glasses One critical event that facilitates bone formation is the interaction of BGs with physiological solutions and the subsequent formation of a Ca-P rich layer on the glass surface, which is important in the development of osteoconductive biomaterials for orthopaedic applications (Leonor et al., 2009). Apatite is the major mineral phase of which the hard tissues such as bone and dentin are composed (Kim et al., 2004; de Arenas et al., 2006). Therefore, the induction of apatite formation is another important approach to modifying the surface of BGs. For this purpose, different solutions have been applied for soaking the BGs in order to facilitate the formation the Ca-P layer on their surface. Simulated body fluid (SBF) (see Table 2.3) is a common solution for soaking BGs and apatite formation, since its ionic concentrations are nearly equal to those of human blood plasma (Kokubo and Takadama, 2006). In addition, Tris(hydroxymethyl) aminomethane (Tris) solution at different pHs (Cerruti et al., 2005) or with electrolytes typical for plasma (TE, a solution with 142.0 mM Na+, 5.0 mM K+, 1.5 mM Mg2+, 2.5 mM Ca2+, 148.8 mM Cl−, 4.2 mM HCO3−, and 1.0 mM HPO42−) can also be used to treat BGs for deposition Table 2.3 Ionic concentrations and pH of simulated body fluid and human blood plasma Concentration (mM)

Na+ K+ Mg2+ Ca2+ Cl− HCO3− HPO42− SO42− pH

Simulated fluid

Blood plasma

142.0 5.0 1.5 2.5 147.8 4.2 1.0 0.5 7.25

142.0 5.0 1.5 2.5 103.0 4.2 1.0 0.5 7.20–7.40

Source: Fujibayashi et al., 2003.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

34 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

of HA on the surface. Adsorption of serum proteins onto HA layer serves to enhance bone cell adhesion, proliferation, and function (El-Ghannam et al., 1997; Radin et al., 2005). Moreover, some investigations have demonstrated that certain proteins incorporated into the soaking solution can regulate the capability of forming Ca-P layer on the surface of BGs thus affecting the bioactivity of the BGs. A study has shown that the addition of fibronectin can reduce the electronegativity of the BG surface and consequently delay the formation of both the amorphous and the crystalline Ca-P layers (Lu et al., 2001). Radin and coworkers studied the influence of serum proteins on the formation of Ca-P layer on the surface of BG granules. Serum can promote porous surface structure of Ca-P layer on the Surface of BGs (Radin et al., 1997, 2000), while recombinant porcine amelogenin rP172 showed a modulation effect on the oriented growth of apatite crystals on 45S5 Bioglass® surface in supersaturated calcifying solution (SCS). Grafting functional groups on the surface of BGs can also alter the bioactivity of BGs. With the addition of the amino and carboxylic groups to mesoporous bioactive glasses (MBGs) through post-grafting process, the nucleation, growth rate and morphology of carbonated HA formed on the surface of the glasses were affected remarkably. MBGs functionalized by amino groups promote the formation of spherical HA particles, while the nucleation and growth rate of HA on MBGs functionalized by carboxylic groups decrease to a large extent with the increasing carboxylic group content (Sun et al., 2008). In brief, a Ca-P layer is osteoconductive, and vital for the bioactivity of BGs (Leonor et al., 2009). Surface modification of BGs by biomineralization of the Ca-P layer and subsequent immobilization of specific proteins can enhance cell attachment, proliferation and differentiation.

2.2.3 The role of surface structure Chemical treatment is an easy way to modify the surface morphologies and increase microroughening of the surface without changing the body properties (Li et al., 2009). BG microspheres (215 to 350 µm) were etched in different etching solutions (from pH = 1.0 to 13) and etching time (soaking time from a few seconds to 30 min) to achieve the desired surface appearance with an average roughness value of 0.35 to 0.52 µm (Itala et al., 2001). The microrough surface can accelerate the formation of Si-gel layer during the first hours of SBF and Tris immersion, and can enhance the attachment of human osteoblast-like MG-63 cells during the first 24 h of incubation in vitro (Itala et al., 2002). An in vivo study in a rabbit model also demonstrated that microroughened surface can enhance osteopromotive properties and bone-bonding response (Itala et al., 2003a). Furthermore, surface microroughness can cause a temporary change in the expression of specific genes (Itala et al., 2003b; Valimaki et al., 2005). Bioleaching is a soft biochemical approach that can obtain nanoscale surface under gentle condition. Borosilicate glasses were treated through fungus-based bioleaching and the surfaces were



35

morpho-chemical modified with monodispersed ultrafine (~5 ± 0.5 nm) silicate nanoparticles through acidolysis (Kulkarni et al., 2008).

2.3

Surface modification of bioactive glasses using organic molecules to improve dispersivity

Although BGs show favorable bioactivity, lack of in situ moldability and relative brittleness limit their applications (Rich et al., 2002). Therefore, polymers/BG composites have been fabricated in order to obtain biomaterials with improved properties (Misra et al., 2008; Maquet et al., 2004; Jiang et al., 2005; Silva et al., 2004). Among the factors affecting the properties of the polymers/inorganic composites, the interface adhesion of inorganic particles to polymer matrix plays an important role (Hong et al., 2005; Supova, 2009). It has been found that some polymers/inorganic composites lost their strength rapidly in a physiological environment, and the failure occurred mainly at the interface between inorganic particles and the polymer matrix (Zhang et al., 2005). The main reason was the tendency for inorganic particles to agglomerate in the polymer matrix owing to their small dimensions and incompatible polarity with polymers (Cheng and Chang, 2006; Borum-Nicholas and Wilson, 2003; Liu et al., 2008a). Therefore, improving poor dispersion of inorganic particles in polymeric matrix was critical for preparing composite materials with improved properties. Several chemical reactions have been employed to modify BGs with organic molecules based on the reaction between Si-OH on the surface of BG particles and the functional group of organic molecules. The isocyanate-ended low-molecular-weight PLLA was reacted with BG particles (Si:P:Ca = 29:13:58 weight ratio) (particle size 40 nm) at 80°C for about 12 h to obtain surface-modified BG particles. The grafting resulted in an increase in phase compatibility, and the consequent improvement of tensile strength, tensile modulus and impact energy. An in vitro bioactivity test showed that, compared to pure PLLA scaffold, the BG/PLLA nanocomposite possessed a greater capability of inducing the formation of an apatite layer on the scaffold surface (Liu et al., 2008a). In another study, silane coupling agent 3-glycidoxypropyltrimethoxysilane was introduced to modify the surface of BG particles through condensation. The filler dispersion and phase compatibility between poly (D,L-lactide) (PDLLA) and BG particles were also improved (Zhang et al., 2009a). Stearic acid is widely used in the surface modification of inorganic materials, because its carboxyl groups can chemically bind with -Si-OH on the surface of BGs, and the long alcohol chain has high compatibility with polymers (Grassi et al., 2003; Saleema and Farzaneh, 2008). Dodecyl alcohol can also be used to modify BGs through an esterification reaction. The esterification of silanols with alcohols is well known as described in the following equation: -Si-OH + R-OH ← → Si-OH

HO-R← → Si-OR + H2O


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

36 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

2.2 The SEM images (×500) of the composite films composed with unmodified and modified bioactive glass particles: (a) 45S5, (b) m-45S5, (c) mesoporous 58S, (d) m-mesoporous 58S, (e) 58S, (f) m-58S (Gao and Chang, 2009).

In this study, three kinds of BGs (45S5, 58S and mesoporous 58S BGs) were surface-modified with dodecyl alcohol through esterification at 260°C to improve the homogeneous dispersion of BG particles in polymeric matrix. A SEM observation (Fig. 2.2) illustrated that the modified BG particles were homogeneously dispersed in the PDLLA matrix. The modified composite films can still induce the formation of HA on its surface after immersion in SBF, and the



37

2.3 SEM images (with ×500 and ×1000 images on the left column and ×60 000 images on the right column) of the surface of films after immersion in SBF for 7 days (a, b) PDLLA/58S composite film, (c, d) PDLLA/m-58S composite film (Gao and Chang, 2009).

distribution of HA was more homogeneous on the film (Fig. 2.3) (Gao and Chang, 2009). However, the disadvantage of the modification with dodecyl alcohol is the decrease of hydrophilicity, which may affect the biocompatibility of the composite materials, since increased surface hydrophilicity is known to be associated with enhanced protein adsorption and consequent cell adhesion and proliferation on biomaterials (Zhang et al., 2009b). Fortunately, this modification is reversible and the dodecyl alcohol can be removed after the achievement of homogenous dispersion of BG particles in composite materials by hydrolytic treatment in hot water. The properties (such as tensile strength) of the composite films after treatment will not be affected. Most importantly, cells on the composite films after hydrolysis show the highest proliferation rate and differentiation level (Fig. 2.4 and Fig. 2.5) (Zhou et al., 2010). In summary, non-toxic organic molecules, especially biocompatible molecules are useful to modify the surface of BGs for the improvement of the dispersivity in polymer matrix.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

38 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

2.4 The dMSC proliferation of the composite films. OD value on y-axis represents the number of living cells (p < 0.05) (Zhou et al., 2010).

2.5 ALP activity of dMSCs after culturing on the composite films for different periods (p < 0.05) (Zhou et al., 2010).



2.4

39

Surface modification of bioinert materials using bioactive glasses

Owing to excellent bulk properties, such as relatively low modulus, good fatigue strength, corrosion resistance and biocompatibility, metallic materials including stainless steel, commercially available pure titanium (Ti), Ti alloys (such as Ti6Al4V) and nickel-Ti etc, are widely used for orthopedic applications (Ramaswamy et al., 2009). However, they are bioinert, owing to their absent or weak interaction with living tissues upon implant (Navarro et al., 2008). Interfacial movement under external stress leads to loosening and deterioration of the mechanical fit, which causes pain and eventually leads to clinical failure of the bioinert implants (Tilocca, 2009; Shi et al., 2002). One of the most general solutions is coating the bioinert substrates with bioactive materials, which can modify the surface bioactivity and maintain the mechanical properties of the metal implants at the same time (Vallet-Regi et al., 2003; Liu et al., 2008b). Several bioactive materials have been evaluated for implant coatings. BGs are one of the most promising candidates for implant coating because of high bonebonding ability through the formation of a HA layer on the surface and excellent resorbability in the body fluids (Lewandowska et al., 2007; Ballo et al., 2008). The bone-bonding ability of BGs is also helpful for providing right fixation of the implants. Three factors are essential for preparing BG coatings onto implants successfully: 1) the thermal expansion coefficients of BGs and the metal substrates should be similar in order to avoid the generation of large thermal stresses that can result in coating cracking or delaminating during fabrication; 2) the coating should have good adhesion to the substrates (Bolelli et al., 2007b); 3) the deposited coatings should be bioactive in SBF (Ma et al., 2006). Various coating methods have been employed to create BGs or BG composite layers onto metallic substrates, including the plasma-sprayed deposition technique, electrophoretic deposition, the sol-gel process and magnetron sputtering, and so on (Verne et al., 2005). Each of these coating techniques has its own strengths and weaknesses. In this section, different coating techniques using BGs are introduced.

2.4.1 Enameling technique The conventional enameling technique is a simple way to coat metal substrates by BGs. The advantages of this technique are low cost, ease of operation and optimization by changing the parameters. Experimentally, it is possible to prepare BG coatings with good adherence to Ti substrate by controlling the composition and process conditions (firing time, temperature and atmosphere) (Bloyer et al., 1999). For example, by carefully controlling the composition of BGs and the treatment of the enameling and glazing process, BG coatings with a thermal expansion coefficient congruent with the alumina or Ti alloy can be obtained (Gomez-Vega et al., 1999; Pazo et al., 1998). These kinds of BG coatings can


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

40 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

induce the precipitation of a thick HA layer when soaked in SBF (Vitale-Brovarone and Verne, 2005). Compared to uncoated implants, samples coated with BGs showed osteoconductivity (Martorana et al., 2009) and enhanced osseointegration (Ignatius et al., 2005). A functionally graded structure has been applied to minimize crack propagation at the interface between the coating and the substrate. The multilayer approach was developed to achieve a compliant material that could withstand the stresses due to the expansion coefficient mismatch between the substrate and the coatings, and maintain the bioactivity of the outer layers (Brovarone et al., 2001). Bioactive glasses with higher silica were used to form the first layer of a composite coating, providing a strong bond with the metal substrates, whereas BGs with lower silica were used to form the outer layer to enhance bioactivity (Rahaman et al., 2008; Lopez-Esteban et al., 2009; Verne et al., 2004). Compared to uncoated Ti alloy, the functionally graded coatings showed good cytocompatibility (Foppiano et al., 2006) and can indirectly induce an increase of expression of Runx-2, a key marker of osteoblast differentiation (Foppiano et al., 2007). The employed intermediate layer between the metal and BGs can also avoid the contamination of the diffused metallic ions. Verne et al. found that an intermediate layer based on SiO2-CaO, which formed undesired additional phases with the metal substrate, was necessary to avoid the diffusion of metal ions (Verne et al., 2005). Besides changing the composition of BGs, other methods are also used to improve the bioactivity of the outer layers. MBGs have shown higher bioactivity than the conventional sol-gel BGs due to the higher specific surface area (Xia and Chang, 2006). A thin film of mesoporous silica was deposited on Ti6Al4V substrate coated with glasses, which induced apatite formation in SBF after 7 days (Gomez-Vega et al., 2001). In brief, it is possible to fabricate a layer of BGs onto bioinert implants with an adequate thermal expansion coefficient, good adherence ability and bioactivity using the enameling method by tailoring the composition of BGs and enameling processing conditions.

2.4.2 Plasma-sprayed deposition Plasma-spraying is a feasible way to coat metallic substrates with glass powders (Bolelli et al., 2007a). The high-temperature plasma (up to 10 000 to 30 000K) melts glass particles into droplets. Then the droplets are ejected at high velocity and sprayed on the substrate with rapid solidification (Kang et al., 2007). The advantages of the plasma-sprayed technique are in creating reasonably high coating bond strength and mechanical properties (Ding et al., 2001). Compared to the enameling technique, BG coating prepared by plasmaspraying has higher superficial mechanical strength. Schrooten et al. used reactive plasma spraying method to deposit the BG layer on a Ti6Al4V rod. This technique proved to produce a high-quality coating with the adhesion strength of 40.1 ± 4.8 Mpa in shear and 69.4 ± 8.4 MPa in tension (Schrooten et al., 1999). The BG



41

coated metal implants showed good biocompatibility in vitro and in vivo. Gabbi et al. prepared amorphous BG-coated titanium and its alloys by plasma-spraying. Both in vitro and in vivo studies confirmed that BG coating was osteoconductive (Gabbi et al., 1995). In order to increase the mechanical properties of the coating, Goller et al. investigated the effect of bond coating layer, which can provide a good thermal expansion match between the substrate and the BG coating. In that study, they used a alumina-titania (60%Al2O3, 40%TiO2) bond coating and found this kind of coating showed better mechanical strength than the pure BG coating (Goller et al., 2003). In general, implants coated by BGs using plasma spraying showed superior properties such as bioactivity and long-term stability, and is generally accepted for orthopedic applications. However, the plasma spraying is a beeline process, so it is challenging to use it for coating complex substrates and to deposit graded coating (Ma et al., 2006).

2.4.3 Electrophoretic deposition (EPD) technique Electrophoretic deposition (EPD) is an electrochemical method which is usually carried out in a two electrode cells. When direct current (DC) electric field is applied, charged particles suspended in a suitable liquid move toward the oppositely charged electrode, and then the particles accumulate at the deposition electrode and create a relatively compact and homogeneous film. Therefore, EPD can be applied to any solid that is available as a fine powder (e.g. 100 µm. These are the preferred conditions for vascularization, cell penetration and tissue in growth [5, 7].

5.3

Bioactive glasses and glass-ceramics in bone tissue engineering

The most widely applied bioactive glasses consist of a silicate network incorporating sodium, calcium and phosphorus in different relative proportions in the SiO2-Na2O-CaO-P2O5 system [8]. The original bioactive glass composition universally known as 45S5 Bioglass® (in wt%: 45% SiO2, 24.5% Na2O, 24.5% CaO and 6% P2O5) has received approval from the US Food and Drug Administration (FDA), and it has found applications in clinical treatments of periodontal diseases as bone filler as well as in middle ear surgery [10]. This glass has also been the subject of intensive research for the development of BTE scaffolds [11]. Other silicate compositions contain no sodium or have additional elements in the silicate network such as fluorine, magnesium, strontium, iron, silver or zinc [12–16]. In addition, a range of silicate glass-ceramics is being investigated to fabricate BTE scaffolds [17–19]. Borosilicate and borate glasses are also receiving increasing attention for the fabrication of BTE [20, 21], as discussed further below. The typical feature common to all bioactive glasses is their positive interaction with living tissues, in particular bone tissue, which is termed ‘bioactivity’ [10, 22,


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

110 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

23]. Bonding to bone is established by the precipitation of a calcium-deficient, carbonated apatite surface layer on the bioactive glass surface when in contact with relevant physiological fluids [10]. When bioactive glasses are applied in vivo, apatite crystals precipitate in the vicinity of collagen fibrils inducing enhanced bone cell attachment and strong bonding to bone. Both micron-sized and nanoscale bioactive glass particles are considered in BTE [24, 25], and also the fabrication of composite materials, e.g. a combination of biodegradable polymers and bioactive glass [26] is being exploited, as discussed further below. There are several other reasons to select bioactive glasses for BTE scaffolds, which are related to the specific effect of dissolution products of bioactive glasses on cellular behavior [27, 28]. For example, it has been shown that dissolution products from 45S5 Bioglass® upregulate the expression of genes that control osteogenesis [27, 29, 30], leading thus to higher rate of bone formation in comparison to hydroxyapatite [31]. There is also increasing evidence that bioactive glass particles added to a biopolymer construct can enhance the angiogenic potential of the scaffold, i.e. increase the secretion of vascular endothelial growth factor (VEGF) in vitro and the enhancement of vascularization in vivo [32–34]. These studies suggest that Bioglass® BTE scaffolds might stimulate neo-vascularization [34, 35], which is required for the development of large tissue-engineered constructs. Bioactive glasses can also serve as platform for the local delivery of selected ions, which can act to control specific cell functions, for example Co addition to suppress cell hypoxia [36]. This is in addition to the well-known development of bioactive glasses with antibacterial properties, for example incorporating Ag [37, 38], and as a delivery platform for other therapeutic ions for the development of multifunctional scaffolds [39]. The range of bioactive glass compositions and morphologies exhibiting these attractive properties for BTE has increased over the years, with the development of new synthesis and fabrication methods. This includes the development of sol-gel based techniques [40] as well as 3D manufacturing methods for bioactive glass and glass-ceramic scaffolds [11, 17–19], the production of bioactive glass nanofibres [41] and nanoparticles [24, 25, 42] as well as a great variety of composites combining bioactive glass and biopolymers [5, 26].

5.4

Bioactive glass-based scaffolds: fabrication technologies

The use of bioactive glasses for BTE scaffold fabrication requires manufacturing complicated 3D porous shapes and at the same time retention of the specific degree of bioactivity. Bioactive glass and glass-ceramic scaffolds exhibiting highly porous structure are being fabricated by a variety of techniques described in detail elsewhere [4]. The fabrication methods can be broadly divided into two groups; (1) methods relying on glass powders and (2) sol-gel approaches. Only powder-based methods will be considered in this chapter.


Bioactive glass and glass-ceramic scaffolds

111

The foam replica technique is one of the most popular methods being applied to fabricate BTE scaffolds from glass powders. It was used for Bioglass® for the first time in 2006 [11] and it is currently one of the methods of choice for fabricating BTE scaffolds leading to foam-like structures of high porosity (>90%) and high pore interconnectivity [43]. The technique involves the use of polymeric sponges as sacrificial templates to prepare cellular bioactive glass (and glass-ceramic) structures of various pore sizes, porosities and chemical compositions. The sacrificial template, e.g. a polyurethane foam, is initially soaked in a glass powder suspension until the foam struts are homogeneously coated with a high concentration of particles. Binders are usually added to the initial suspension in order to prevent cracking of the struts during drying and upon the subsequent heat treatment. In the following step, the polymer template is burnt out by a controlled heat treatment and the glass or glass-ceramic structure is finally densified by sintering at high temperatures. If the fabrication of the scaffold proceeds from a bioactive glass powder, such as in the foam replica method [11], then the viscosity–temperature and crystallization characteristics of the glass must be known because they define the sintering conditions. Depending on the polymer sponge used as template, pore sizes between 200 µm and 3 mm can be obtained. Numerous types of bioactive glass-ceramic [11, 17–19] foams have been produced by the replica method using polyurethane sponges as templates. Figure 5.2a shows the macroscopic pore structure of a Bioglass®-based glass-ceramic scaffold (SEM image) fabricated by this technique [11] whilst Fig. 5.2b is a higher magnification SEM image showing the crosssection of the scaffold struts. Table 5.1 includes a summary of bioactive glass and glass-ceramic scaffolds that have been fabricated by the foam replica technique indicating also the mechanical properties achieved and other typical scaffold characteristics investigated. Another method developed to produce porous scaffolds from bioactive glass powder is the sacrificial porogen technique. This method involves the preparation of a composite comprising a sacrificial phase mixed with glass particles. The sacrificial phase is extracted (usually thermally) from the partially consolidated matrix to generate pores within the microstructure. The mechanical strength of structures made by the sacrificial template method is usually higher than that of scaffolds fabricated by the replica method. However porosity and pore interconnectivity are substantially lower than in scaffolds made by the replica technique. Another advantage of the foam replica method is the possibility of developing scaffolds with graded or layered porosity [47, 48]. These structures are created by modifying the shape of the starting polyurethane sponges as discussed elsewhere [47]. To improve the mechanical properties of highly porous scaffolds made from bioactive glasses, crystallization of the scaffold struts, effectively developing glass-ceramics, is a suitable approach [17–19]. In the case of porous scaffolds extensive sintering of the bioactive glass is required in order to densify the struts


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

112 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

5.2 SEM images showing the pore structure of a Bioglass® scaffold fabricated by the foam replica method (a) and the microstructure of the strut cross-section (b).



81.5

85

48

68

54

Calcium silicate (CaSiO3, WT)

Glass 13–93, Composition: 53SiO2-6Na2O-12K2O-5MgCaO-4P2O5 wt%

Sol-gel bioglass powder. Stoichiometric ratio SiO2:CaO:P2O5 = 70:26:4

Glass: 50SiO2–22.6CaO–5.9 Na2–4P2O5–12K2–5.3MgO– 0.2B2O3 (wt%)

Molar composition: 45% SiO2, 3% P2O5, 26% CaO, 7% MgO, 15% Na2O, 4% K2O

89–90

45S5 Bioglass®

100–500

240

50

100–500

300–500

510–720

Porosity (%) Pore size (µm)

Composition

0.4*

0.40

–

11

0.3

0.27–0.42

Compressive strength (MPa)

Reference

Apatite formation after 28 d in SBF * Scaffolds pre-treated in SBF – Study proliferation of human osteoblasts

Apatite formation after 4 d in SBF

Apatite formation after 1 week in SBF

Apatite formation after 7d in SBF MC3T3-E1 attachment and proliferation

Apatite formation after 14 d in SBF HBDC attachment-viability

[45] *[46]

[44]

[19]

[18]

[17]

Apatite formation after [11] 28 d in SBF osteoblastlike cells (MG 63) attachment-proliferation*

Study of bioactivity SBF or cell culture

Table 5.1 Different silicate scaffolds fabricated by the foam replica technique, and their principal characteristics

Ca3Mg(SiO4)2 and Ca2MgSi2O9

Amorphous

Quartz (SiO2), wollastonite (CaSiO3), apatite

Amorphous

α-CaSiO3

Na2Ca2Si3O9

Crystalline phases

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

114 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

and create a structure able to withstand mechanical loading during manipulation by surgeons and during application in vivo. In most cases reported, silicate bioactive glasses crystallize to wollastonite, quartz, apatite, calcium silicate, sodium calcium silicate or combeite crystalline phases (or combinations of these phases). It has been also confirmed recently that glass crystallization, for example in Bioglass®-derived glass-ceramic scaffolds, does not suppress bioactivity, it only retards the formation of the surface hydroxyapatite layer when the scaffold is immersed in body fluid [11, 49]. In this regard, a very important characteristic of Bioglass®-derived glassceramics is the transformation of the Na2Ca2Si2O9 crystalline phase to an amorphous matrix containing hydroxyapatite crystallites after immersion in simulated body fluid for 28 days [11]. The kinetic of HA formation can be tailored through the fabrication process by changing the sintering conditions. Fine crystals of Na2Ca2Si2O9 grow and almost complete densification of the scaffold struts occur when scaffolds are sintered at temperatures >1000°C. These conditions confer the scaffold the best possible compressive strength for the given porosity. It should therefore be emphasized that the final goal in the design of BTE scaffolds is attaining adequate structural integrity to warrant mechanical support of the surrounding tissue whilst maintaining bioactive behavior. Then, in the later stages of the tissue regeneration process the scaffold should biodegrade at a set rate matched to the rate of formation of new bone tissue, as mentioned above. Bioactive glass and glass-ceramic scaffolds fabricated by the replica technique (Table 5.1) are able to provide this function and they remain the subject of extensive research efforts considering their promising properties in the context of BTE. In the following sections two exemplary cases will be considered to illustrate the development of BTE scaffolds based on bioactive glasses and their fabrication by the foam replica manufacturing technique: (1) bioactive glass scaffolds made from a novel boron-containing glass, and (2) polymer-coated 45S5 Bioglass® based scaffolds.

5.5

Scaffolds from boron-containing bioactive glass

5.5.1 Fabricating bioactive glass scaffolds with different compositions Many parameters are involved in the selection of the bioactive glass for BTE scaffolds, which are dictated by the target application and the processing characteristics chosen for the scaffold fabrication. The main factors to be considered in relation to the scaffold requirements discussed in Section 5.2 are: the extent of bioactive response (bioreactivity), the kinetics of biodegradability and osteoconductivity as well as appropriate mechanical strength and structural integrity over a period of time [9]. In addition, in the case of silicate systems, the chemical composition will have an effect on the viscosity–temperature relationship



115

and the crystallization potential of bioactive glasses, which are two important considerations in relation to the ability to process scaffolds of complex porous morphology from glass powders. Finding optimal chemical compositions of bioactive glasses for BTE scaffolds requires knowledge of the effect of addition or suppression of different oxides, in relation to the basic 45S5 Bioglass® composition, on relevant properties [50, 51]. It is well known that the properties of silicate glasses can be adjusted in a more or less controlled fashion by changes in chemical composition [52]. For conventional glasses, the property and composition optimization is supported by well-established relationships, and these are also relevant for bioactive glasses [50, 51, 53]. The composition also affects the processing characteristics of scaffolds considering that scaffold fabrication demands a wide working range and, preferably, a low tendency to devitrify. It has been discussed that bioactive glass compositions intended for sintered porous bodies should have preferably devitrification temperatures above 800°C and they should be chosen within the calcium silicate (e.g. wollastonite) primary crystalline phase, as these compositions can better sustain higher temperatures without crystallizing [50]. Generally, the working range of glasses increases with rising silica content and with a simultaneous decrease in in vitro bioactivity. The bioactivity has been found to depend more on the total amount of alkalis and alkaline earths than on the types of oxides, i.e. sodium and potassium, or calcium and magnesium, in each group. However it has been observed that with incorporation of potassium oxide, magnesium oxide or boron oxide, the forming properties of the glasses can be adjusted [53, 54]. In particular, the incorporation of boron oxide represents an interesting approach owing to the documented osteogenic and angiogenic effect of B2O3 containing bioactive glasses, which is attributed to the effect of the boron ion [21]. Boron-containing silicate glasses are therefore being increasingly considered to fabricate BTE scaffolds, as discussed next.

5.5.2 Boron-containing bioactive glass scaffolds The incorporation of boron in silicate bioactive glasses has been reported to lead to enhanced bone formation [21]. In addition, silica-free borate glasses have been shown to possess low chemical durability, and to convert rapidly to calcium phosphate (or hydroxyapatite) in physiological media [55] and to bond directly to bone in a manner comparable to the silicate-based 45S5 Bioglass®. Boron containing silicate glasses are being currently investigated considering the reported positive effect of the B ion but also based on the fact that these glasses can be heat treated (sintered) at lower temperatures than 45S5 Bioglass® and do not crystallize [20, 44, 56]. The borosilicate bioactive glass ‘code 0106’ with nominal composition (in wt%): 50 SiO2, 22.6 CaO, 5.9 Na2O, 4 P2O5, 12 K2O, 5.3 MgO and 0.2 B2O3 developed at the Process Chemistry Centre, Åbo Akademi University, Turku


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

116 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Finland [57] has been recently used to fabricate BTE scaffolds by the foam replica method. Powder of this glass with average particle size of 36 µm was used. The heat treatment involved sintering for 3 to 10 hours at temperatures in the range 630 to 680°C (considerably lower than temperatures used for 45S5 Bioglass® scaffolds mentioned above) [44]. The heat treatment leading to scaffolds with the most promising characteristics involved sintering at 670°C for 5 hours. The morphology of the scaffolds is shown in Fig. 5.3. The scaffold struts were seen to be well sintered, which should lead to favorable mechanical properties. The average pore size of the fabricated scaffolds was 240 µm, while the average porosity was 68%. This porosity value is lower than that of scaffolds fabricated previously using Bioglass® [49], which exhibited porosity of up to 90%. However the pore size is within the desired range (>200 µm) and the pore structure is highly interconnected (Fig. 5.3). In vitro bioactivity studies in simulated body fluid (SBF) showed that a carbonate hydroxyapatite (HCAp) layer was deposited on scaffolds after only 4 days of immersion in SBF, demonstrating the high in vitro bioactivity, which is comparable to that of 45S5 Bioglass®. The FTIR characterization results shown in Fig. 5.4 confirm that scaffolds exhibit marked bioactive behavior given by the significant and rapid formation of the HCAp surface layer [44]. Figure 5.5 shows SEM images illustrating the transformation of the microstructure of samples during the formation of HCAp. After one day’s immersion in SBF the onset of a degradation process is clearly visible and the development of a uniform, thick HCAp layer on the surface of the scaffolds is confirmed after only four days of immersion in SBF. X-Ray diffraction analysis showed that as-sintered scaffolds consisted of an amorphous structure. The fabrication of bioactive scaffolds from bioactive glasses in powder form, to retain an amorphous structure, is quite unusual as most known bioactive glasses crystallize during the high-temperature fabrication process, as mentioned in the

5.3 SEM images showing the morphology of borosilicate bioactive glass scaffolds at low (left) and high (right) magnifications (sintering temperature and time: 670°C, 5 hours, respectively) [44] (reproduced with permission of IOP).



117

5.4 FTIR spectra of borosilicate bioactive BTE scaffolds before and after immersion in SBF for different soaking times [44]. The transmittance spectrum of HAp is also presented for comparison (reproduced with permission of IOP).


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

118 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

5.5 SEM images showing the surface of borosilicate bioactive scaffolds after one and four days of immersion in SBF (modified from [44], reproduced with permission of IOP).

previous section. An amorphous phase is less stable than its crystalline counterpart in terms of Gibbs free energy, which is expected to result in a better bioactive behavior in comparison to partially crystallized scaffolds. The degradation rate of the boron-containing scaffolds in relevant in vitro or in vivo conditions has not been investigated to date, and it remains a key study to confirm the positive effect



119

of boron (released during scaffold degradation) on both osteogenic and angiogenic potential predicted by other authors [20, 21].

5.6

Polymer-coated composite scaffolds

5.6.1 Advantages of composite systems The major disadvantage of bioactive glasses is their low fracture toughness and brittleness. For applications such as BTE, bioactive glasses are often used in combination with biodegradable polymers to achieve the best possible mechanical and biological performance [5, 6]. In the development of composites for tissue engineering scaffolds, two main approaches are being followed; the first approach considers the incorporation of bioactive glass particles as inclusions into polymer structures, e.g. foams [5]; and the second approach considers the incorporation of polymer coatings onto a 3D porous bioactive glass based scaffold [58]. In this section specific examples of the second approach will be discussed while composite scaffolds based on biopolymers containing bioactive glass particulate inclusions are presented in Chapter 7. The introduction of a polymer into the structure of bioactive glass scaffolds, whether as a coating or by infiltrating into the strut structure, has been investigated by several authors [58, 59]. It is hypothesized that by introducing a polymer into the structure of struts, the polymer can fill micro-cracks, bridge large cracks and form a polymer-bioactive glass composite network, behaving in a similar way to collagen fibers in bone. It is well established that collagen fibers bridge cracks in bone, to increase bone fracture toughness [62, 63]. If the polymer just forms a layer that rests on top of the scaffold strut and does not infiltrate the material structure then the polymer will not improve the mechanical properties as well as if the polymer infiltrated the micro-cracks and remaining pores of the scaffold struts. This behavior is illustrated schematically in Fig. 5.6a and 5.6b. Processes developed to fabricate both polymer coated bioactive glass scaffolds and polymer-bioactive glass scaffolds with interpenetrating network microstructure are based on infiltrating a sintered (or partially sintered) bioactive glass or glassceramic scaffold with the biodegradable polymer in solution [58]. A novel method recently developed to coat 3D scaffolds with polymers is Matrix Assisted Pulsed Laser Evaporation (MAPLE) [64]. This technique is often preferred over other film deposition methods since it provides high control over film characteristics. It was shown that this technique can be used to produce PDLLA (poly(D,L lactide))coated Bioglass® scaffolds [65]. The alternative approach of fabricating hybrid polymer–ceramic composite scaffolds, e.g. exploiting the molecular mixing of inorganic and organic phases for example in sol-gel based techniques, has also been explored [66], however for the sake of brevity, these hybrid materials will not be considered in the present chapter.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

120 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

5.6 Schematic diagram illustrating how a polymer lies on the surface of a cracked (or partially sintered) scaffold strut (coating) (a); or how it can infiltrate the micro-cracks on the surface of the struts (b).

Table 5.2 Compressive strength values of a variety of polymer-coated bioactive glass composites and relevant bone tissues Inorganic phase

Polymer coating

Type

Compressive strength (MPa)

Reference

45S5 Bioglass® 45S5 Bioglass® 45S5 Bioglass®

No coating P(3HB) P(3HB) microspheres PDLLA – –

Scaffold Scaffold Scaffold

0.1–0.3 ~1.5 ~0.4

[11] [67] [68]

Scaffold Bone Bone

0.2–0.65 4–12 130–180

[59] [69] [69]

45S5 Bioglass® Cancellous bone Cortical bone

Chen et al. [59] developed Bioglass®-based scaffolds coated with PDLLA for the first time. In those polymer-coated scaffolds bioactivity was not impaired as demonstrated by tests in simulated body fluid. Polyhydroxyalkanoate (P(3HB) has been investigated as an alternative coating material for tissue-engineering scaffolds [67], which leads to a significant improvement of the work of fracture in compression, as discussed below. Although polymer/bioactive glass composite scaffolds developed so far have not quite reached the mechanical requirements of the surrounding host tissue, as shown in Table 5.2, they are much closer to the desired mechanical properties than uncoated scaffolds. Polymers have the added advantage that they bind the structure of the scaffold together making the scaffold tougher and providing it with extra mechanical stability during the in vivo tissue regeneration process. In addition, the polymer can have other functions in the scaffold, such as being a carrier for therapeutic drugs, growth factors and even antibacterial metal ions; this should increase the functionality of the scaffolds [70]. There are several polymer–bioceramic composites that demonstrate the evolution of this hypothesis and are reviewed in the literature [58]. In the following section, BTE scaffolds in



121

the system poly(3-hydroxybutyrate) coated Bioglass® will be discussed as a typical example of this type of composite.

5.6.2 Bioglass® scaffolds coated with poly(3-hydroxybutyrate) Poly(3-hydroxybutyrate) (P(3HB)), a member of the polyhydroxyalkanoate family, is being considered for the development of composites for tissue engineering [71]. This polymer is naturally thermoplastic and is produced by many types of micro-organisms [72]. Bretcanu et al. [73] first used this bacteria derived polymer to infiltrate 45S5 Bioglass®-derived scaffolds. P(3HB), when used as a thin coating material, does not affect the interconnectivity of the scaffold porous structure and coated scaffolds were seen to maintain a high porosity of 85% [67, 73]. It was also found that when applied as a coating on a 3D porous Bioglass® scaffold, P(3HB) does not form a fully homogeneous layer on the struts, as shown in Fig. 5.7. However, this lack of coating connectivity is useful for the scaffold to retain a high bioactivity because the underlying Bioglass® surface will be in direct contact with the biological environment promoting the occurrence of the typical surface bioreactions of bioactive glasses [10]. When the compressive strength of P(3HB)-coated Bioglass® scaffolds is compared to that of uncoated Bioglass®-derived scaffolds, there is a considerable improvement; from 0.74MPa to ∼1.5MPa [67], which makes up for the reduction in scaffold porosity. This composite is comparable in terms of mechanical properties (and again a vast improvement on the plain biomaterial), to other polymer-coated bioactive glass composites described above, as illustrated in Table 5.2. There is increasing interest in using these scaffolds to produce drug delivery devices, e.g. incorporating growth factors, antibiotics and other additives into the polymer coating to improve the scaffold multi-functionality.

5.7 SEM images showing the microstructure of a 45S5 Bioglass®/ P(3HB) composite scaffold (a); and the discontinuous nature of the P(3HB) coating on the scaffold (b) [73].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

122 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Another method of incorporating the polymer onto a scaffold is in the form of microspheres, as discussed by Francis and Meng et al. [68]. This approach has two effects on the overall performance of the composite scaffold, the first being increased functionality of the scaffold as the microspheres can be filled with antibiotics or growth factors, and secondly it should improve the properties of the scaffold in the same manner as the coating with P(3HB) films discussed above. The microspheres are formed using the solid-in-oil-water emulsion technique, which produces microspheres with a mean diameter ranging from 1.5 to 2.0 µm. Pre-formed 45S5 Bioglass® scaffolds produced using the foam replication technique as described above are coated with P(3HB) microspheres by pipetting drops of a slurry containing the microspheres onto the scaffolds. It should be noted that the incorporation of the additive (the antibiotic Gentamicin in the case reported by Francis and Meng et al. [68]) into the microspheres occurs before the microspheres are used to coat the scaffolds. Figure 5.8 indicates that the microsphere coating of the scaffolds is extensive but not completely homogeneous: there are some gaps in the layer of microspheres, and this will help the composite scaffold maintain its high level of bioactivity because strut surfaces will be directly exposed to the biological environment. Similarly, the addition of the polymer coating (in the form of microspheres) was seen to reduce the overall porosity of the scaffold from ∼90% down to ∼70%; this reduction in porosity has a positive effect on the compressive strength

5.8 SEM image showing the microstructure of a 45S5 Bioglass® scaffold coated with P(3HB) microspheres [68].



123

of the scaffolds, which doubles compared to uncoated scaffolds [68]. The addition of the microspheres also increased the work of fracture of the composite scaffolds, i.e. they are more resistant to crack propagation compared to uncoated scaffolds [68]. It was also demonstrated that addition of P(3HB) microspheres did not inhibit the formation of hydroxyapatite on the surface of the scaffold in contact with SBF, which is important as this surface bioreaction is the main function provided by the 45S5 Bioglass® scaffold to ensure bone cell attachment [11]. Consequently, it can be concluded that there is not much difference between the two different applied techniques of incorporating P(3HB) into Bioglass®-based scaffolds, as the two methods, namely film and microsphere coating, lead to scaffolds with increased compressive strength, and in both cases multifunctional BTE scaffolds incorporating P(3HB) that retain the basic 45S5 Bioglass® bioactivity can be fabricated.

5.7

Conclusions

An overview on the development of BTE scaffolds based on (melt-derived) bioactive glasses and glass-ceramics has been provided in this chapter highlighting the manufacturing of 3D scaffolds by the foam replica technique. The increasing interest in the application of bioactive glasses and glass-ceramics for fabricating BTE scaffolds is based on the attractive biological properties of bioactive glasses, which include not only the classical concept of bioactivity based on the bioreactivity of the material in contact with relevant fluids, but also the recognized effect of ionic glass dissolution products on osteogeneis and angiogenesis. The utilization of bioactive glasses as platforms for the controlled release of certain metallic ions, beyond Si, Ca, and P, to stimulate specific cellular responses can be highlighted as an important research area likely to concentrate research efforts in the near future. In this chapter, two specific types of scaffolds were discussed in some detail: (1) BTE scaffolds fabricated from boron-containing bioactive glass, which leads to non-crystalline scaffolds of high bioactivity with the potential advantage of the release of boron to promote bone formation; and (2) polymer-coated scaffolds of enhanced mechanical behavior and adequate bioactivity as basic structures for the future development of multifunctional scaffolds with drug delivery capability. In this context, a significant toughening effect by polymer incorporation, especially in scaffolds exhibiting interpenetrating network microstructure, represents an important improvement in scaffold design. The addition of a polymer phase might have extra functions, e.g. the biodegradable polymer can act as carrier for biomolecules, growth factors and antibiotics, hence increasing the capability of the BTE constructs. The optimization of scaffolds for BTE strategies is still one of the key tasks to be accomplished in order to realize the promise that tissue engineering holds to


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

124 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

improve the life of millions of humans suffering bone and muscoloskeletal diseases. Bioactive glasses, and the possibility of new compositions incorporating specific therapeutic ions and their combination with biopolymers to form composites, will continue to play a significant role in the field as one of the materials of choice for scaffold developments.

5.8

References

1. Hutmacher D. W., Schantz J. T., Lam C. X. F., Tan K. C., Lim T. C., State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective. J Tissue Eng Regen Med 2007; 1:245–260. 2. Ikada Y., Challenges in tissue engineering. J R Soc Interface, 2006; 3:589–601. 3. Moroni L., De Wijn J. R., Van Blitterswik C. A., The integrating novel technologies to fabricate smart scaffolds. J BioMater Sci Polymer Edition 2008; 19:543–572. 4. Jones J. R., Boccaccini A. R., Cellular Ceramics in Biomedical Applications: Tissue Engineering, Chapter 5.8 in P. Colombo and M. Scheffler (eds) Cellular Ceramics. Structure, Manufacturing, Properties and Applications, Wiley-VCH, Weinheim 2005, pp. 547–570 (ISBN: 3-527-31320-6). 5. Rezwan K., Chen Q. Z., Blaker J. J., Boccaccini A. R., Biodegrdable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 2006; 27:3413–3431. 6. Guarino V., Causa F., Ambrosio L., Bioactive scaffolds for bone and ligament tissue, Expert Rev Medical Devices 2007; 4 (3):405–418. 7. Karageorgiou V., Kaplan D., Porosity of 3D biomaterial scaffolds and osteogenesis. Biomaterials 2006; 26:5474–5491. 8. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K., Bonding mechanism at the interface of ceramic prosthetic materials. J Biomed Mater Res 1971; 2:117–127. 9. Hutmacher D. W., Polymeric scaffolds in tissue engineering bone and cartilage. Biomaterials 2000; 21:2529–2543. 10. Hench L. L., Bioceramics. J Amer Ceram Soc 1998; 81:1705–1728. 11. Chen Q. Z., Thompson I. D., Boccaccini A. R. 45S5 Bioglass®-derived glass-ceramic scaffolds for bone tissue engineering. Biomaterials 2006; 27:2414–2425. 12. Vallet-Regi M., Salinas A. J., Roman J., Gil M., Effect of magnesium content on the in vitro bioactivity of CaO-MgO-SiO2-P2O5 sol-gel glasses. J Mater Chem 1999; 9 (2):515–518. 13. Lao J., Nedelec J. M., Jallot E., New strontium-based bioactive glasses: physicochemical reactivity and delivering capability of biologically active dissolution products. J Mater Chem 2009; 19:2940–2949. 14. Haimi S., Gorianc G., Moimas L., Lindroos B., Huthala H. et al., Characterization of zinc-releasing three-dimensional bioactive glass scaffolds and their effect on human adipose stem cell proliferation and osteogenic differentiation, Acta Biomater 2009; 5:3122–3131. 15. Cannillo V., Sola A., Potassium-based composition for a bioactive glass. Ceram Int 2009; 35:3389–3393. 16. Zreiqat H., Ramaswany Y., Wu C., Paschalidis A., Lu Z., James B., Birke O., McDonald M., Little D., Dunstan C. R., The incorporation of strontium and zinc into a calcium-silicon ceramic for bone tissue engineering. Biomaterials 2010, in press.



125

17. Wu C., Ramaswamy Y., Boughton P. and Zreiqat H., Improvement of mechanical and biological properties of porous CaSiO3 scaffolds by poly(D,L-lactic acid) modification Acta Biomaterialia 2008; 4:343–353. 18. Fu Q., Rahaman M. N., Bal B. S., Brown R. F. and Day D. E., Mechanical and in vitro performance of 13-93 bioactive glass scaffolds prepared by a polymer foam replication technique. Acta Biomaterialia 2008; 4:1854–1964. 19. Rainer A., Giannitelli S. M., Aqbbruzzese F., Traversa E., Licoccia S., Trombetta M., Fabrication of bioactive glass-ceramic foams mimicking human bone portions for regenerative medicine. Acta Biomaterialia 2008; 4:362–369. 20. Liu X., Huang W., Fu H., Yao A., Wang D., Pan H. and Lu W. W., Bioactive borosilicate glass scaffolds: improvement on the strength of glass-based scaffolds for tissue engineering. J Mater Sci Mater Med 2009; 20:365–372. 21. Gorustovich A. A., Porto Lopez J. M., Guglielmotti M. B., Cabrini R. L., Biological performance of boron-modified bioactive glass particles implanted in rat tibia bone marrow. Biomed Mater 2006; 1:100–105. 22. Livingstone T., Ducheyne P., Garino J., In vivo evaluation of a bioactive scaffold for bone tissue engineering. J Biomed Mater Res 2002; 62:1–13. 23. Ylänen H., Karlsson K. H., Itälä A., Aro H. T., Effect of immersion in SBF on porous bioactive bodies made by sintering bioactive glass microspheres. J Non-Cryst Sol 2000; 275:107–115. 24. Misra S. K., Mohn D., Brunner T. J., Stark W. J., Philip S. E., Roy I., Salih V., Knowles J. C., Boccaccini A. R., Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB)/Bioglass® composites. Biomaterials 2008; 29(12): 1750–1761. 25. Hong Z., Reis R. L., Mano J. F. Preparation and in vitro characterization of novel bioactive glass-ceramic nanoparticles. J Biomed Mater Res A 2009; 88 (2):304–313. 26. Boccaccini A. R., Maquet V., Bioresorbable and Bioactive Polymer/Bioglass® Composites with tailored pore structure for tissue engineering applications. Comp Sci Technol 2003; 63:2417–2429. 27. Xynos I. D., Edgar A. J., Buttery L. D. K., Hench L. L., Polak J. M., Gene expression profiling of human osteoblasts following treatment with the ionic products of Bioglass® 45S5 dissolution. J Biomed Mater Res 2001; 55:151–159. 28. Hench L. L., Thompson I. D., Twenty-first century challenges for biomaterials. J Royal Soc Interface 2010, in press. 29. Xynos I. D., Hukkanen M. V. J., Batten J. J., Buttery L. D., Hench L. L., Polak J. M. Bioglass® 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: Implications and applications for bone tissue engineering. Calcified Tissue Int 2000; 67 (4):321–329. 30. Jell G., Stevens M. M., Gene activation by bioactive glasses. J Mater Sci Mater Med 2006; 17:997–1002. 31. Wheeler D. L., Montfort M. J., McLoughlin S. W., J Biomed Mater Res 2001; 55:603–610. 32. Day R. M., Boccaccini A. R., Shurey S., Roether J. A., Forbes A., Hench L. L., Gabe S. M., Assessment of polyglycolic acid mesh and bioactive glass for soft-tissue engineering scaffolds. Biomaterials 2004; 25:5857–5866. 33. Leu A., Leach J. K., Proangiogenic potential of a collagen/bioactive glass substrate. Pharm Res 2008; 25:1222–1229. 34. Gorustovich A. A., Roether J. A., Boccaccini A. R., Effect of bioactive glasses on angiogenesis: In vitro and in vivo evidence. A review. Tissue Eng Part B 2010; 16:199– 207.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

126 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

35. Deb S., Mandegaran R., Di Silvio L., A porous scaffold for bone tissue engineering/45S5 Bioglass® derived porous scaffolds for co-culturing osteoblasts and endothelial cells. J Mater Sci Mater Med 2010; 21:893–905. 36. Azevedo M., Jell G., Hill R., Stevens M. M., Novel hypoxia mimicking bioactive materials for tissue engineering. Tissue Eng A 2008; 14:889–889 (Meeting Abstract: P288). 37. Bellantone M., Williams H. D., Hench L. L., Broad-spectrum bactericidal activity of Ag2O-doped bioactive glass. Antimicrob Agents Chemother 2002; 46:1940–1945. 38. Vitale-Brovarone C., Miola M., Balagna C., Verne E., 3D-glass-ceramic scaffolds with antibacterial properties for bone grafting. Chemical Eng J 2008; 137:129–136. 39. Mouriño V., Boccaccini A. R., Bone tissue engineering therapeutics: controlled drug delivery in three-dimensional scaffolds. J R Soc Interface 2010; 7:209–227. 40. Li R., Clark A. E., Hench L. L., An investigation of bioactive glass powders by sol-gel processing. J App Biomater 1991; 2 (4):231–239. 41. Quintero F., Pou J., Comesana R., Lusquinos F., Riveiro A., Mann A. B., Hill R. G., Wu Z. Y., Jones J. R., Laser spinning of bioactive glass nanofibers. Adv Funct Mater 2009; 19:1–7. 42. Brunner T. J., Grass R. N., Stark W. J., Glass and bioglass nanopowders by flame synthesis. Chem Commun 2006; 13:1384–1386.74. 43. Zenati R., Fantozzi G., Chevallier J., Mourad A., Porous Bioglass and Preparation Thereof. French Patent no FR2005/001921, 2005. 44. Mantsos, T., Chatzistavrou X., Roether J. A., Hupa L., Arstila H., Boccaccini A. R., Non-crystalline composite tissue engineering scaffolds using boron-containing bioactive glass and poly(D,L-lactic acid) coatings. Biomed Mater 2009; 4:055002 (12 pp). 45. Renghini C., Komlev V., Fiori F., Verne E., Baino F., Micro-CT studies on 3-D bioactive glass-ceramic scaffolds for bone regeneration. Acta Biomater 2009; 5: 1328–1337. 46. Vitale-Brovarone C., Verne E., Robiglio L., Appendino P., Bassi F., Martinasso G., Muzio G., Canuto, R. Development of glass-ceramic scaffolds for bone tissue engineering: characterization, proliferation of human osteoblasts and nodule formation. Acta Biomater 2007; 3:199–208. 47. Bretcanu O., Samaille C., Boccaccini A. R., Simple methods to fabricate Bioglass®derived glass-ceramic scaffolds exhibiting porosity gradient, J Mater Sci 2008; 43:4127–4134. 48. Vitale-Brovarone C., Baino F., Verne E., Feasibility and tailoring of bioactive glassceramic scaffolds with gradient of porosity for bone grafting. J Biomater Appl 2010; 24:693–712. 49. Boccaccini A. R., Chen Q. Z., Lefebvre L., Gremillard L., Chevalier J., Sintering, crystallisation and biodegradation behaviour of Bioglass®-derived glass-ceramics. Faraday Discuss 2007; 136:27–44. 50. Arstila H., Vedel E., Hupa L., Hupa M., Factors affecting crystallization of bioactive glasses. J Eur Ceram Soc 2007; 27:1543–1546. 51. Andersson Ö. H., Karlsson K. H., On the bioactivity of silicate glass. Journal of Non-Crystalline Solids 1991; 129: 145–151. 52. Westerund T., Hatakka L., Karlsson K. H., A model for optimizing. Glass batch compositions. J Amer Ceram Soc 1983; 66 (8):574–579. 53. Brink, M., The influence of alkali and alkaline earths on the working range for bioactive glasses. J Biomed Mater Res 1997; 36:109–117.



127

54. Itälä A., Koort J., Ylänen H. O., Hupa M., Aro H. T., Biologic significance of surface microroughening in bone incorporation of porous bioactive glass implants. J Bio Mater Res A 2003; 67 (2):496–503. 55. Day D. E., White J. E., Brown R. F., McMenamin K. D., Transformation of borate glasses into biologically useful materials. Glass Technol 2003; 44 (2):75–81. 56. Rahaman M. N., Day D. E., Brown R. F., Fu Q., Jung S. B., Nanostructured bioactive glass scaffolds for bone repair, Advances in Bioceramics and Porous Ceramics, ed. R. Narayan and P. Colombo (Westerville, OH: American Ceramic Society) 2008; pp. 211–222. 57. Vedel E., Zhang D., Arstila H., Hupa L., Hupa M., Predicting physical and chemical properties of bioactive glasses from chemical composition. Part IV: tailoring composition with desired properties. Glass Technol: Eur J Glass Sci Technol A 2009; 50:9–16. 58. Yunos M. D., Bretcanu O., Boccaccini A. R., Polymer–bioceramic composites for tissue engineering scaffolds. J Mater Sci 2008; 43:4433–4442. 59. Chen Q. Z., Boccaccini A. R., Poly(D,L-lactide) coated 45S5 Bioglass®based scaffolds: processing and characterisation. J Biomed Mater Res 2006; A 77: 445–452. 60. Miao X., Tan L. P., Tan L. S., Huang X., Porous calcium phosphate ceramics modified with PLGA-bioactive glass. Mater Sci Eng C 2007; 27:274–280. 61. Kim H. W., Knowles J. C., Kim H. E., Hydroxyapatite porous scaffold engineered with biological polymer hybrid coating for antibiotic Vancomycin release. J Mater Sci Mater Med 2005; 16:189–195. 62. Wang X., Bank R. A., Tekoppele J. M., Agrawal C. M., The role of collagen in determining bone mechanical properties. J Orthopaed Res 2001; 19:1021–1026. 63. Nalla R. K., Kinney J. H., Ritchie R. O., Effect of orientation on the in vitro fracture toughness of dentin: the role of toughening mechanisms. Biomaterials 2003; 24:3955–3968. 64. Califano V., Bloisi F., Vicari L. R. M., Bretcanu O. and Boccaccini A. R., Matrixassisted pulsed laser evaporation of poly(D,L-lactide) for biomedical applications: effect of near infrared radiation. J Biomed Opt 2008; 13:014028. 65. Califano V., Bloisi F., Vicari L. R. M., Yunos D. M., Chatzistavrou X., Boccaccini A. R., Matrix assisted pulsed laser evaporation (MAPLE) of Poly(D,L lactide) (PDLLA) on three-dimensional Bioglass® structures. Adv Eng Mater 2009; 11: 685–689. 66. Catauro M., Raucci M. G., De Gaetano F., Buri A., Marotta A., Ambrosio L., Sol-gel synthesis, structure and bioactivity of polycaprolanctone/CaO.SiO2 hybrid material. J Mater Sci Mater Med 2004; 15:991–995. 67. Bretcanu O., Misra S., Roy I., Renghini C., Fiori F., Boccaccini A. R., Salih V., In vitro biocompatibility of 45S5 Bioglass®-derived glass-ceramic scaffolds coated with poly(3-hydroxybutyrate). J Tissue Eng Regen Med 2009; 3:139–148. 68. Francis L., Meng D., Knowles J. C., Roy I., Boccaccini A. R., Multifunctional P(3HB) microsphere/45S5 Bioglass®-based composite scaffolds for bone tissue engineering. Acta Biomater 2010; 6:2773–2786. 69. Gibson L. J., Ashby M. F., Cellular solids: structure and properties. 2nd edn. Oxford: Pergamon 1999, pp. 429–452. 70. Baroli B., From natural bone graft to tissue engineering therapeutics: brainstorming on pharmaceutical formulative requirements and challenges. J Pharm Sciences 2009; 98:1317–1375.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

128 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

71. Misra S. K., Valappil S. P., Roy I., Boccaccini A. R., Polyhydroxyalkanoate (PHA)/ inorganic phase composites for tissue engineering applications. Biomacromolecules 2006; 7 (8):2249–2258. 72. Chen G. Q., Wu Q., The application of polyhydroxyalkanoates as tissue engineering materials. Biomaterials 2005; 26:6565–6571. 73. Bretcanu O., Chen Q. K., Misra S., Boccaccini A. R., Roy I., Verne E., Brovarone C. V., Biodegradable polymer coated 45S5 Bioglass-derived glass-ceramic scaffolds for bone tissue engineering. Glass technology-Europ. J Glass Sci Technol Part A 2007; 48 (5):227–234.


6 Nanoscaled bioactive glass particles and nanofibres M. EROL , Istanbul Technical University, Turkey and A. R. BOCCACCINI, University of Erlangen-Nuremberg and Imperial College London, UK

Abstract: Nanoscale bioactive glasses of silicate composition are gaining increasing attention due to their superior bioactivity, enhanced osteoconductivity and antibacterial properties when compared to conventional (micron-sized) bioactive glasses. In this chapter, we present an overview of the technology, characterization and applications of nanoparticle and nanofibrous bioactive silicate glasses. Novel fabrication technologies are presented, covering sol-gel routes, microemulsion techniques, the gas phase synthesis method (flame spray synthesis), laser spinning and electro-spinning and the synergistic effects that determine the final properties of these materials are discussed. The advantages of nanoscaled bioactive glasses compared to conventional bioactive glasses are also discussed based on available literature evidence. The focus of the chapter is the diverse application areas of nanoscale bioactive glasses ranging from tissue engineering scaffolds to drug delivery and dentistry, considering also polymer/ bioactive glass nanocomposites as promising materials for implants, bone fillers and bioactive coatings for orthopaedic applications. The nanofeatures characteristic of this type of bioactive glass are discussed and the possibilities of expanding the use of these nanomaterials in other nanotechnology approaches aiming at advanced biomedical applications (nanomedicine) are also highlighted. Key words: bioactive glass nanoparticles, bioactive glass nanofibres, nanocomposites, tissue engineering, bioactive glass coatings, nanomedicine.

6.1

Introduction

Silicate bioactive glasses were developed for the first time by Hench and co-workers in 1969 [1]. These highly surface reactive inorganic materials are able to bond to bone tissue in a physiological environment [2]. The most widely used group of bioactive glasses for applications in the biomedical field, related to the first chemical composition developed by Hench et al. [1], consist of a silicate network incorporating sodium, calcium and phosphorus in different proportions. The traditional 45S5 bioactive glass (45S5 Bioglass®) of composition in wt%: 45 SiO2, 24.5 Na2O, 24.5 CaO and 6 P2O5, has received approval from the US Food and Drug Administration (FDA) for clinical use in the treatment of periodontal diseases as bone filler as well as in middle ear surgery [2]. A wide range of bioactive glass compositions is now available that have additional elements incorporated such as fluorine, magnesium, strontium, iron, silver, boron, potassium or zinc [3–9]. More recently, a range of other clinical applications of bioactive glasses was proposed, for example in periodontology [10, 11], endodontology [12, 13] or as coating on 129 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

130 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

metallic orthopaedic implants [14, 15]. Moreover bioactive glasses are finding application in the development of scaffolds for tissue engineering and regenerative medicine [16, 17–20]. In these approaches, both micron-sized and nanoscale particles are considered for fabrication of composite scaffolds, e.g. combining biodegradable polymers and bioactive glass [16, 18, 21]. In this context, bioactive silicate glasses have advantages in comparison to other bioactive ceramics, e.g. sintered hydroxyapatite, considering that their dissolution products have been shown to upregulate the expression of genes controlling osteogenesis [17]. There is also emerging evidence in the literature showing the potential angiogenic effects of bioactive glasses, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularisation in vivo [22–24]. The specific properties of bioactive glasses mentioned above can be influenced and possibly controlled to a greater extent in the nanoscale, e.g. developing nanoparticles or nanofibres, for a range of applications including coatings of biomedical devices, fillers in composite materials for biodegradable implants, dental fillers, tissue engineering scaffolds or drug delivery vehicles. Nanotechnology approaches are being proposed for a wide variety of medical engineering and biomedical applications in what is termed nanomedicine. Nanoscience and nanotechnology are particularly attractive in tissue engineering since the interactions between cells and biomaterial surfaces occur firstly in the nanoscale and the components of biological tissues have nanoscale dimensions [25, 26]. In the nanometer scale, key properties determining the cell–biomaterial interaction such as surface area, surface roughness, hydrophilicity and wettability, which influence cell adhesion and provide bonding properties to host (bone) tissue for long-term functionality, are completely different from the conventional (micrometer) scale [27, 28]. Nanoscaled biomaterials have in general a low defect concentration and a high ratio of surface area to volume [29, 30], which results in greater surface energy and surface bioreactivity compared with micro-scale biomaterials. In the case of bioactive glass nanoparticles, a superior in vitro bioactive behaviour in comparison with µm-scale particulate glasses is expected, mainly due to enhanced textural properties (higher surface area) than the micrometer-sized counterparts. It has been reported that the larger specific surface area of the nanoparticles allows not only for a faster release of ions but also a higher protein adsorption, which improves bioactivity [27, 31]. In addition, the use of nanosized bioactive glass particles enhances osteoblast adhesion, proliferation and differentiation, and induces an increase in the biomineralization process [27, 32]. Furthermore, the greater specific surface area of the nanosized bioactive glass particles when used as filler in biopolymers will lead to higher interface effects resulting in improved mechanical properties of the materials, compared with micrometer size particles, provided a homogeneous dispersion of the nanoparticles in the polymer matrices is achieved [33–37]. A range of techniques has emerged recently to fabricate nanoscale bioactive glasses including sol-gel [31, 38, 39], laser spinning [40], microemulsion [41] and



131

gas-phase synthesis [42–45]. The produced nanoscale bioactive glasses are being proposed for the design of numerous nanomaterials for biomedical applications, including: combination of nanofibres or nanoparticles with polymeric matrices to produce nanocomposites [33, 36, 37, 46], incorporation of nanoparticles or nanofibres into porous 3D scaffolds [34, 47, 48], nanoparticle coatings on implant surfaces [49–51] and production of non-porous materials containing nanoparticles in the form of gels, injectable materials or hard devices [52]. In this chapter, we present the current state of-the-art fabrication technologies for nanoscaled bioactive glasses (nanoparticles and nanofibres) and the application of the produced nanomaterials in the biomedical field. Section 6.2 discusses the specific nanoscale-related features of bioactive glasses. Different synthesis methods for bioactive glass nanoparticles and nanofibres are reviewed in Section 6.3. In Section 6.4, the application of bioactive glass nanoparticles and nanofibres is comprehensively reviewed. Finally in Section 6.5, a summary of the topic is presented and areas for future research are highlighted.

6.2

Characteristics of nanoscale bioactive glasses

The higher specific surface area of nanoscale bioactive glasses (nanoparticles and nanofibres) is expected to enable a faster solubility of the material (higher ion release rate) and also higher protein adsorption and thus enhanced bioactivity. Faster deposition or mineralization of tissues such as bone or teeth is possible when the tissues are in contact with nanoscale particles [31, 53]. This effect is related to the bone structure, which can be considered as a nanostructured composite mixture of collagen fibrils and carbonate hydroxyapatite nanocrystals [31, 53]. Mimicking the nanofeatures of bone on the surface of synthetic bone implants, for example, has been shown to increase bone-forming cell adhesion and proliferation [31]. Cells in their natural environment are surrounded by nanostructures in contact with other cells and with the extra-cellular matrix (ECM), formed by biomolecules configured in different geometrical arrangements and (nano)structures. In the case of bone tissue applications, nanoscale-related features of the materials, such as textural properties, surface energy and chemistry, nanotopography and wettability, control protein interactions modulating subsequent osteoblast adhesion and long-term functionality [54, 55]. Moreover nanotopography (roughness, shape/size of surface features, geometric vs. random distribution, etc.) affects cell interactions and it is expected to alter cellular behaviour when compared to conventional (µm-sized) topography [54, 56, 57]. At the nanoscale, experimental evidence indicates that cell types react specifically and in a differential manner to topographical surfaces [55]. Moreover, nanotopography enhances osteoblastic differentiation which could also promote stability and change the biomechanical environment for healing [58]. Surface energy is another important factor that regulates cell response to biomaterials and it can be altered by incorporating


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

132 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

nanoparticles or nanofibres. Surface energy modulates protein adsorption, which further regulates cell adhesion, cell spreading and proliferation [59], with higher surface energy enhancing the interaction between an implant surface and its biological environment [60, 61]. A direct relationship exists between roughness and surface energy of materials [62]. Moreover it has been demonstrated that the polar component of surface energy increases significantly with roughness and that cell adhesion enhancement is related to the degree of roughness and hydrophobicity [63]. It is also evident that the hydrophobicity or hydrophilicity of a surface can significantly alter cell behaviour. Hydrophilic materials that exhibit higher surface energy are preferred for cell attachment and proliferation [64]. Moreover, the wettability of a surface can allow for ranking materials with regards to their hydrophobic or hydrophilic category. For example, the surface wettability of alumina and titanium were shown to be enhanced by reducing particle size and this resulted in improved attachment and proliferation of preosteoblasts [65, 66]. As reported in the literature, features affected in the nanoscale, such as surface energy, nanotopography and wettability, have direct effects on cell orientation, morphology and cytoskeleton arrangements [54, 55, 67]. Although most previous results have been obtained on TiO2, alumina and hydroxyapatite, the findings should be directly applicable to nanoscaled bioactive glasses, as discussed in this chapter. For bone tissue engineering purposes, where scaffolds made of polymer/bioactive glass composites are applicable [16, 18, 68], the use of nanoscale bioactive glasses is expected to improve both the scaffolds’ mechanical and biological properties. As mentioned above, the surface bioreactivity of nanoparticles is higher than that of µm-sized particles. In addition, bioactive glass nanoparticles will induce nanostructured features on scaffold surfaces, which are likely to improve osteoblast cell attachment and subsequent cell behaviour, following the discussion above. Other advantages of the reduced size of inorganic particles include the possibility to use them to reinforce polymers in the form of nanofibres, and to process thin bioactive coatings and nanoscaled injectable systems [69].

6.3

Fabrication of bioactive glass nanoparticles and nanofibres

6.3.1 Sol-gel methods The sol-gel process is commonly applied in the context of silicate systems and other oxides and it is a technology widely used for inorganic material synthesis, for example for the fabrication of thin films, powders, nanoparticles and fibres [34, 70–72]. Li et al. [70] showed the sol-gel synthesis of silicate bioactive glasses using metal alkoxides as precursors. Typical precursors for sol-gel bioactive glasses are tetraethyl orthosilicate, calcium nitrate and triethylphosphate. After hydrolysis and polycondensation reactions a gel is formed, which after calcination at 600 to 700°C



133

forms the glass. Recent work on fabricating bioactive silicate glass nanoparticles by a sol-gel process has been carried out by Hong et al. [38]. A schematic diagram showing the sol-gel synthesis process developed by Hong et al. [34, 38] is presented in Fig. 6.1. The morphology and size of bioactive glass nanoparticles could be tailored by varying the production conditions and the feeding ratio of reagents [39, 52]. Not only the size but also the morphology of bioactive glass nanoparticles (NBG) is relevant in order to achieve the desired behaviour when applied in the biomedical field. For example, the effects of different morphologies on the in vitro bioactivity of nanosized bioactive glass particles in the system CaO-P2O5-SiO2 has been investigated [73]. In addition, surface-modified bioactive glass nanoparticles have been developed by sol-gel to improve their dispersibility by using a wet mechanical grinding technique [74]. It was reported that a layer of silane could prevent agglomeration of sol-gel derived bioactive glass nanoparticles [74]. It is usually difficult to synthesize silicate glasses in a nanosize scale with multicomponent chemical composition, e.g. by addition of specific metallic ions.

6.1 Schematic diagram for the sol-gel synthesis process of bioactive glass nanoparticle. Drafted according to the methods described in refs. [34, 38].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

134 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Delben et al. [75] have developed sol-gel derived bioactive glass nanoparticles doped with silver with a mean particle size of 100 nm. In an expansion of the sol-gel method to develop nanoscale bioactive glass, the method has been combined with electrospinning to fabricate silicate fibres [76]. Electrospinning of sol-gel precursors can result in bioactive glass fibres with diameters 1000 K s−1) and short residence times (1 ms) enable nanoparticle formation. In contrast to wet phase processes, gas phase synthesis allows generally higher production rates. Flame spray synthesis was originally developed for manufacturing carbon black [82], and it is nowadays routinely used to produce large amounts of silica and titania nanoparticles. Adapting the process to allow the use of organic liquid precursors loaded with metals instead of gaseous precursors led to increased versatility [83–85]. In this context, the metal carboxylate system is a very convenient precursor, because it allows the synthesis of oxide nanoparticles of almost any composition [85]. The process enables production of numerous nanoparticulate-mixed oxides with high chemical homogeneity. Moreover, and depending on the composition, fast quenching can


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

136 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

6.3 TEM image showing the diameter distribution of bioactive glass particles produced by microemulsion method (reprinted from ref. [41] with permission).

preserve the amorphous state of the material [42, 43]; consequently, by using flame spray synthesis, the preparation of nanoparticles of different bioactive glass compositions has become possible. Mixtures of 2-ethylhexanoic acid salts of calcium and sodium, hexamethyldisiloxane, tributyl phosphate and fluorobenzene to introduce fluorine have been employed [42].

6.3.4 Laser spinning methods Laser spinning enables the production of glass fibres with a wide range of diameters (from the nanometre to micrometre scale) [86]. Large quantities of nanofibres can be produced with specific, controllable chemical compositions without the necessity of any chemical additives or post-heat treatments. The process is very efficient; nanofibres are produced in several microseconds. The laser spinning technique essentially involves the quick heating and melting of a small volume of the precursor material up to high temperatures using a high power laser. At the same time, a supersonic gas jet is injected into the melt volume



137

to blow the molten material. The molten material is then quickly stretched and cooled by the supersonic gas jet [86]. Long fibres with high length-to-diameter ratios can be produced, and the obtained material is in amorphous form because of the high cooling speed. Glass fibres in the form of a disordered net of intertwined amorphous micro- and nanofibres have been produced [87, 88]. More recently, Quintero et al. [40] developed the first bioactive glass nanofibres by the laser spinning technique. They produced bioactive glass nanofibres in the 45S5 Bioglass® and 52S4.6 silicate compositions with average diameter in the range 200 to 300 nm. Bioactivity testing demonstrated that the nanofibres were covered by a foamy and porous layer of amorphous calcium phosphate after immersion in SBF for five days. The laser spinning technique is thus proposed as an effective method of producing bioactive glass nanofibres in desired compositions. As a summary of the fabrication methods discussed in this section, their advantages and disadvantages are listed in Table 6.1. Table 6.1 Methods of fabrication for nano-scaled bioactive glasses showing their advantages and disadvantages Shape of the nano-structure

Method of fabrication

Advantages

Disadvantages

Particle

Sol-gel

Controlled porosity, controlled pore size and surface area, homogeneous products with high purity, simple operation, low cost of precursors

Problems with residual solvent and water, high calcination temperature. Relatively time consuming and limited to some compositions

Microemulsion

Ability to synthesize nanosized particles of organic and inorganic composition with minimum agglomeration, low processing temperature

Low production yield and the use of a large quantity of oil and surfactant phases

Gas phase synthesis method

High chemical homogeneity, High processing higher production rates, no temperatures need to additional source of energy

Sol-gel and electrospinning

Homogeneous products with high purity, simple operation

Difficulties in controlling many parameters that affect the fibre diameter

Laser spinning

Higher production rates, controllable chemical compositions, very fast process, high chemical homogeneity, long fibres with extraordinarily high length to diameter ratios

High processing temperatures

Fibre


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

138

Bioactive glasses

6.4

Applications of nanoscale bioactive glasses

6.4.1 Composites for tissue engineering scaffolds Inorganic phases can be added to different polymer matrices in the form of micron-sized or nanoscale particles or fibres to form bioactive composite materials [16]. The size of the filler particles affects the effective mechanical and biological properties of the composites. For example, the introduction of nanoscale fillers with desired morphology should lead to a higher increase of mechanical strength and stiffness in comparison to neat polymers and to composites with micronsize reinforcement, provided the nanoparticles or nanofibres are dispersed homogeneously in the matrix [89]. Additionally, in the case of bioactive silicate glass nanoparticles and nanofibres, they will induce higher alkalinity when compared to commercially available (µm-sized) bioactive glass particles [90]. The larger specific surface area of the nanoparticles will increase interface-related effects, contributing to improved bioactivity. Further advantages of the use of nanoscale fillers in composites for medical applications, in particular in relation to surface effects for tissue engineering scaffolds, were discussed in Section 6.2. Synthetic polymer/nanoparticulate bioactive glass composites Poly(3-hydroxybutyrate) (P(3HB))/nanoparticulate bioactive glass composites with different filler concentrations have been fabricated by solvent casting [30]. The addition of nanoparticles was shown to have a significant stiffening effect in comparison with the µm-sized counterparts. Moreover surface effects induced by the nanoparticles (nanotopography) considerably improved total protein adsorption compared to the unfilled polymer and to composites containing micron-sized bioactive glass particles [30]. In addition, cell proliferation investigations using osteoblast-like cells confirmed enhanced cytocompatibility of the P(3HB)/bioactive glass composites [30]. Misra et al. [91] have also reported that the ALP activity of MG-63 cells on nanoparticulate bioactive glass/P(3HB) composites was considerably higher than that on the control surface confirming the suitability of the composites for bone tissue engineering. Zheng et al. [92] developed composites using poly(hydroxybutyrate-2-co2-hydroxyvalerate) (PHBV) containing biomimetically synthesized nanosized bioactive glass (BMBG) (CaO-P2O5-SiO2). The porous composites were shown to be bioactive, and cell attachment studies indicated that the material has attractive biomineralization and cell biocompatibility [92]. Composites combining poly(L-lactic acid) and sol-gel-derived bioactive glass-ceramic (BGC) nanoparticles were fabricated by Hong et al. [34] using thermally induced phase separation. It was shown that composites containing BGC nanoparticles with lower phosphorous and higher silicon content had higher bioactivity than that of the BGC with lower silicon and higher phosphorous content [34]. The effect of nanoparticle content on the properties of nanocomposite



139

scaffolds has been also investigated [35] and an improvement of the mechanical properties was measured. A similar system was developed by El-Kady et al. [93] using sol-gel derived bioactive glass nanoparticles and poly(L-lactide) (PLA). In this system, it was shown that the scaffold’s pore size decreased with the increase of glass nanoparticles content. A recent approach to improve the mechanical properties of nanoparticulate bioactive glass/PLLA composites by using solvent evaporation has been reported by Liu et al. [36, 37]. It was shown that surface modification of nanosized bioactive glass particles by grafting organic molecules or polymers is a convenient solution to improve the mechanical properties of the composites. Moreover surface modified bioactive glass/PLLA composites exhibited much better cell proliferation ability than non-modified bioactive glass/ PLLA composites or pure PLLA [36, 37]. Natural polymer/bioactive glass nanocomposites Polysaccharides (starch, chitin, chitosan) and proteins (silk, collagen) are candidate natural polymers for preparing nanocomposites for biomedical applications. Peter et al. [94, 95] have synthesized α-chitin/sol-gel derived bioactive glass-ceramic nanoparticle and chitosan/sol-gel derived bioactive glassceramic nanoparticle composite scaffolds by using lyophilization technique. Macroporous composite scaffolds with pore size in the range 150 to 300 µm were fabricated (Fig. 6.4) [95]. In vitro studies showed the deposition of apatite and the attachment of osteoblast-like cells (MG-63) on the surface of the composite scaffolds [94, 95]. More recently, the same authors [96] have fabricated chitosan– gelatine/nano-sized bioactive glass-ceramic nanocomposite scaffolds by through freezing and lyophilization technique. It was reported that high surface area of

6.4 SEM images showing the macro porous microstructure of composite scaffold synthesized from chitosan/sol-gel derived bioactive glass-ceramic nanoparticle by using lyophilization technique (a). Pore size ranged from 150 to 300 µm, and nBGC particles were on the chitosan matrix (b) (reproduced from ref. [95] with permission).


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

140 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

bioactive glass-ceramic nanoparticles resulted in increasing protein adsorption especially adhesive proteins [96]. New porous bioactive nanocomposites combining sol-gel derived bioactive glass nanoparticles (BG), collagen (COL), hyaluronic acid (HYA) and phosphatidylserine (PS) have been fabricated by a combination of sol-gel and freeze-drying methods [97]. A bioactive nanocomposite was also synthesized by cross-linking collagen and HYA by using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS). Biomineralization, degradation in SBF and mechanical strength of the EDC/NHS-cross-linked BG-COL-HA-PS composite scaffolds were better than those of the scaffolds without HYA, PS and the cross-linking process. PS and HYA can contribute to regulating the biomineralization process, inducing HA to precipitate on the surface of the composites. The in vivo bone regeneration ability of the EDC/NHS-cross-linked BG-COL-HA-PS composite scaffolds was investigated by Xie et al. [98] using a rabbit radius defect model. X-ray and histological studies showed the bone regeneration ability of both nanocomposites and for nanocomposites combined with growth factors (BMP). Moreover, nanocomposites containing BMP showed better ectopic bone formation ability [98]. In related studies, Luz and Mano [99] have proposed a new composite membrane combining chitosan with sol-gel derived bioactive glass nanoparticles based on both ternary (SiO2-CaO-P2O5) and binary (SiO2-CaO) systems. It was reported that bioactive glass nanoparticles could be distributed homogeneously in the biodegradable polymeric membrane because of their reduced size, thus promoting bioactivity [99]. Nanocomposites containing bioactive glass nanofibres Kim et al. [47] have developed PLA composites filled with sol-gel-derived bioactive glass as a nanoscale composite fibre using electrospinning (ES). The in vitro bioactivity and osteoblast responses of the developed nanocomposites were studied [48]. These nanocomposites showed excellent bioactivity, inducing CaP precipitation within 24 hours of immersion in SBF. The results have also been confirmed by Noh et al. [100] in a similar study. Kim et al. [101] also developed BGNF-collagen nanocomposite both in the form of a thin membrane and as macroporous scaffold. BGNF-collagen nanocomposites were seen to exhibit high bioactivity, which was assessed by the rapid formation of bonelike apatite minerals on their surfaces when immersed in SBF. Moreover, the nanocomposites enhanced the adhesion and growth of human osteoblast-like cells [101]. In related research, Lee et al. [102] have produced poly(e-caprolactone) (PCL)/ sol-gel-derived BGNF nanocomposites. The glass nanofibres were distributed homogeneously in the PLC matrix, showing a much rougher surface than the pure PCL. The precipitated apatite covered the surface of the nanocomposite membrane almost completely after immersion in SBF for 14 days. Osteoblastic cells



141

6.5 Surface and cross-sectional images of PCL/bioactive glass nanocomposites containing 20 wt% BGP (a, b) and 20 wt% BGNF (c, d) (reproduced from ref. [46] with permission).

(MC3T3-E1) were seen to spread better and to grow with many cytoplasmic extensions, showing improved proliferation behaviour in comparison to those on the pure PCL membrane [102]. Jo et al. [46] have fabricated (PCL)/sol-gel derived BGNF composites exhibiting a highly homogeneous BGNF distribution (Fig. 6.5). This microstructure resulted in a significant improvement of the biological and mechanical properties of the PCL/BGNF composites, compared to that of the micron-sized ones. Multifunctional composite scaffolds The desired combination of biocompatibility of biodegradable polymers and bioactivity of bioceramics can be achieved by preparation of porous polymer/ ceramic composites by different methods, as mentioned above [16]. However, the need for advanced scaffold systems has prompted the addition of different functionalities into the materials to be able to closely mimic the natural bone’s structure and properties. Properties such as bioactivity, mechanical competence, electrical or magnetic conduction, growth factor and drug delivery, antioxidative effects and antibactericidal behaviour are being considered for designing new multifuncational scaffolds. Misra et al. [103] have prepared P(3HB)/micro scaled Bioglass® 3D composite scaffolds using the conventional solvent casting/


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

142 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

particulate leaching technique and by employing commercially available sugar cubes as the porogen preform. Functionalized P(3HB)/micro scaled Bioglass® 3D composite scaffolds were fabricated by the systematic addition of appropriate amounts of vitamin E and multiwalled carbon nanotubes (MWCNTs) in the polymer solution and by sonicating the mixture before impregnating it onto the sugar sacrificial preforms. Studies indicated that each additive was able to contribute to the overall performance of the composite, i.e. Bioglass® particles imparting bioactivity, vitamin E improving protein adsorption and finally MWCNTs inducing electrical conductivity. P(3HB) composite scaffolds were also produced, using nanoscale bioactive glass particles to enhance the antimicrobial properties. Bactericidal studies revealed that nanoscale bioactive glass particles ensure both biocompatibility and enhanced antimicrobial properties. It was also concluded that the higher specific surface area of NBG particles played a vital role in imparting the desired antimicrobial activity [103]. The novel properties of nanoscaled bioactive glasses make them good candidates for the further development of advanced multifunctional 3D scaffolds. However, it is clear that the present status of research and development in the field of multifunctional 3D scaffolds is still at the starting point for overcoming some of the limitations of the biomaterials currently used in bone tissue engineering. In future, it should be possible to design multifunctional 3D composites with incorporation of NBG (nanoparticles, nanofibres or both) for a variety of healthcare applications beyond the general field of tissue engineering therapeutics [104], i.e. ranging from drug delivery to biosensing devices.

6.4.2 Applications in dentistry Bioactive glasses have been used in dentistry for de- and remineralization of dentin, root canal disinfection, restorative dental applications, augmenting alveolar ridges and for treating periodontal pockets and dentin hypersensitivity [10–13, 105–107]. Such favourable properties of bioactive glasses as bioactivity, ability to mineralize dentine and their antimicrobial effects make them interesting materials for dental applications. Bioactive glasses are for example promising candidates for the remineralization of human dentin and have potential as a filler component in mineralizing restorative materials [108]. However, the application of bioactive glass as a remineralization agent in dental practice has been limited owing to its relatively long reaction times. An alternative approach is to decrease glass particle size and thus to increase the surface area, which should enhance dissolution of ions from the glass resulting in an acceleration of the remineralization of dentin. Vollenweider et al. [90] investigated the in vitro remineralization capabilities of bioactive glass nanoparticles (45S5 composition) fabricated by flame spray synthesis (see Section 6.3.3). It was reported that NBG treatments of 10 and 30 days resulted in a markedly higher



143

mineral content compared to equivalent treatments with suspensions of conventional (µm-size) bioactive glass particles. The mineral content of dentin remineralized for 30 days with NBG was equivalent to natural dentin. Kernen et al. [109] also reported the rapid remineralization rate of flame spray-derived nanoparticulate 45S5 bioactive glass. It was found that NBG-treated discs incubated for 7 days showed a degree of mineralization comparable to the native control disc [109]. Bioactive glasses are also interesting materials in dentistry because of their antimicrobial effect in closed systems [106, 107]. The antimicrobial effect of bioactive glasses has been attributed to their ability to raise pH in an aqueous environment [110]. Studies with conventional bioactive glasses showed that their antibacterial efficacy in human teeth is inferior to that of calcium hydroxide, the gold standard material [105]. Some attempts have been made to increase the antimicrobial efficacy of bioactive glasses for example incorporating silver into the glass network [111–113]. As an alternative way, Waltimo et al. [108] used nanometric 45S5 bioactive glass to improve the antibacterial properties against different Enterococcus faecalis strains in a direct exposure model. The antibacterial effect of nanoparticulate bioactive glass appears to be directly linked to its high surface area, and thus the resulting release of ionic components in solution [108]. Gubler et al. [114] also investigated the antibacterial efficiency of nanometric bioactive glasses with the compositions 28S5, 45S5 and 77S. Studies showed that in addition to the high surface area of bioactive glass nanoparticles, the antimicrobial effect is also related to the amount of sodium and thus the resulting alkaline environment. Microorganisms remaining or re-entering the root canal system are the main cause of post-treatment disease following root canal treatment [115]. Alkaline capacity and the antibacterial properties of bioactive glasses make them potential candidates for root canal disinfection treatments. The possibility of using 45S5 bioactive glass nanoparticles as root canal disinfectants has been investigated [13]. It was reported that not only the specific surface area, but also the total mass of material per volume in bioactive glass slurries is important for root canal disinfectant treatments [13]. More recently, Mortazavi et al. [116] have synthesized 58S, 63S, and 72S bioactive glass nanopowders using the sol-gel technique, and investigated the antibacterial effects of bioactive gel glass nanoparticles on aerobic bacteria. It was found that 58S and 63S bioactive glass nanopowders had antibacterial activity even at concentrations lower than those currently used in clinical applications. However, 72S bioactive glass nanopowder showed no antibacterial effect because of the high SiO2 content. It was concluded that bioactive glass nanopowders with antibacterial properties could be considered for the treatment of oral bone defects and disinfection of the root canal [116]. Besides the alkaline capacity and antibacterial properties, a certain level of radiopacity is necessary for a clear distinction between the material and the surrounding tissues in clinical practice. To the authors’ knowledge, Mohn et al.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

144 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

6.6 Transmission electron microscopy images of classical nanometric bioactive glass (a) and bioactive glass (flame derived) containing 50 wt% bismuth oxide (b) (reproduced from ref. [45] with permission).

[45] were the first to develop flame spray-derived bioactive glass nanoparticles including bismuth oxide as a radiopacifier. Flame spray synthesis resulted in novel bioactive glass nanoparticles with sticking bismuth oxide (2 to 7 nm) onto the surface (Fig. 6.6) [45]. The studies revealed that bioactive glasses with bismuth oxide have high in vitro bioactivity, alkaline capacity and radiopacity, which make them potential bioactive root canal dressing or filling materials [45]. The various primary positive results regarding the bioactivity and antibacterial properties of bioactive glass nanoparticles make them attractive materials in dentistry, notably in dentin regeneration and root canal disinfection, as briefly discussed in this section. Furthermore, bioactive glass particles can be potential candidates as nanofillers in synthesizing new dentin composites to improve the bioactivity, radiopacity and mechanical properties of the composite systems.

6.4.3 Bioactive coatings and other orthopaedic applications The clinical applications of bioactive glasses have been limited largely to nonload bearing parts due to their inferior mechanical properties. In order to solve the lack of strength and fracture toughness of bulk bioactive glasses for loadbearing applications, coatings of these materials on metallic prostheses have been developed [14, 15]. Several techniques have been investigated for the preparation of glass coatings, such as enamelling [117], plasma spraying [118], ion beam sputtering [119], sol-gel [120], pulsed-laser deposition (PLD) [121], electrophoretic deposition (EPD) [122–124] and laser cladding [125]. Several



145

factors influence the properties of bioactive glass coatings such as coating thickness, crystallinity, chemical purity, porosity and adhesion [126]. Novel bioactive glass nanoparticles represent a significant opportunity to develop improved bioactive and nanostructured orthopaedic coatings. As discussed above, bioactive glass nanoparticles have better bioactivity, antibacterial properties and enhanced bone-forming cell adhesion and proliferation compared with micronsized particles. Recently, Fathi et al. [49] prepared bioactive glass nanopowders for coating 316L stainless steel by the sol-gel technique. The particle size of the bioactive glass was less than 100 nm. Nanopowders were seen to attach very well to the stainless steel substrate and crack-free homogeneous coatings were obtained. The most important finding was the improved corrosion resistance, biocompatibility and bone bonding ability of the metallic substrate [49]. Sol-gel derived bioactive glass nanoparticles have also been used to coat different materials, with the intention of combining improved mechanical properties and high bioactivity in one material [33, 49]. Bioactive glass nanoparticle coating using the sol-gel technique has been applied for example on the struts of porous HA scaffolds by Esfahani et al. [33], in order to improve the mechanical properties of the scaffold. It was shown that the compressive strength of scaffolds increased and a new crystalline phase was detected with the increase in sintering temperature. According to Esfahani et al. [33], crystallization occurred in bioactive glass nanoparticles resulting in an improvement of the mechanical properties. Fathi et al. [50] prepared a sol-gel derived bioactive glass nanopowder coating for treating oral bone defects. It was reported that crack-free and homogeneous bioactive glass coatings were achieved with no observable defects. In vitro studies showed that the bioactive glass coating induced the formation of a weakly crystalline hydroxyapatite-rich layer on the bioactive glass surface as an indication of bioactivity [50]. Furthermore, novel nanoscaled coatings can be used for dental hard tissues to improve the aesthetic appearance and to protect them against ageing processes such as wear and cracking [127]. In related investigations, Couto et al. [51] have fabricated polycation (chitosan) and bioactive glass nanoparticle multilayer coatings by the layer-by-layer (LbL) technique. In this method, a glass panel was first dipped in a 1% (v/v) acetic acid solution containing 0.4% (w/v) of chitosan, then into water, then into a 0.4% (w/v) bioactive glass aqueous solution and finally into water again. The multilayer was achieved by repeating the dipping process sequentially. The scheme of this process is shown in Fig. 6.7 [51]. Spherical nanoparticles with sizes in the range 30 to 100 nm were dispersed homogeneously on the surface of the multilayered coatings. The organic component, chitosan, provided viscoelastic properties to the final coating, and the bioactive glass induced the required bioactivity for improving attachment of the coating to the bone. The coating was evaluated by immersing specimens in simulated body fluid, and the formation of a surface layer of hydroxyapatite, as a marker of bioactive behaviour, was confirmed.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

146 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

6.7 Scheme of the layer-by-layer coating procedure for production of chitosan/nano-bioactive glass (BG) coatings (a). After each cycle it is expected that a complete nanostructured bi-layer will be formed, resulting in a multilayered film adsorbed at the surface after repeated layering, according to Couto et al. (b) (adapted from ref. [51]).

6.4.4 Applications in drug delivery and nanomedicine Research efforts are increasingly devoted to the development of new drug-delivery systems with greater efficiency, lower toxicity, controlled and prolonged drug release capability, predictable therapeutic response and safety [128]. Drug-delivery systems based on particulate carriers will benefit from nanoscale dimensions, as particles need to reach the given location in the body based on the size of vessels of the human circulatory system in order to transport the desired drug molecules to the targeted cells or tissues and to release them in a controlled manner [128, 129]. Ordered mesoporous bioactive glasses and composites of bioactive glasses and polymers have been proposed as delivery systems for antibiotics and other antibacterial agents, anti-inflammatory drugs, fluoride ions, vascular endothelial growth factor, proteins and peptides [20, 22–24, 130–136]. Ionic products of bioactive glass dissolution also increase the proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression as well as protein synthesis [117, 137]. It is now well established that ionic dissolution products released from bioactive glass upregulate seven families of genes that control osteogenesis [17, 138, 139] and promote angiogenesis [22–24, 132]. Besides these advantages,



147

however, the drug-storage capacity of conventional bioactive glasses is generally relatively low and the drug delivery is difficult to control [140]. An alternative way of overcoming these disadvantages is the synthesis of nanoscaled bioactive glasses because these offer a number of distinct advantages for drug adsorption and delivery over conventional (µm-sized) particles. Nanoparticles have in general relatively higher intracellular uptake compared to microparticles as discussed above [141]. The efficiency of uptake of nanoparticles has been reported to be 15to 250-fold greater than that of µm-sized particles [142]. Due to their sub-micron size, nanoparticles can penetrate deep into tissues through fine capillaries, and are generally taken up efficiently by cells [143]. This behaviour enables the efficient delivery of therapeutic agents to target sites in the body [144]. Nanosized particles in particular can affect the efficiency of the immune response [145]. Moreover, nanoparticles show larger storage capacity than microparticles because of their high surface area. Yun et al. [146] have synthesized mesoporous bioactive glass (MBG) nanospheres of SiO2-CaO-P2O5 composition, which exhibit large specific surface area and pore volume. The size of the nanospheres was shown to depend on the amount of CaO incorporated and it can be controlled over the range of diameters from 20 to 200 nm (Fig. 6.8) [146]. In vitro studies indicated that the MBG nanospheres with high amount of CaO show good in vitro bone-forming bioactivity as well as favourable in vitro biocompatibility. Cytotoxicity tests of MBG nanospheres with a high amount of CaO revealed that they have no negative effect on macrophage cell behaviour. It was concluded that MBG nanospheres are promising materials for drug delivery applications [146]. Further research is focusing on the development of nano-carriers to enhance the effective targeting in specific cancer regions without destroying or affecting the viability of nearby normal tissues [147]. It was shown that ferromagnetic bioactive glasses and glass-ceramics containing magnetite could be used for hyperthermia treatment of cancer [148]. Recently, Wang et al. [149] developed a novel magnetic degradable material, adding Fe ions to bioactive glass (Na2O-CaO-P2O5-SiO2) as thermoseed for hyperthermia cancer therapy using the sol-gel method. The particle size was measured in the range 50 to 100 nm and spherical and rod shape particles were developed; the morphology was dependent on the amount of iron ion added. In a biocompatibility test, these magnetic bioactive glasses had no significant influence on cell viability and mediated low cytotoxicity when cultured with fibroblasts. The material was cultured in vitro with either human Caucasian lung carcinoma (A549) or normal HFL1 fibroblast cells to examine the hyperthermia effect. In vitro cell culture studies clearly indicated that after exposure to an alternating magnetic field, the cell number of human Caucasian lung carcinoma cells (A549) significantly decreased, while normal HFL1 fibroblasts were still alive without severe damage. It was stated that magnetic degradable bioactive glass incorporated with Fe ions would be a potential candidate for tumour hyperthermia treatment in future [149].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

148 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

6.8 FEG-SEM images of MBG-Ca4 (a), MBG-Ca8 (b), MBG-Ca16 (c) and a comparison of their particle sizes as a function of the amount of calcium (d), according to Yun et al. (reproduced from ref. [146] with permission).

Smart hydrogels as injectable scaffolds are gaining relevance for several tissue engineering applications since they enable the encapsulation of cells and bioactive agents in a biodegradable matrix to be delivered through minimally invasive procedures [150, 151]. To the authors’ knowledge, Couto et al. [52] were the first to develop injectable biodegradable materials with bioactive glass nanoparticles in order to produce thermo-responsive hydrogels for orthopaedic reconstructive and regenerative medicine applications. In their study, chitosan-β-glycerophosphate salt formulation was combined with sol-gel derived bioactive glass nanoparticles to synthesize novel thermo-responsive hydrogels. The inner structure of the hydrogels was characterized by using cryogenic scanning electron microscopy (cryoSEM) and it was found that bioactive glass nanoparticles were well dispersed in the organic matrix (Fig. 6.9) [52]. In vitro bioactivity tests showed that the bioactive glass nanoparticles incorporated in the chitosanbased thermo-responsive system induced the formation of bone-like apatite clusters that are well integrated in the hydrogel organic structure. It was also observed that the density of the apatite precipitates increased with increasing © Woodhead Publishing Limited, 2011


149

6.9 CryoSEM images of the surface fracture of the BG-30% hydrogel: global view, where the arrows indicate some clusters of BG particles (a); magnified image, where the arrows indicate small agglomerates of BG nanoparticles deposited on the hydrogel walls (b): in the upper right-hand side there is part of a cluster of nanoparticles involved with chitosan (reprinted from ref. [52] with permission).

NBG content and soaking time in SBF, which indicates that the system could promote a positive contact with surrounding tissue upon injection in a bone defect. The high surface area, enhanced bioactivity and antibacterial properties of nanoscaled bioactive glasses make them promising materials for slow and targeted drug delivery systems. However, although significant advantages of nanoscaled bioactive glasses are expected in combined drug delivery and regenerative medicine approaches, as discussed above, only limited work has been reported to date. Moreover a large amount of biological information is necessary to fully understand the drug-delivery functions of these novel systems. It is clear that using bioactive glass nanoparticles in drug delivery systems and nanomedicine is in its early years and comprehensive research should be carried out to measure and assess potential exposure risks of nanoparticles to patients. In the future, the challenge would be to develop the next generation of advanced smart biomaterials with the integration of extended functions such as antioxidative and anticancer functions and in several of these novel applications nanoscale bioactive glasses will be the materials of choice. Table 6.2 summarizes recent investigations on the biomedical application areas of nanoscale bioactive glasses. The novel properties of nanoscale bioactive glasses make them promising materials for a variety of healthcare applications, many of which are bound to expand in the near future as more in-depth and relevant knowledge from basic research is generated.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Method of fabrication

Solvent casting

Thermally-induced phase-separation

Solvent evaporation

Freezing/lyophilization

Freeze-drying

Freeze-drying

Sol-gel/ES

Sol-gel/ES

Sol-gel/ES

Sol-gel/ES

Solvent casting/particulate leaching technique

Flame spray synthesis

Composition

P(3HB)/NBG

PLA/NBG

PLLA/surface modified NBG

Chitin/NBG and chitosan/NBG

NBG-COL-HYA-PS

EDC/NHS-crosslinked NBG-COL-HA-PS

PLA/BGNF

Collagen/BGNF

PCL/BGNF


PCL/BGNF

P(3HB)/NBG

NBG

NBG particles

3D porous scaffold

Porous matrix

Membrane

Membrane and porous scaffold

3D porous scaffold

Porous scaffold

Porous scaffold

Porous scaffold

Porous scaffold

3D porous scaffold

2-D films

End product shape

Rapid remineralization rate

Enhanced biocompatibility and antimicrobial properties

Significant improvement of the biological and mechanical properties and in vivo animal test results showed bone-forming ability

Rough surface improved proliferation behaviour

High bioactivity and good cell adhesion and growth

Excellent bioactivity and good osteoblast response

Ability of bone regeneration

Better biomineralization, mechanical strength, cell attachment and proliferation ability

Improvement in cell adhesion and proliferation and increasing protein adsorption

Less nanoparticle aggregation, improved mechanical properties, bioactivity, cell adhesion and growth

Improved bioactivity and mechanical properties

Enhanced bioactivity, cell adhesion and growth

Key results achieved

[109]

[103]

[46]

[102]

[101]

[47, 48]

[98]

[97]

[94, 95, 96]

[36, 37]

[34, 35]

[30, 91]

Reference

Table 6.2 Selected biomedical studies carried out on nanoscaled bioactive glasses, covering a wide range of both in vivo and in vitro investigations

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X




Sol-gel

Sol-gel

Sol-gel

LbL technique

Sol-gel

Sol-gel

Combining sol-gel derived NBG with chitosan–βglycerophosphate salt formulation

NBG

NBG

NBG

NBG

NBG

Chitosan/NBG

NBG

DP-NBG with Fe

Chitosan/NBG

Hydrogel

Sphere and rod shape particles

Nanosphere

Nanoparticle multilayer coatings

Nanopowder coating

Porous scaffold

Nanopowder coating

NBG particles

NBG particles

[108, 114]

[33]

[49]

[45]

Promoting a positive contact with surrounding tissue upon injection in a bone defect

Potential candidate for tumour hyperthermia treatment

Good in vitro bone-forming bioactivity and biocompatibility

Provided viscoelastic properties and bioactivity to the final coating

[52]

[149]

[146]

[51]

Promote bone formation in osseous defects and [50] bone grafting to improve the long-term prognosis of dental implants

Improved mechanical properties without using any polymeric materials

Improvement of the corrosion resistance, biocompatibility and bone bonding ability of the metallic substrate

High in vitro bioactivity, alkaline capacity and radiopacity

Improved antibacterial properties

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

152

Bioactive glasses

6.5

Conclusions

The preparation of bioactive glasses in nanoparticle and nanofibre form has recently become feasible by advances in a variety of processing methods. Nanoscale particulate and nanofibre bioactive glasses have shown advantages over conventional (micron-sized) bioactive glasses due to their large surface area and enhanced solubility as well as bioreactivity. In addition, their application leads to nanotopographic surface features in biomaterials, e.g. scaffolds and coatings. These nanomaterials and their advantages have also inspired researchers to investigate a range of new biomedical applications. In this chapter, new developments in the fabrication techniques of nanoscale bioactive glass particles and fibres and the biomedical applications of these novel materials have been reviewed and comprehensively discussed. Substantial advantages of nanoscale bioactive glasses compared to conventional (µm-scale) bioactive glasses were demonstrated in particular for bone tissue engineering, dentistry, orthopaedic coatings, antibacterial materials and drug delivery systems. Current research shows that it is possible to design improved 3D composite scaffolds by incorporating nanoscaled bioactive glasses (and other additives), which exhibit multifunctionalities such as bioactivity, electrical conduction and antibacterial behaviour. These 3D multifunctional composite scaffolds will have applications in tissue engineering therapeutics including drug delivery and biosensing functions. Moreover, the intrinsic properties of nanoscaled bioactive glasses make them promising materials in dentistry, especially for dentin regeneration, root canal disinfection and as nanofillers in new dentin composites to improve the bioactivity, radiopacity, antibacterial effect and mechanical properties of the materials. Other applications of nanoscale bioactive glasses in nanomedicine are being explored. The new challenge in this field is to develop the next generation of smart biomaterials with the integration of extended functions such as antioxidative and anti-cancer functions, where nanoscaled bioactive glasses may represent a key component for novel biomedical devices. It is anticipated that further research efforts will lead to relevant basic knowledge and fundamental information on these novel nanomaterials, e.g. cell biologydriven research, which will prompt the further expansion of their biomedical applications.

6.6

Acknowledgment

ME gratefully acknowledges the financial support from The Scientific and Technological Research Council of Turkey (TUBITAK), Turkey.

6.7

References

1. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K. Bonding mechanisms at the interface of ceramic prosthetic materials. J Biomed Mater Res 1971; 5(6):117–141.



153

2. Hench L. L. Bioceramics. J Am Ceram Soc 1998; 81(7):1705–1728. 3. Sepulveda P., Jones J. R., Hench L. L. Characterization of melt-derived 45S5 and solgel-derived 58S bioactive glasses. J Biomed Mater Res 2001; 58(6):734–740. 4. Saravanapavan P., Jones J. R., Pryce R. S., Hench L. L. Bioactivity of gel-glass powders in the CaO-SiO2 system: A comparison with ternary (CaO-P2O5-SiO2) and quaternary glasses (SiO2-CaO-P2O5-Na2O). J Biomed Mater Res A 2003; 66A(1):110–119. 5. Oki A., Parveen B., Hossain S., Adeniji S., Donahue H. Preparation and in vitro bioactivity of zinc containing sol-gel-derived bioglass materials. J Biomed Mater Res A 2004; 69A(2):216–221. 6. Vallet-Regi M., Salinas A. J., Roman J., Gil M. Effect of magnesium content on the in vitro bioactivity of CaO-MgO-SiO2-P2O5 sol-gel glasses. J Mater Chem 1999; 9(2):515–518. 7. Liu X., Huang W., Fu H., Yao A., Wang D., Pan H., Lu W. W., Jiang, X., Zhang X. Bioactive borosilicate glass scaffolds: in vitro degradation and bioactivity behaviours. J Mater Sci Mater Med 2009; 20:1237–1243. 8. Lao J., Nedelec J. M., Jallot E. New strontium-based bioactive glasses: physicochemical reactivity and delivering capability of biologically active dissolution products. J Mater Chem 2009; 19:2940–2949. 9. Varanasi V. G., Saiz E., Loomer P. M., Ancheta B., Uritani N., Ho S. P., Tomsia A. P., Marshall S. J., Marshall G. W. Enhanced osteocalcin expression by osteoblast-like cells (MC3T3-E1) exposed to bioactive coating glass (SiO2-CaO-P2O5-MgO-K2ONa2O system) ions. Acta Biomater 2009; 5:3536–3547. 10. Zamet J. S., Darbar U. R., Griffiths G. S., Burgin W., Newman H. N. Particulate bioglass (Perioglas®) in the treatment of periodontal intrabony defects. J Dent Res 1997; 76:2219. 11. Gatti A. M., Simonetti L. A., Monari E., Guidi S., Greenspan D. Bone augmentation with bioactive glass in three cases of dental implant placement. J Biomater Appl 2006; 20(4):325–339. 12. Zehnder M., Soderling E., Salonen J., Waltimo T. Preliminary evaluation of bioactive glass S53P4 as an endodontic medication in vitro. J Endodont 2004; 30(4):220–224. 13. Waltimo T., Mohn D., Paque F., Brunner T. J., Stark W. J., Imfeld T., Schaetzle M., Zehnder M. Fine-tuning of bioactive glass for root canal disinfection. J Dent Res 2009; 88(3):235–238. 14. Gomez-Vega J. M., Saiz E., Tomsia A. P., Marshall G. W., Marshall S. J. Bioactive glass coatings with hydroxyapatite and Bioglass® particles on Ti-based implants. 1. Processing. Biomaterials 2000; 21(2):105–111. 15. Kitsugi T., Nakamura T., Oka M., Senaha Y., Goto T., Shibuya T. Bone-bonding behavior of plasma-sprayed coatings of Bioglass®, AW-glass-ceramic, and tricalcium phosphate on titanium alloy. J Biomed Mater Res 1996; 30(2):261–269. 16. Rezwan K., Chen Q. Z., Blaker J. J., Boccaccini A. R. Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 2006; 27(18):3413–3431. 17. Xynos I. D., Hukkanen M. V. J., Batten J. J., Buttery L. D., Hench L. L., Polak J. M. Bioglass® 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: Implications and applications for bone tissue engineering. Calcified Tissue Int 2000; 67(4):321–329. 18. Yao J., Radin S., Leboy S., Ducheyne P. The effect of bioactive glass content on synthesis and bioactivity of composite poly(lactic-co-glycolic acid)/bioactive glass substrate for tissue engineering. Biomaterials 2005; 26(14):1935–1945.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

154 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

19. Hench L. L., Polak J. M. Third-generation biomedical materials. Science 2002; 295(5557):1014–1017. 20. Chen Q. Z. Z., Thompson I. D., Boccaccini A. R. 45S5 Bioglass®-derived glassceramic scaffolds for bone tissue engineering. Biomaterials 2006; 27(11):2414–2425. 21. Mano J. F., Sousa R. A., Boesel L. F., Neves N. M., Reis R. L. Bioinert, biodegradable and injectable polymeric matrix composites for hard tissue replacement: state of the art and recent developments. Comp Sci Tech 2004; 64(6):789–817. 22. Day R. M., Boccaccini A. R., Shurey S., Roether J. A., Forbes A., Hench L. L., Gabe S. M. Assessment of polyglycolic acid mesh and bioactive glass for soft-tissue engineering scaffolds. Biomaterials 2004; 25:5857–5866. 23. Gorustovich A. A., Roether J. A., Boccaccini A. R. Effect of bioactive glasses on angiogenesis: in vitro and in vivo evidence. A review. Tissue Eng Part B 2010; 16(2):199–207. 24. Leu A., Leach J. K. Proangiogenic potential of a collagen/bioactive glass substrate. Pharm Res 2008; 25:1222–1229. 25. Fisher J. P., Mikos A. G., Bronzino G. D. Tissue Engineering: Nanocomposite Scaffolds for Tissue Engineering. CRC Press 2007; p. 11. 26. Hajiali H., Karbasi S., Hosseinalipour M., Rezaie H. R. Preparation of a novel biodegradable nanocomposite scaffold based on poly(3-hydroxybutyrate)/bioglass nanoparticles for bone tissue engineering. J Mater Sci: Mater Med 2010; 21:2125–2132. 27. Webster T. J., Ergun C., Doremus R. H., Siegel R. W., Bizios R. Enhanced functions of osteoblasts on nanophase ceramics. Biomaterials 2000; 21:1803–1810. 28. Wei G., Ma P. X. Nanostructured biomaterials for regeneration. Adv Funct Mater 2008; 18:3568–3582. 29. Loher S., Reboul V., Brunner T. J., Simonet M., Dora C., Neuenschwander P., Stark W. J. Improved degradation and bioactivity of amorphous aerosol derived tricalcium phosphate nanoparticles in poly(lactide-co-glycolide). Nanotechnology 2006; 17(8):2054–2061. 30. Misra S. K., Mohn D., Brunner T. J., Stark W. J., Philip S. E., Roy I., Salih V., Knowles J. C., Boccaccini A. R. Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB)/Bioglass® composites. Biomaterials 2008; 29(12): 1750–1761. 31. Palin E., Liu H. N., Webster T. J. Mimicking the nanofeatures of bone increases boneforming cell adhesion and proliferation. Nanotechnology 2005; 16(9):1828–1835. 32. Ballard J. D., Dulgar-Tulloch A. J., Siegel R. W. Nanophase materials. In Akay M., ed., Wiley Encyclopedia of Biomedical Engineering, vol. 6. New York: John Wiley 2006. 33. Esfahani S. I. R., Tavangarian F., Emadi R. Nanostructured bioactive glass coating on porous hydroxyapatite scaffold for strength enhancement. Mater Lett 2008; 62:3428–3430. 34. Hong Z., Reis R. L., Mano J. F. Preparation and in vitro characterization of novel bioactive glass-ceramic nanoparticles. J Biomed Mater Res A 2009; 88(2):304–313. 35. Hong Z., Reis R. L., Mano J. F. Preparation and in vitro characterization of scaffolds of poly(l-lactic acid) containing bioactive glass-ceramic nanoparticles. Acta Biomaterialia 2008; 4:1297–1306. 36. Liu A., Hong Z., Zhuang X., Chen X., Cui Y., Liu Y., Jing X. Surface modification of bioactive glass nanoparticles and the mechanical and biological properties of poly(Llactide) composites. Acta Biomaterialia 2008; 4(4):1005–1015. 37. Liu A., Wei J., Chen X., Jing X., Cui Y., Liu Y. Novel composites of poly(l-lactide) and surface modified bioactive SiO2-CaO-P2O5 gel nanoparticles: Mechanical and biological properties. Chinese J Polym Sci 2009; 27(3):415–426.



155

38. Hong Z., Liu A., Chen L., Chen X., Jing X. Preparation of bioactive glass-ceramic nanoparticles by combination of sol-gel and coprecipitation method. J Non-Cryst Solids 2009; 355(6):368–372. 39. Hong Z., Merino E. G., Reis R. L., Mano J. F. Novel rice-shaped bioactive glass ceramic nanoparticles. Adv Eng Mater 2009; 11(15):B25–29. 40. Quintero F., Pou J., Comesana R., Lusquinos F., Riveiro A., Mann A. B., Hill R. G., Wu Z. Y., Jones J. R. Laser spinning of bioactive glass nanofibres. Adv Funct Mater 2009; 19:1–7. 41. Zhao N. R., Wang Y. J., Chen X. F., Yang Y. X., Wei K., Wu G. Preparation of bioactive nanoparticles in the system CaO-P2O5-SiO2 using microemulsions. Key Eng Mater 2005; 288–289:179–182. 42. Brunner T. J., Grass R. N., Stark W. J. Glass and bioglass nanopowders by flame synthesis. Chem Commun 2006; 13:1384–1386. 43. Loher S., Stark W. J., Maciejewski M., Baiker A., Pratsinis S. E., Reichardt D., Maspero F., Krumeich F., Günther D. Fluoro-apatite and calcium phosphate nanoparticles by flame synthesis. Chem Mater 2005; 17(1):36–42. 44. Brunner T. J., Stark W. J., Boccaccini A. R., Nanoscale bioactive silicate glasses in biomedical applications, Chapter 6 in Challa S. S. R. Kumar, ed., Nanomaterials for the Life Sciences vol. 2: Nanostructured Oxides, Weinheim, Germany: Wiley-VCH GmbH & Co. KGaA 2009, pp. 203–220. 45. Mohn D., Zehnder M., Imfeld T., Stark W. J. Radio-opaque nanosized bioactive glass for potential root canal application: Evaluation of radiopacity, bioactivity, and alkaline capacity. Int End J 2010; 43:210–217. 46. Jo J. H., Lee E. J., Shin D. S., Kim H. E., Kim H. W., Koh Y. H., Jang J. H. In vitro/in vivo biocompatibility and mechanical properties of bioactive glass nanofibre and poly(e-caprolactone) composite materials. J Biomed Mater Res B: Appl Biomaterials 2009; 91B(1):213–220. 47. Kim H. W., Kim H. E., Knowles J. C. Production and potential of bioactive glass nanofibres as a next-generation biomaterial. Adv Funct Mater 2006; 16:1529– 1535. 48. Kim H. W., Lee H. H., Chun G. S. Bioactivity and osteoblast responses of novel biomedical nanocomposites of bioactive glass nanofibre filled poly(lactic acid). J Biomed Mater Res A 2008; 85A(3):651–663. 49. Fathi M. H., Doostmohammadi A. Bioactive glass nanopowder and bioglass coating for biocompatibility improvement of metallic implant. J Mater Process Tech 2009; 209:1385–1391. 50. Fathi M. H., Mortazavi V., Doostmohammadi A. Bioactive glass nanopowder for the treatment of oral bone defects. J Dentist 2007; 4(3):115–122. 51. Couto D. S., Alves N. M., Mano J. F. Nanostructured multilayer coatings combining chitosan with bioactive glass nanoparticles. J Nanosci Nanotechno 2009; 9:1741–1748. 52. Couto D. S., Hong Z., Mano J. F. Development of bioactive and biodegradable chitosan-based injectable systems containing bioactive glass nanoparticles. Acta Biomater 2009; 5(1):115–123. 53. Kay S., Thapa A., Haberstroh K. M., Webster T. J. Nanostructured polymer/nanophase ceramic composites enhance osteoblast and chondrocyte adhesion. Tissue Eng 2002; 8:753–761. 54. Curtis A. Cell response to nanofeatures in biomaterials. In D. L. D. Silvio, ed., Cellular Response to Biomaterials. London: Woodhead Publishing Limited 2009; p. 429.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

156 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

55. von der Mark K., Park J., Bauer S., Schmuki P., Nanoscale engineering of biomimetic surfaces-cues from the extracellular matrix. Cell Tissue Res 2010; 339:131–153. 56. Engel E., Michiardi A., Navarro M., Lacroix D., Planell J. A., Nanotechnology in regenerative medicine: the materials side. Trends in Biotechnology 2008; 26(1): 39–47. 57. Laurencin C. T., Kumbar S. G., Nukavarapu S. P. Nanotechnology and orthopedics: a personal perspective, Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology 2008; 1(1), pp. 6–10. 58. Mendonça G., Mendonça D. B. S., Aragaõ F. J. L., Cooper L. F. Advancing dental implant surface technology – from micronto nanotopography. Biomaterials 2008; 29:3822–3835. 59. Kennedy S. B., Washburn N. R., Simon Jr C. G., Amis E. J. Combinatorial screen of the effect of surface energy on fibronectin-mediated osteoblast adhesion, spreading and proliferation. Biomaterials 2006; 27:3817–3824. 60. Baier R. E., Meyer A. E., Natiella J. R., Natiella R. R., Carter J. M. Surface properties determine bioadhesive outcomes: methods and results. J Biomed Mater Res 1984; 18:327–355. 61. Zhaoa G., Rainesa A. L., Wielandb M., Schwartza Z., Boyan B. D. Requirement for both micron- and submicron scale structure for synergistic responses of osteoblasts to substrate surface energy and topography. Biomaterials 2007; 28:2821–2829. 62. Charles-Harris M., Navarro M., Engel E., Aparicio C., Ginebra M. P., Planell J. A. Surface characterization of completely degradable composite scaffolds. J Mater Sci: Mater Med 2005; 16:1125. 63. Neo M., Nakamura T., Ohtsuki C., Kokubo T., Yamamuro T. Apatite formation on three kinds of bioactive material at an early stage in vivo: A comparative study by transmission electron microscopy. J Biomed Mater Res 1993; 27:999. 64. Wei J., Chen Q. Z., Stevens M. M., Roether J. A., Boccaccini A. R. Biocompatibility and bioactivity of PDLLA/TiO2 and PDLLA/TiO2/Bioglass® nanocomposites. Mater Sci Eng C 2008; 28(1):1–10. 65. Webster T. J., Ergun C., Doremus R. H., Siegel R. W., Bizios R. Specific proteins mediate enhanced osteoblast adhesion on nanophase ceramics. J Biomed Mater Res 2000; 51(3):475–483. 66. Faghihi S., Azari F., Zhilyaev A. P., Szpunar J. A., Vali H. et al. Cellular and molecular interactions between MC3T3-E1 pre-osteoblasts and nanostructured titanium produced by high-pressure torsion. Biomaterials 2007; 28(27):3887–3895. 67. Lehnert D., Wehrle-Haller B., David C., Weiland U., Ballestrem C., Imhof B. A., Bastmeyer M. Cell behaviour on micropatterned substrata: limits of extracellular matrix geometry for spreading and adhesion. J Cell Sci 2004; 117:41–52. 68. Guarino V., Causa F., Ambrosio L. Bioactive scaffolds for bone and ligament tissue. Expert Rev Medical Devices 2007; 4(3):405–418. 69. Alves N. M., Leonor I. B., Azevedo H. S., Reis R. L., Mano J. F. Designing biomaterials based on biomineralization of bone, submitted. 70. Li R., Clark A. E., Hench L. L. An investigation of bioactive glass powders by sol-gel processing. J App Biomater 1991; 2(4):231–239. 71. Scherrer G. W., Brinker C. J. Sol-gel Science: the Physics and Chemistry of Sol-gel. Boston, MA: Academic Press 1990. 72. Xia, W., Chang, J. Preparation and characterization of nano-bioactive-glasses (NBG) by a quick alkali-mediated sol-gel method. Mater Lett 2007; 61(14–15):3251– 3253.



157

73. Chen X., Lei B., Wang Y., Zhao N. Morphological control and in vitro bioactivity of nanoscale bioactive glasses. J Non-Cryst Solids 2009; 355:791–796. 74. Chen X., Guo C., Zhao N. Preparation and characterization of the sol-gel nanobioactive glasses modified by the coupling agent gamma-aminopropyltriethoxysilane. Appl Surf Sci 2008; 255:466–468. 75. Delben J. R. J., Pimentel O. M., Coelho M. B., Candelorio P. D., Furini L. N., Santos F. A., Vicente F. S., Delben A. A. S. T. Synthesis and thermal properties of nanoparticles of bioactive glasses containing silver. J Therm Anal Calorim 2009; 97:433–436. 76. Kim H. W., Kim H. E., Knowles J. C. Production and potential of bioactive glass nanofibres as a next-generation biomaterial. Adv Func Mater 2006; 16(12): 1529–1535. 77. Xia W., Zhang D. M., Chang J. Fabrication and in vitro biomineralization of bioactive glass (BG) nanofibres. Nanotechnology 2007; 18:135601(7 pp.). 78. Lu H., Zhang T., Wang X. P., Fang Q. F. Electrospun submicron bioactive glass fibres for bone tissue scaffold. J Mater Sci: Mater Med 2009; 20:793–798. 79. Arriagada F. J. Synthesis of nanosize silica in a nonionic water-in-oil microemulsion. J Colloid Interf Sci 1999; 211:210–220. 80. Lim G. K., Wang J., Ng S. C., Gan L. M. Processing of fine hydroxyapatite powders via an inverse microemulsion route. Mater Lett 1996; 28:431–436. 81. Paul B. K., Moulik S. P. Microemulsions: An overview. J Disper Sci Technol 1997; 18(4):301–367. 82. Pratsinis S. E. Flame aerosol synthesis of ceramic powders. Prog Energ Combust Sci 1998; 24(3):197–219. 83. Stark W. J., Madler L., Maciejewski M., Pratsinis S. E., Baiker A. Flame synthesis of nanocrystalline ceria-zirconia: effect of carrier liquid. Chem Commun 2003; 5: 588–589. 84. Stark W. J., Mädler L., Pratsinis S. E. Metal oxides prepared by flame spray pyrolysis. WO 2004/005184, 2004. 85. Stark W. J., Pratsinis S. E. Metal delivery system for nanoparticle manufacture. WO 2004/103900A1, 2004. 86. Quintero F., Mann A. B., Pou J., Lusquinos F., Riveiro A. Appl Phys Lett 2007; 90:153109. 87. Quintero F., Pou J., Lusquinos F., Riveiro A. Experimental analysis of the production of micro- and nanofibres by laser spinning. Appl Surf Sci 2007; 254:1042–1047. 88. Quintero F., Dieste O., Pou J., Lusquiños F., Riveiro A. On the conditions to produce micro- and nanofibres by laser spinning. J Phys D: Appl Phys 2009; 42:065501 (10 pp). 89. Koo, J. H. Polymer Nanocomposites. USA: McGraw-Hill 2006. 90. Vollenweider M., Brunner T. J., Knecht S., Grass R. N., Zehnder M., Imfeld T., Stark W. J. Remineralization of human dentin using ultrafine bioactive glass particles. Acta Biomaterialia 2007; 3(6):936–943. 91. Misra S. K., Ansari T., Mohn D., Valappil S. P., Brunner T. J., Stark W. J., Roy I., Knowles J. C., Sibbons P. D., Jones E. V., Boccaccini A. R., Salih V. Effect of nanoparticulate bioactive glass particles on bioactivity and cytocompatibility of poly(3-hydroxybutyrate) composites. J Royal Soc Interface 2010; 7:453–465. 92. Zheng H., Yingjun W., Chunrong Y., Xiaofeng C., Naru Z. Investigation on the Porous biomaterial for bone reconstruction with addition of bio-mimetic nano-sized inorganic particles. Key Eng Mater 2007; 336–338:1534–1537.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

158 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

93. El-Kady A. M., Ali A. F., Farag M. M. Development, characterization, and in vitro bioactivity studies of sol-gel bioactive glass/poly(L-lactide) nanocomposite scaffolds. Mater Sci Eng C 2010; 30(1):120–131. 94. Peter M., Kumar P. T. S., Binulal N. S., Nair S. V., Tamura H., Jayakumar R. Development of novel a-chitin/nanobioactive glass-ceramic composite scaffolds for tissue engineering applications. Carbohyd Polym 2009; 78(4):926–931. 95. Peter M., Binulal N. S., Soumya S., Nair S. V., Furuike T., Tamura H., Jayakumar R. Nanocomposite scaffolds of bioactive glass-ceramic nanoparticles disseminated chitosan matrix for tissue engineering applications. Carbohyd Polym 2010; 79(2):284–289. 96. Peter M., Binulal N. S., Nair S. V., Selvamurugan N., Tamura H., Jayakumar R. Novel biodegradable chitosan–gelatin/nano-bioactive glass-ceramic composite scaffolds for alveolar bone tissue engineering. Chem Eng J 2010; 158:353–361. 97. Wang Y., Yang C., Chen X., Zhao N. Development and characterization of novel biomimetic composite scaffolds based on bioglass-collagen-hyaluronicacidphosphatidylserine for tissue engineering applications. Macromol Mater Eng 2006; 291:254–262. 98. Xie E., Hu Y., Chen X., Bai J., Ren L., Zhang Z. A novel nanocomposite and its application in repairing bone defects. Proceedings of the 3rd IEEE Int Conf on Nano/ Micro Engineered and Molecular Systems, Sanya, China, 2008; pp. 943–946. 99. Luz G. M., Mano J. F. New composite membranes containing bioactive glassceramic nanoparticles and chitosan for biomedical applications. Mater Sci Forum 2010; 636–637:31–35. 100. Noh K. T., Lee H. Y., Shin U. S., Kim H. W. Composite nanofibre of bioactive glass nanofiller incorporated poly(lactic acid) for bone regeneration. Mater Lett 2010; 64:802–805. 101. Kim H. W., Song J. H., Kim H. E. Bioactive glass nanofibre–collagen nanocomposite as a novel bone regeneration matrix. J Biomed Mater Res A 2006; 79(3):698–705. 102. Lee H. H., Yu H. S., Jang J. H., Kim H. W. Bioactivity improvement of poly(ecaprolactone) membrane with the addition of nanofibrous bioactive glass. Acta Biomaterialia 2008; 4:622–629. 103. Misra S. K., Ansari T. I., Valappil S. P., Mohne D., Philip S. E., Stark W. J., Roy I., Knowles J. C., Salih V., Boccaccini A. R. Poly(3-hydroxybutyrate) multifunctional composite scaffolds for tissue engineering applications. Biomaterials 2010; 31: 2806–2815. 104. Mouriño V., Boccaccini A. R. Bone tissue engineering therapeutics: controlled drug delivery in three-dimensional scaffolds. J R Soc Interface 2010; 6, 7(43):209–227. 105. Forsback A. P., Areva S., Salonen J. I.. Mineralization of dentin induced by treatment with bioactive glass S53P4 in vitro. Acta Odont Scandinavica 2004; 62:14–20. 106. Stoor P., Soderling E., Salonen J. I. Antibacterial effects of a bioactive glass paste on oral microorganisms. Acta Odont Scandinavica 1998; 56:161–165. 107. Sepulveda P., Jones J. R., Hench L. L. In vitro dissolution of melt-derived 45S5 and sol-gel derived 58S bioactive glasses. J Biomed Mater Res 2002; 61:301–311. 108. Waltimo T., Brunner T. J., Vollenweider M., Stark W. J., Zehnder M. Antimicrobial effect of nanometric bioactive glass 45S5. J Dent Res 2007; 86(8):754–757. 109. Kernen F., Waltimoc T., Deyhlea H., Beckmannd F., Starke W., Müller B. Synchrotron radiation-based micro computed tomography in the assessment of dentin de- and re-mineralization. In Stuart R., ed., Stock Developments in X-Ray Tomography VI, Proc of SPIE: 7078: 70780M, 2008.



159

110. Allan I., Newman H., Wilson M. Antibacterial activity of particulate Bioglass® against supra- and subgingival bacteria. Biomaterials 2001; 22:1683–1687. 111. Bellantone M., Williams H. D., Hench L. L. Broad-spectrum bactericidal activity of Ag2O-doped bioactive glass. Antimicrob Agents Chemother 2002; 46:1940–1945. 112. Vitale-Brovarone C., Miola M., Balagna C., Verne E. 3D-glass-ceramic scaffolds with antibacterial properties for bone grafting. Chemical Eng J 2008; 137:129–136. 113. Haimi S., Gorianc G., Moimas L., Lindroos B., Huthala H., Räty S., Kuokkanen H., Sándor G. K., Schmid C., Miettinen S., Suuronen R. Characterization of zincreleasing three-dimensional bioactive glass scaffolds and their effect on human adipose stem cell proliferation and osteogenic differentiation. Acta Biomater 2009; 5:3122–3131 114. Gubler M., Brunner T. J., Zehnder M., Waltimo T., Sener B., Stark W. J. Do bioactive glasses convey a disinfecting mechanism beyond a mere increase in pH? Inter Endodon J 2008; 41:670–678. 115. Nair P. N. R. On the causes of persistent apical periodontitis: a review. Inter Endodon J 2006; 39:249–281. 116. Mortazavi V., Nahrkhalaji M. M., Fathi M. H., Mousavi S. B., Esfahani B. N. Antibacterial effects of sol-gel-derived bioactive glass nanoparticle on aerobic bacteria. J Biomed Mater Res A 2010; 94A(1):160–168. 117. Cannillo V., Sola A. Different approaches to produce coatings with bioactive glasses: Enamelling vs plasma spraying. J Euro Ceram Soc 2010; 30:2031–2039. 118. Verné E., Ferraris M., Ventrella A., Paracchini L., Krajewski A., Ravaglioli A. Sintering and plasma spray deposition of bioactive glass–matrix composites for medical applications. J Euro Ceram Soc 1998; 18:363–372. 119. Wang C. X., Chen Z. Q., Wang M. Fabrication and characterization of bioactive glass coatings produced by the ion beam sputter deposition technique. J Mater Sci: Mater Med 2002; 13:247–251. 120. Durán A., Conde A., Gómez-Coedo A., Dorado T., García C., Ceré S. Sol-gel coatings for protection and bioactivation of metals used in orthopaedic devices. J Mater Chem 2004; 14:2282–2290. 121. Liste S., Serra J., Gonzalez P., Borrajo J. P., Chiussi S., Leon B., Pérez-Amor M. The role of the reactive atmosphere in pulsed laser deposition of bioactive glass films. Thin Solid Films 2004; 453–454:224–228. 122. Boccaccini A. R., Peters C., Roether J. A., Eifler D., Misra S. K., Minay E. J., Electrophoretic deposition of polyetheretherketone (PEEK) and PEEK/Bioglass® coatings on NiTi shape memory alloy wires. J Mater Sci 2006; 41:8152–8159. 123. Schausten M. C., Meng D., Telle R., Boccaccini A. R. Electrophoretic deposition of carbon nanotubes and bioactive glass particles for bioactive composite coatings. Ceram Inter 2010; 36:307–312. 124. Pishbin F., Simchi A., Ryana M. P., Boccaccini A. R. A study of the electrophoretic deposition of Bioglass® suspensions using the Taguchi experimental design approach. J Euro Ceram Soc in press. 125. Comesaña R., Quintero F., Lusquiños F., Pascual M. J., Boutinguiza M., Durán A., Pou J. Laser cladding of bioactive glass coatings. Acta Biomater 2010; 6:953–961. 126. Dorozhkin S. V. Bioceramics of calcium orthophosphates Biomaterials 2010; 31:1465–1485. 127. Moszner N., Klapdohr S. Nanotechnology for dental composites. Int J Nanotechnology 2004; 1:130–156.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

160 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

128. Zhang C., Li C., Huang S., Hou Z., Cheng Z., Yang P., Peng C., Lin J. Self-activated luminescent and mesoporous strontium hydroxyapatite nanorods for drug delivery. Biomaterials 2010; 31:3374–3383. 129. Slowing I. I., Trewyn B. G., Giri S., Lin V. S. Y. Mesoporous silica nanoparticles for drug delivery and biosensing applications. Adv Funct Mater 2007; 17:1225–1236. 130. Ladron de G. F. S., Ragel C. V., Vallet-Regi M. Bioactive glass-polymer materials for controlled release of ibuprofen. Biomaterials 2003; 24:4037–4043. 131. Domingues Z. R., Cortes M. E., Gomes T. A., Diniz H. F., Freitas C. S., Gomes J. B., Faria A. M. C., Sinisterra R. D. Bioactive glass as a drug delivery system of tetracycline and tetracycline associated with betacyclodextrin. Biomaterials 2004; 25:327–333. 132. Leach J. K., Kaigler D., Wang Z., Krebsbach P. H., Mooney D. J. Coating of VEGFreleasing scaffolds with bioactive glass for angiogenesis and bone regeneration. Biomaterials 2006; 27:3249–3255. 133. Bergeron E., Marquis M. E., Chrétien I., Faucheux N. Differentiation of preosteoblasts using a delivery system with BMPs and bioactive glass microspheres. J Mater Sci: Mater Med 2007; 18:255–263. 134. Li S., Nguyen L., Xiong H., Wang M., Hu T. C., She J. X., Serkiz S. M., Wicks G. G., Dynan W. S. Porous-wall hollow glass microspheres as novel potential nanocarriers for biomedical applications. Nanomedicine: Nanotechnology, Biology, and Medicine 2010; 6:127–136. 135. Wu C., Ramaswamy Y., Zhu Y., Zheng R., Appleyard R., Howard A., Zreiqat H. The effect of mesoporous bioactive glass on the physiochemical, biological and drugrelease properties of poly(DL-lactide-co-glycolide) films. Biomaterials 2009; 30:2199–2208. 136. Chen Q. Z., Rezwan K., Armitage D., Nazhat S. N., Boccaccini A. R. The surface functionalization of 45S5 Bioglass-based, glass-ceramic scaffolds and its impact on bioactivity, J Mater Sci: Mater Med 2006; 17:979–987. 137. Xynos I. D., Edgar A. J., Buttery L. D. K., Hench L. L., Polak J. M. Ionic products of bioactive glass dissolution increase proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression and protein synthesis. Biochem Bioph Res Co 2000; 276:461–465. 138. Xynos I. D., Edgar A. J., Buttery L. D., Hench L. L. Gene-expression profiling of human osteoblasts following treatment with the ionic products of Bioglass® 45S5 dissolution. J Biomed Mater Res 2001; 55:151–157. 139. Hench L. L. Genetic design of bioactive glass. J Euro Ceram Soc 2009; 29:1257– 1265. 140. Vallet-Regi M., Balas F., Arcos D. Mesoporous materials for drug delivery, Angewandte Chemie International Edition 2007; 46:7548–7558. 141. Desai M. P., Labhasetwar V., Walter E., Levy R. J., Amidon G. L. The mechanism of uptake of biodegradable microparticles in Caco-2 cells is size dependent. Pharm Res Adv 1997; 1568–1573. 142. Desai M. P., Labhasetwar V., Amidon G. L., Levy R. J. Gastrointestinal uptake of biodegradable microparticles: effect of particle size. Pharm Res 1996; 13:1838–1845. 143. Vinagradov S. V., Bronich T. K., Kabanov A. V. Nanosized cationic hydrogels for drug delivery: preparation, properties and interactions with cells. Adv Drug Del Rev 2002; 54:223–233. 144. Panyama J., Labhasetwar V. Biodegradable nanoparticles for drug and gene delivery to cells and tissue. Adv Drug Del Rev 2003; 55:329–347.



161

145. Witasp E., Kupferschmidt N., Bengtsson L., Hultenby K., Smedman C., Paulie S., Garcia-Bennett A. E., Fadee B. Efficient internalization of mesoporous silica particles of different sizes by primary human macrophages without impairment of macrophage clearance of apoptotic or antibody-opsonized target cells. Toxicol Appl Pharmacol 2009; 239:306–319. 146. Yun H., Kim S., Lee S., Song I. Synthesis of high surface area mesoporous bioactive glass nanospheres. Mater Lett in press. 147. Kukowska-Latallo J. F., Candido K. A., Cao Z., Nigavekar S. S., Majoros I. J., Thomas T. P., Balogh L. P., Khan M. K., Baker J. R. Jr. Nanoparticle targeting of anticancer drug improves therapeutic response in animal model of human epithelial cancer. Cancer Res 2005; 65:5317–5324. 148. Bretcanu O., Spriano S., Vitale C. B., Verne E. Synthesis and characterization of coprecipitation-derived ferrimagnetic glass-ceramic. J Mater Sci 2006; 41: 1029–1037. 149. Wang T. W., Wu H., Wang W. R., Lin F. H., Lou P. J., Shieh M. J., Young T. H. The development of magnetic degradable DP-Bioglass for hyperthermia cancer therapy. J Biomed Mater Res A 2007; 83A(3):828–837. 150. Roy I., Gupta M. N. Smart polymeric materials: emerging biochemical applications. Chem Biol 2003; 10:1161–1171. 151. Phong A. T., Sarin L., Hurtb R. H., Webster T. J. Opportunities for nanotechnologyenabled bioactive bone implants. J Mater Chem 2009; 19:2653–2659.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

7 Bioactive glass containing composites for bone and musculoskeletal tissue engineering scaffolds S. VERRIER, AO Research Institute Davos, Switzerland, J. E . GOUGH, University of Manchester, UK and A. R. BOCCACCINI , University of Erlangen-Nuremberg, Germany and Imperial College London, UK

Abstract: Composites developed by combining bioactive glasses and biodegradable polymers are attractive materials for use in scaffolds for musculoskeletal tissue engineering due to their ability to be tailored for different applications. Scaffolds must be osteoconductive and osteoinductive in order to guide and encourage new bone formation. Additionally, neo-vascularization of the construct is required. This chapter focuses on bioactive glass-containing composites for tissue engineering, with emphasis on the in vitro and in vivo performance of these scaffolds. A wide range of cell types (primary or cell lines) from different origins have been considered for studies in vitro. Several studies have also been carried out in specific animal models (in vivo). Recent studies report the angiogenic potential of such composites in both cases. The evidence indicates that these composite materials are promising and of high interest for bone tissue engineering and musculoskeletal tissue regeneration. However, in vitro and in vivo understanding of these scaffolds is still limited, especially regarding their long-term degradation and ion release effects on the biological environment. Key words: bioactive glasses, composites, scaffolds, tissue engineering, in vitro, in vivo, angiogenesis, musculoskeletal tissue.

7.1

Introduction

One of the key areas within tissue engineering (TE) and regenerative medicine gaining increasing attention is related to bone and musculoskeletal tissue engineering and regeneration [1]. It is well known that critical size bone defects that occur due to trauma or disease are very difficult to repair via the natural growth of host tissue. A common TE approach to restore function to diseased or damaged bone tissue is to design combinations of functional cells and biodegradable scaffolds made from engineered biomaterials [1–3]. In musculoeskeletal tissue regeneration, for example intervertebral disc regeneration, scaffolds made of suitable engineered biomaterials are also being investigated in combination with relevant cells to tackle the increasing medical need in this area [4]. In common tissue engineering strategies, scaffolds are highly porous biomaterial engineered structures that serve as temporary 3D templates for cell adhesion, 162 © Woodhead Publishing Limited, 2011

Bioactive glasses for tissue engineering scaffolds

163

proliferation, migration, and ultimately for supporting the formation of new tissue [2, 3]. The viability of this approach for growing new functional bone depends not only on the osteogenesis potential of the construct but also on the ability to induce rapid vascularization upon implantation [5, 6]. The geometry of the biomaterial scaffold is very important for the success of this strategy: when cells have attached to the surface there must be enough space and open channels to allow for nutrient ingress, waste delivery, protein transport and vascular growth to occur, functions that are obtainable with an interconnected network of pores. Suitable biomaterials for the development of bone TE scaffolds are those exhibiting bioactive properties [1]. Bioactive materials show high surface reactivity, leading to strong chemical interactions with relevant physiological fluids that induce the development of tenacious bonds to bone through the biological interaction of collagen with the material surface. In this way, bioactive materials can transfer loads to and from living bone. The most investigated bioactive materials for bone TE are bioceramics, including specific compositions of silicate glasses and glass-ceramics, as well as hydroxyapatite (HA) and related calcium phosphate ceramics [7]. Silicate bioactive glasses were first developed by Hench and co-workers 40 years ago [8]. The first bioactive glass composition, universally known as 45S5 Bioglass® (BG) (composition in wt%: 45% SiO2, 24.5% Na2O, 24.5% CaO and 6% P2O5), has the approval of the US Food and Drug Administration (FDA). This glass has been successfully used, for example, in clinical treatments of periodontal diseases as bone filler as well as in middle ear surgery [7]. Other clinical applications of bioactive glasses have been proposed, for example in dentistry [9] or as coating on metallic orthopaedic implants [10, 11]. In the last 40 years tens of different bioactive glass compositions for different biomedical applications have been developed, and several traditional and modern applications of bioactive glasses are described in other chapters of this book. The common characteristic of these glasses is their high surface reactivity and their ability to bond to bone in a physiological environment [7]. More recently, the application of bioactive glasses in bone TE and regenerative medicine has received marked impulse from the TE research community [12–15] and it can be stated that bone TE represents one of the most exciting future clinical applications of bioactive glasses, e.g. to fabricate optimal scaffolds with osteogenic and angiogenic potential [16]. In the context of bone TE, bioactive silicate glasses have several attractive advantages in comparison to other bioactive ceramics, e.g. sintered hydroxyapatite. One important finding in 2000 [17] is related to the effect of dissolution products from bioactive glasses on the upregulation of the expression of genes that control osteogenesis, which is being actively investigated [16, 18–20]. Further in vivo and in vitro studies using 45S5 BG particles have shown encouraging results regarding the potential angiogenic effects of Bioglass®, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularization in vivo [21–23]. Bioactive glasses can also serve as vehicle for the controlled


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

164 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

delivery of selected ions, which can act to control specific cell functions. In this context, bioactive glass compositions incorporating additional elements in the silicate network such as magnesium, strontium, boron, iron, silver, potassium or zinc are being developed [24–28]. Like most inorganic materials, the major disadvantage of bioactive glasses is their limited fracture strength and low fracture toughness (i.e. high brittleness). This problem can be tackled by combining bioactive glasses with biopolymers forming bioactive composite materials [29]. In the most common approach, bioactive glass particles or fibres are incorporated in biodegradable polymer matrices [29–32], this being the type of composite to be discussed in the present chapter, as illustrated in Fig. 7.1. Other possibilities include the coating of (porous) polymer scaffolds with bioactive glass particles [33] or the coating and impregnation of 3D bioactive glass scaffolds with biodegradable polymers (discussed in Chapter 5 of this book). Since the requirements for optimal scaffolds are manifold [34], it can be stated that a combination of degradable polymers and bioactive glasses represents an optimal approach in terms of achievable mechanical and biological performance towards improved TE scaffolds [29, 30]. A wide range of biodegradable synthetic polymers is available for developing composite scaffolds in combination with bioactive glasses [35, 36]. For example, poly(lactic-acid) (PLA), poly(glycolic-acid) (PGA) and their co-polymers poly(lactid-co-glycolic-acid) (PLGA) [37, 38] are being highly investigated. These polymers have extensive FDA approval history. Also

7.1 Schematic diagram showing combinations of biodegradable polymers and bioactive glasses to form composites [46].



165

polyhydroxyalkanoates (PHAs), a family of microbial polyesters, are increasingly finding application for composite scaffolds development, in particular poly(3hydroxybutyrate) and poly(3-hydroxybutyrate-co-hydroxyvalerate) [39, 40]. PCL is another popular polymer being combined with bioactive glasses for composite scaffolds [41]. In addition, composite scaffolds are also being developed combining natural polymers, including collagen, chitosan, gelatine or alginate, with bioactive glasses (e.g. [42, 43]). For bone tissue engineering, porosity of ∼90% and pore size >100 µm are desirable, as well as high pore interconnectivity, in order to facilitate the attachment and proliferation of cells and the ingrowth of new tissue into the scaffold, as well as to enable mass transport of oxygen, nutrition and waste products [34, 44]. This chapter will discuss the materials science and technology of composites based on the combination of biodegradable polymers (specifically synthetic polymers) and bioactive glass particles for bone tissue engineering scaffolds. To demonstrate the versatility of the materials in musculoskeletal tissue engineering, the development and application of polymer/Bioglass® scaffolds for intervertebral disc regeneration will be also presented. In Section 7.2 the basics of the composite materials approach to tissue engineering scaffolds are described, including a discussion of typical composite scaffold fabrication technologies, scaffold microstructure and relevant properties achieved. Section 7.3 discusses the specific results of the in vitro and in vivo applications of these scaffolds in relation to bone regeneration and intervetebral disc regeneration, while Section 7.4 presents a discussion of the latest developments on cell/tissue response of bioactive composite scaffolds, highlighting the effect of the presence of bioactive glass particles within polymer matrices on cell growth and differentiation. The chapter finishes with a summary and the scope for future developments in the field (Section 7.5).

7.2

Composite materials approach to tissue engineering scaffolds

7.2.1 Advantages of composite materials Composite materials are made by combining two or more chemically distinct materials (metallic, ceramic, or polymeric) in the micron or nanoscale, which are separated by an interface. Fillers of different morphologies can be used to fabricate composite materials with biopolymer matrices, such as particulates, short fibres, continuous fibres and nanofillers (e.g. nanoparticles, nanofibres) [45] (Fig. 7.1). For applications in tissue engineering scaffolds, composites must exhibit a set of desired properties, such as adequate mechanical strength, tailored initial elastic modulus close to the elastic modulus of the tissue of concern and controlled degradation behaviour in vivo. Composite scaffolds must degrade at a predefined rate and must retain their structural integrity in vivo providing the necessary


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

166 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

mechanical support for cell attachment and proliferation. Polymers are generally flexible exhibiting relatively low compression strength and stiffness, being in cases too flexible to meet the mechanical demands in surgery and in in vivo situations. On the other hand, bioactive glasses are too stiff and brittle compared to bone. Several reasons have been put forward for suggesting the combination of biodegradable polymers and bioactive glasses for tissue engineering applications [29–32]. In addition to the possibility of increasing the compression strength and stiffness of polymer scaffolds by incorporating bioactive glass inclusions, the availability of effective processing methods for porous polymer structures can also be exploited. For example, polymers can be easily fabricated into complex shapes and porous structures by a variety of well-established techniques [46]. Moreover, the addition of bioactive phases to bioresorbable polymers can be used to control polymer degradation behaviour, e.g. by buffering the pH of the nearby solution and hence controlling the acidic degradation of the polymer, in particular in case of polylactic acid. Here, dissolution products from bioactive glasses can alter the autocatalytic effect of the acidic end groups resulting from hydrolysis of the polymer chains. In addition, bioactive glass inclusions contribute to water absorption in the scaffold due to the increased number of internal interfaces formed between the polymer and the hydrophilic bioactive glass particles [47]. The incorporation of bioactive glass into a biodegradable polymer matrix will allow the composite to interact with the surrounding bone tissue inducing the formation of a strong bond with bone via the growth of a carbonate hydroxyapatite layer, as mentioned above. Composite materials represent thus a convenient alternative for fabricating tissue engineering scaffolds considering that their properties can be engineered to suit the mechanical and physiological demands of the host tissue, which can be controlled by varying the volume fraction, morphology and arrangement of the bioactive glass inclusions [29, 31, 32]. Two most commonly used bioactive glass inclusions in biopolymers for biomedical composites are fibres and particulates [29]. With the availability of bioactive glass nanoparticles [48] (see also Chapter 6) nanocomposites are also being considered for bone tissue engineering [49]. It has been shown that increased volume fraction and higher surface area-to-volume ratio of inclusions favour bioactivity [29]. For certain applications the incorporation of (nano)fibres is preferred to particles [50]. In addition, composite mechanical properties are determined by the inclusion shape and size as well as by the quality of the distribution of the particles or fibres in the matrix. Moreover, as in all composite systems, the reinforcement-matrix interfacial bonding is of major importance in influencing the final composite properties. In the case of highly porous scaffolds, particularly of interest for bone tissue engineering, is the scaffold porosity, e.g. pore volume, size, shape, orientation and connectivity, that strongly affects the mechanical properties and structural integrity of scaffolds. The porosity structure is related mainly to the fabrication method used, as discussed next.



167

7.2.2 Processing methods Tissue engineering scaffolds must mimic the numerous functions of the natural extracellular matrix, including providing support for cell adhesion and migration, and organizing cells into 3D structure [34]. Scaffolds for bone and muscoloeskeletal tissue engineering are highly porous, thus porosity and pore structure significantly affect the characteristics of a scaffold [29, 51]. A minimum pore size is required for tissue ingrowth, and high 3D interconnectivity is necessary for access of nutrients, transport of waste products, better cell spreading and vascularization [29–31]. The pore structure and properties of scaffolds are dictated by the manufacturing process employed. There has been a broad range of scaffold manufacturing techniques developed recently, which can be applied for biopolymer/bioactive glass composites, including well known methods such as the use of porogens, chemical segregation, microsphere sintering, solvent casting, particulate (salt or sugar) leaching and thermally induced phase separation (TIPS), as well as a series of computer-assisted rapid prototyping techniques, such as three-dimensional printing and fused deposition modelling. In addition, electrospinning is being increasingly considered for use in forming nanofibrous structures. Each of the developed techniques has the ability to produce scaffolds with a different pore architecture, but they also have limitations with respect to specific properties than can be achieved, as discussed elsewhere [29, 34, 52]. Solvent casting with and without particle leaching [53, 54], thermally induced phase separation (TIPS) [55, 56] and solid free form fabrication methods [2, 57, 58] have been applied successfully to manufacture synthetic biopolymer-ceramic composite scaffolds. Solvent casting for production of composite scaffolds involves dissolving the polymer in an organic solvent, mixing with bioactive glass particles and casting the solution into a predefined 3D mould [54]. The solvent is subsequently allowed to evaporate. This is a simple processing technique not requiring specialized or expensive equipment. The disadvantages of the method include poor pore interconnectivity especially at low porosities and a difficulty in generating large (3D) structures (over 3 mm thick). Composite constructs are also fabricated by combining solvent casting, particle leaching and microsphere packing methods [54]. Polymer microspheres are first formed from traditional water oil/water emulsions, and scaffolds are then developed by mixing solvent, salt particles (porogens), bioactive glass particles and pre-hardened microspheres [59]. The sintered scaffolds present a wellintegrated microstructure with a porosity of 40%. The mechanical properties of these composites were found to be similar to those of cancellous bone. Biodegradable polymer scaffolds with very high porosities (∼97%) can be produced using the thermally induced phase separation (TIPS) technique [31]. The obtained scaffolds exhibit pores with anisotropic tubular morphology and high pore interconnectivity. The TIPS process has been used to produce composite


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

168 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

scaffolds based on PLGA and PDLLA foams containing Bioglass® particles [31, 47, 60]. Micrographs showing the pore structure of PDLLA and PDLLA/ Bioglass® scaffolds developed by the TIPS method are shown in Fig. 7.2 [60]. Scaffolds obtained by TIPS usually exhibit oriented tubular pores of diameters of several hundred microns (>100 µm) and an isotropic pore network of smaller pore size (∼10 µm) connecting the large tubular pores. Solid freeform fabrication (SFF) techniques including 3D printing, selective laser sintering, multi-phase jet solidification, and fused deposition modelling (FDM) have been developed to manufacture tissue scaffolds for bone tissue engineering with specific structure and properties [61, 62]. SFF allows a high degree of pore interconnectivity and controlled morphology. The methods enable the fabrication of tailored scaffolds by incorporating patient-specific information as well as a designed microenvironment. Tissue geometry can be extracted from patient’s computed tomography (CT) or magnetic resonance imaging (MRI) data and reconstructed as a 3D model. However, a shortcoming of these methods is increased fabrication time and the requirement of relatively complex equipment compared with the methods described above.

7.2 Scanning electron microscopy images of the transversal section of TIPS foams, showing the typical homogeneous regions of (a) pure PDLLA foam, (b) PDLLA/2 vol% Bioglass® foam, (c) PDLLA/15 vol% Bioglass® foam (reproduced from [60] with permission from Elsevier).



169

7.2.3 Mechanical properties Bioactive glass inclusions incorporated in biodegradable polymer matrices positively affect the mechanical properties leading to reinforcement of the scaffold structure. The enhancement of mechanical properties (compression strength, elastic constants) depends strongly on particle shape and size distribution; as well as on the quality of the inclusion distribution in the matrix and, crucially, on the strength of the inclusion-matrix interface. Current porous scaffolds incorporating bioactive glass particles (size powder > PCL for cell viability

[41]

Human osteosarcoma cell line (MG63)

10% BG => less growth than for 0%BG. Material colonization

[80]

Human osteosarcoma cell line (MG63) Human osteosarcoma cell line (MG63)

Cytocompatibility, BG contents dependent (10%>20%>30%) Negative effect of 20% BG on cell growth, positive effect of VitE + BG

[63]

Human osteosarcoma cell line (HOS-TE85)

BG scaffold – 2 porosities – cell growth on composite > neat BG

[82]

P(3HB)


[81]


171

Table 7.1 Continued Polymer matrix

ECM comp

Cell type

Outcomes/remarks

References

Human osteosarcoma cell line (MG63)

Composite films – cell attachment, proliferation and differentiation

[83]

Human MSC

BG < BG + Coll +/− HYA PCL alone and empty defect

P(3HB)

Rat-subcutaneous

Thin encapsulation, fibroblasts growth, collagen, newly formed capillaries

[80]

ECMcomp

Rat calvaria

Collagen sponges – higher vascularization, higher faster bone formation compared to neat collagen

[100]

[79]

Notes: BG: Bioactive glass PLGA: poly-(D/L-lactic-co-glycolic) acid PGA: polyglycolic acid PLA: polylactic acid PLLA: poly-(L-lactic) acid PDLLA: poly-(D/L-lactic) acid PHB: polyhydroxybutyrate PCL: poly-caprolactone BMSC: Bone marrow stromal cells OB: osteoblast ECM: Extra cellular matrix VEGF: Vascular endothelial growth factor ECM comp: Extra cellular matrix component (e.g. type I collagen, hyaluronic acid, demineralized bone matrix) PS: phosphatidylserine P(3HB): poly(3-hydroxybutyrate).

Bioglass® (type 45S5) is the most investigated silicate bioactive glass system for bone tissue engineering [21, 33, 67–69, 71–77, 85–87] and, as shown in Tables 7.1 and 7.2, the majority of the composites comprising Bioglass® are prepared with either PLGA or PDLLA biodegradable polymer matrices. Several studies have investigated the biological response of different cell types to composite materials made of polylactide-co-glycolide and bioactive glass. For example, Lu et al. [71] found extensive growth of SaOS-2 (human osteosarcoma cell line) at the surface and within the pores of PLGA/BG composites in comparison to neat PLGA and TCP scaffolds. Higher ALP activity and collagen I synthesis were also observed. Lu et al. [72] also described a dose–response effect according to BG content, for both physico-chemical characteristics and cellular



173

response points of view. After 7 days of culture on PLGA, PLGA/BG 10 wt% or PLGA/BG 50 wt%, SaOS-2 cell proliferation was found to be higher on the wt10% material compared to the 0 or 50 wt% BG composites. Although similar ALP activity was found in the 0 and 10 wt% BG materials, mineralization was higher for the PLGA/BG 10 wt% scaffold. Similar observations were made by Yao et al. [68], using rat MSC. Further studies [74–77] using PDLLA scaffolds also showed a dose-related influence of the bioactive glass filler on cellular behaviour. Tsigkou et al. [76], for example, investigated differentiation of human foetal osteoblast cultured on PDLLA, PDLLA containing 5 wt% BG and 40 wt% BG. Despite the fact that the cultures were performed in absence of osteogenic factor supplements, significantly higher cell maturation and differentiation on the composite materials were measured when compared to the cell culture plastic (control) and PDLLA alone. Moreover, even if significant differences were observed between cells grown on materials containing 5 wt% or 40 wt% BG (ALP activity and osteoblastic gene expression), the osteocalcin secretion and matrix mineralization were found to be higher on PDLAA/BG 5 wt% compared to PDLLA/BG 40 wt%. Using comparable scaffolds (PDLLA alone, PDLLA/BG5 wt% and PDLLA/ BG40 wt%), Yang et al. [75] found very similar results using human MSC or Stro-1 positive fractions of MSC. Namely, significant increase of ALP activity was determined for the material containing 5 wt% BG compared to the neat scaffold or the 40 wt% BG scaffold. As for the previous mentioned study, it was shown that a BG high content has a negative effect on the biological response compared to a lower amount (40 wt% v. 5 wt%). But looking at the influence of pre-treatment in medium containing 20% Fetal Calf Serum (FCS), Yang et al. [75] came to the conclusion that this might be due to the higher and extended ion release in the culture medium in the case of higher amounts of Bioglass® present. In the in vivo part of their study however, the materials were not pre-treated before implantation, and no significant differences were observed whether PDLLA contained BG or not. More recently, a few studies reported the use of polycaprolactone (PCL) associated to bioactive glass particulates [41, 79]. In their work using MC3T3-E1 cell line (mouse pre-osteoblast), Jo et al. [41] showed a significant higher number of cells on the composite containing bioactive glass nanofibres compared to the PCL only or containing bioactive glass powder. The same trend was observed when the composite materials were implanted in a rat calvarial defect. The same composite (PCL/BG) was also used by Erdemli et al. [79], to which a third component, namely bone demineralized matrix (BDM) or calcium sulphate (CS), was added. From the in vitro and in vivo studies carried out, no major beneficial effect of these new scaffolds on Saos-2 cell proliferation was found, however in their in vivo experiment, the authors observed a better tendency of bone ingrowth when composites supplemented with BMD were implanted [79]. Another group of polymers, poly(3-hydroxybutyrate) (P(3HB)) has recently gained the attention of materials scientists interested in bone tissue engineering


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

174 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

[39]. P(3HB) belongs to the family of polyhydroxyalkanoate, which exhibits a much longer degradation time compared to other polymers groups such as for example PLA or PLGA. Chen and co-workers [88] already studied the potential of P(3HB) polymers for tissue engineering application and found adequate biodegradability properties and good biocompatibility with several cell lines. From 2006, Misra and colleagues [89] started to investigate the use of P(3HB) extensively as the polymeric phase of composites for tissue engineering scaffolds. Bioactive glasses were added to P(3HB) in different amounts, forms and sizes (e.g. nanofibres, micro- and nanoparticulates). Looking at the protein adsorption on the different materials, the same team [81] described a significant supremacy of the composites containing 20 and 30 wt% nanoscale Bioglass® compared to the other samples (neat or micro-scale). However, following MG63 cell proliferation, they showed a rather negative effect of increasing amounts of Bioglass® incorporated (10, 20 or 30 wt%) compared to the neat material, and also for both nano- or micro-scale particles, confirming the observations of several groups mentioned in the previous paragraph. Cell growth was observed in all cases, following the proliferation profile of cells on positive control surfaces (cellculture plastic). Later, further in vitro studies [80, 81], showed that, besides increasing cell proliferation, the addition of anti-oxidant Vitamin E to the P(3HB)/ Bioglass® composite also increased the hydrophilicity (and therefore protein adsorption) on these scaffolds. The antibacterial properties of these composites have been also investigated [80]. When implanted subcutaneously in the abdominal region of Sprague Dawley rats for one week, the two types of materials (neat or containing Bioglass®) induced the formation of a thin capsule containing proliferating fibroblasts, collagen fibres, some macrophages and more interestingly, capillaries sprouts, indicating the ability of those scaffolds to enable cell colonization, and neo-vascularization, which is a key aspect in tissue engineering.

7.3.2 Bioactive glass containing composites: new developments in angiogenesis As described previously in this chapter, a variety of composite materials have been developed and show promising potential for bone tissue engineering applications. However, during the past 10 years, even if tremendous progress has been made in the development of adequate matrices (scaffolds) for bone tissue engineering, materials scientists together with cell biologists have been essentially concentrating their efforts on the osteoconductivity and osteoinductivity properties of the scaffolds. But, unlike organ transplants where there is a pre-existing vasculature, synthetic tissue engineered bone constructs are devoid of this. Back in 1963, Trueta [90], published the importance of the vasculature in osteogenesis; in a review published in 2003, Carano and Filvaroff [5] underlined the critical importance of vascularization following the implantation of a tissue-engineered scaffold for its survival, integration to the surrounding tissues and its functionality



175

(i.e. supporting and promoting neo-tissue formation). In recent years, researchers have started to address this problem by improving the rather poor angiogenic potential of tissue-engineered scaffolds by either changing their physico-chemical properties or by supplementation with angiogenic factors [91–93]. Indeed, recent in vitro studies [21, 69, 94–96] (see also ref. [23] for a recent review) using different cell types (e.g. fibroblastic cell line, intestinal epithelial cells) cultured in the presence of bioactive glasses (e.g. 45S5 Bioglass®) in different forms (e.g. coated cell culture dishes, PGLA based composites) have shown the potential angiogenic effect of Bioglass®. Specifically, a significant up-regulation of angiogenic growth factors such as VEGF (Vascular Endothelial Growth Factor) or bFGF (basic Fibroblast Growth Factor) has been found in Bioglass®-containing samples compared to Bioglass®-devoid samples at both protein secretion and gene expression levels (Table 7.3). The effect of 45S5 Bioglass® on VEGF secretion was first shown by Day et al. [21] using a rat fibroblast cell line (208F). While investigating the effect of the incorporating bioactive glass particles (0.01 to 10% w/v) into polyglycolic acid (PGA) scaffolds on cell adhesion and proliferation, the authors observed a strong inhibition of cell proliferation for percentages higher than 0.2% w/v of BG. However, VEGF secretion increased after 24, 48 and 72 hours of incubation for percentages below 0.2% w/v. In related investigations using a human myofibroblast cell line [96], or mice fibroblast [69], an up-regulation of VEGF secretion was also observed when cells were cultured in PLGA microporous spheres [96] or foams [69] containing Bioglass®, when compared to neat materials. Some other studies [94] used conditioned medium from fibroblastic cells cultured on such BG containing composites to treat endothelial cells. After 11 days of culture, a higher cell proliferation accompanied with formation of tubular structures was observed using fibroblast conditioned medium when compared to classical angiogenic medium [94]. Following the same aim, Leu and Leach [101] investigated the direct effect of bioactive glasses on neo-angiogenesis by using composite materials made of bovine collagen and 45S5 Bioglass®. Human microvascular endothelial cells (HMVEC) showed a higher cell proliferation rate compared to basic medium, which was accompanied with a higher tubular structure formation and VEGF mRNA up-regulation. However, again, a BG dose-dependent effect could be observed with 1.2 mg BG/scaffolds showing optimal responses while 0.6 mg or lower and 6 mg or higher BG concentration showed a rather negative effect. More recently, in their in vivo study using a rat calvaria implantation model, Leu et al. [100] demonstrated the angiogenic potential of BG-containing collagen scaffolds, leading to greater neo-vascularization and further bone healing when compared to the neat collagen sponges. These findings confirmed those made earlier by Day et al. [21, 69], who, using PLGA based composites, showed the presence of a significantly higher level of blood vessel invasion of materials containing bioactive glass compared to PLGA only, in both rat and mouse subcutaneous implantation models.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

176 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Table 7.3 Angiogenic indicators stimulated in response to bioactive glass Composite system Cell type

Outcomes/remarks

References

BG-coated culture Rat fibroblasts (208F) VEGF secretion wells up-regulated

[21]

Human fibroblasts (CCD-18Co)

VEGF up-regulation

[94]


VEGF, bFGF up-regulation, [94, 102] endothelial cells proliferation

Mouse fibroblasts (L929)

Disks – VEGF up-regulation, BG dose effect, cell proliferation


Porous microspheres – VEGF [96] up-regulation

Human microvascular endothelial cells (HMVEC)

Scaffolds – cell proliferation

Alginate/BG


Alginate encapsulation + BG, [95] or BG coating – VEGF ↑, less cells, dose BG effect => conditioned medium => ↑ of human endothelial-, and microvascular endothelial cells proliferation

ECM-comp

Human endothelial cell and HMVEC

BG scaffolds + collagen – VEGF up-regulation, cell proliferation

[101]

Endothelial cells/ fibroblast co-culture

BG treated aortic rings – endothelial tubule formation

[101]

PLGA/BG

[69]

[97]

Notes: BG: Bioactive Glass PLGA: poly-(D/L-lactic-co-glycolic) acid PGA: polyglycolic acid PLA: polylactic acid PLLA: poly-(L-lactic) acid PDLLA: poly-(D/L-lactic) acid PHB: polyhydroxybutyrate PCL: poly-caprolactone BMSC: Bone Marrow Stromal Cells OB: osteoblast ECM: Extra Cellular Matrix VEGF: Vascular Endothelial Growth Factor ECM comp: Extra Cellular Matrix component (e.g. type I collagen, hyaluronic acid, demineralized bone matrix) PS: phosphatidylserine P(3HB): poly(3-hydroxybutyrate).



177

7.3.3 Polymer/Bioglass® composites in musculoskeletal tissue engineering: intervertebral disc regeneration The versatility of biopolymer/Bioglass® composites in terms of structure and properties achieved enables their application in the wide musculoskeletal tissue engineering field. The case of intervertebral tissue engineering will be described here as a case in point to demonstrate the applicability of Bioglass®-containing composites in the field. In recent developments, Helen and Gough [103–105] investigated the response of annulus fibrosus cells from the intervertebral disc according to the Bioglass® particulate content in PDLLA matrix composites. Intervertebral discs (IVDs) are cartilage-like in structure and function and are situated between vertebrae, which provide a cushioning effect and allow mobility in the vertebral column. Due to the previously-reported collagenstimulating effects of bioactive glasses [106] it was hypothesized that including bioactive glasses in a composite may stimulate collagen production and benefit the repair of the collagen-rich annulus (outer structure of the IVD). Initially films were prepared containing 0, 5 and 30 wt% Bioglass® and the responses of bovine annulus fibrosus (bAFs) cells were characterized [103]. Sulphated glycosaminoglycan production was found to be higher on the scaffolds containing Bioglass® compared to PDLLA alone. 30 wt% Bioglass® inclusion gave the highest level of sGAG production after three weeks of in vitro culture, but by 4 weeks there was no significant difference between 5 and 30 wt%. The inclusion of 30 wt% Bioglass® particles also resulted in higher collagen production over the 4-week culture period. This increase in extracellular matrix production was then investigated in 3D porous scaffolds produced using TIPS [104]. These foams are highly porous with pore sizes in the range of 100 µm diameter with interconnected pores of approximately 10 to 50 µm diameter (similar to those shown in Fig. 7.2). SEM images in Fig. 7.3 show the extensive cellular colonization of bAFs cells of the scaffolds, increasing with Bioglass® content [104]. Cell numbers were increased over a 4-week culture period on foams containing 5 or 30 wt% Bioglass® as was sGAG and collagen production, with 30 wt% resulting in the highest levels of collagen production. Western blotting analysis showed that both collagen types I and II were produced, both of which are essential components of the IVD. The response of human AF (hAF) cells was also determined [105]. Similar results were obtained with the bAFs where the inclusion of Bioglass® particles increased the sGAG and collagen production. Immunostaining revealed presence of both collagen types I and II. More recently in unpublished data, hAF cells increased expression of SOX-9, collagen I and II and human mesenchymal stem cells were shown to increase expression of SOX-9, collagen I, II and aggrecan in response to inclusion of Bioglass® particles in the porous foams. These results suggest that bioactive glasses may be beneficial in regenerating the annulus of degenerate IVDs.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

178 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

7.3 SEM micrographs of bovine annulus fibrosus cells (bAFs) cultured on PDLLA/Bioglass® composite scaffolds with different Bioglass® content at different time points: neat PDLLA foams ((a): 2 weeks), PDLLA/5 wt% Bioglass® composite ((b): 2 weeks, (c): 4 weeks) showing extensive cellular colonization (reproduced from [104] with permission from Elsevier).



7.4

179

Discussion

The promising and well-explored field of tissue engineering aims to repair a diseased or a damaged tissue, by restoring both its structural and functional entities. Suitable biomaterials are required to confer to the engineered tissue adapted structural features, while cells are necessary to restore the functional activity of a given tissue. In the field of biomaterials for tissue engineering purposes, synthetic and naturally occurring polymers combined with bioactive glasses, forming composites, offer a wide range of freedom in terms of physical and chemical properties, thus the composites are highly versatile for bone and musculoskeletal tissue engineering [29, 30]. Bone and musculoskeletal tissue engineering scaffolds are generally three-dimensional, showing highly interconnected porosity in which the pore size can be adapted. In the studies cited in this chapter, which focus only on biopolymer/bioactive glass composites, two aspects have been looked at. On the one hand, the importance of the material itself, in terms of design, structure or mechanical properties and their effects on the cell/tissue response, has been highlighted. On the other hand, the positive influence of bioactive glass particulates present in the different polymer matrices on cell growth and differentiation has been considered in detail. Researchers have shown that the presence of these particulates can stimulate cell proliferation and differentiation. However, differences in the biological responses have been observed according to cell types studied, as well as to the method used for composite preparation, which leads not only to different pore morphologies but also to different bioactive glass exposition to the surrounding fluid environment. Variations in the level of exposure of the bioactive glass particulates to the biological environment, which influences the composite degradation and the ion exchange mechanism at the interface between the composite and the surrounding environment, are also of importance. Mineral deposition has been shown to slow down the polymer-based scaffold degradation and concurrently increase its mechanical properties. The presence of a mineral phase within a polymer scaffold always induces an increase of the ALP activity and improves the overall osteo-inductive and osteo-conductive properties of a composite. However a BG dose-dependent effect response has been underlined in several investigations using different cell types. Osteoblastic differentiation and fibroblastic angiogenic factor secretion appeared to occur at an optimal content between ‘too low’ and ‘too high’. The optimal content however was found lower for fibroblastic cell types compared to cells from the osteoblastic lineage. Several hypotheses to explain this effect have been suggested, including the effect of BG particle inclusion on the material wettability, or micro-/macro-surface topography, both influencing protein adsorption and subsequent cell adhesion mechanisms. Another significant parameter that has also been suggested to affect the biological behaviour of scaffolds is the pH of the culture media and cell environment, which tends to acidify owing to polymer degradation. The inclusion of bioactive glass


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

180 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

particulates within biodegradable polymer scaffolds will not only slow down the scaffold resorption, but also the alkalinity of the BG dissolution should neutralize the acidic degradation products of many polymers. In the same way, the formation of a calcium phosphate layer on the surface of BG goes through a series of ion exchange steps and precipitation/dissolution reactions, which will induce an increase of the local pH. Xynos et al. [17, 20] have investigated the influence of ion release during BG degradation in vitro, and found an important direct influence of Bioglass® dissolution products on human osteoblast cell proliferation. A summary of previous work on the gene expression effects of dissolution products of bioactive glasses has been published recently [16]. Further studies with Bioglass®-containing composites have shown encouraging results regarding the potential angiogenic effects of Bioglass® dissolution products, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularization in vivo, suggesting that scaffolds containing controlled concentrations of Bioglass® might stimulate neo-vascularization, which is beneficial to large tissue engineered constructs (Table 7.3) [21]. A recent review [23] has comprehensively discussed the current experimental evidence (in vitro and in vivo) of the angiogenic effect of bioactive glasses in the context of tissue engineering.

7.5

Conclusions and future trends

The repair of large bone defects remains a major clinical problem, and after blood, bone is the most implanted tissue. Because of the multi-aspect of its function (biological, structural and mechanical), bone also constitutes one of the most challenging domains of tissue engineering and regeneration. Musculoskeletal tissue engineering is, in general, an area of research requiring the availability of suitable scaffolds. A number of biodegradable polymers and bioactive ceramic combinations have been studied in this field of tissue engineering. In this group of materials, composites formed by combining bioactive glasses and biodegradable polymers seem to represent the materials of choice due to their adequate properties and versatility to be tailored to different applications. However, even if most of these composites can be tailored to meet the required structural/mechanical properties, and to show adequate biodegradability and cytocompatibility, the question of bioactivity of these materials is still under examination. It is important for scaffolds for bone tissue engineering to be osteoconductive and osteoinductive, in order to guide and encourage neo-bone formation. But also neo-vascularization of the tissue-engineering construct is required, representing a critical contribution to the success of regeneration and growth of new tissues as it provides cells with oxygen and nutrients. In this chapter we reviewed the relevant field, focusing on bioactive glass-containing composites for bone and musculoskeletal tissue engineering, reporting on the in vitro and in vivo performance of this family of scaffold materials. Since bioactive glasses have been developed in 1969 and have exhibited successful clinical applications as bone filler material, an increasing amount of studies have



181

been underlining the interest of these BG particulates in tissue engineering. The bioactivity of BG-based composites has been shown using a wide range of cell types (primary or cell lines) from different animal origins in vitro, but also in vivo. More recent studies have reported the angiogenic potential of such composites in both in vitro and in vivo approaches. However, the extended (clinically relevant) in vitro and in vivo understanding of these composite scaffolds is still limited, especially regarding their long-term behaviour related to the kinetics of degradation and ion release effects from bioactive glasses on the biological environment. In future studies, the specific influence of metallic ions released during degradation of the bioactive glass component of the scaffolds on bone formation and angiogenesis must be clarified and quantitatively investigated at a fundamental level. The influence of using micron-sized or nanoscale bioactive glass fillers in composites on bone and blood vessel formation will need to be investigated too. Further investigations of the scaffold bioactivity in relation to surface modification strategies, for example through the use of protein adsorption or plasma treatment, to provide more cues to cell attachment and response, will be relevant. The results of such research will enable a better understanding of the synergetic effect of bioactive glass on osteogeneis and angiogenesis, leading to control of the mineralization and neo-vascularization of the construct. This knowledge will also support the design of specific compositions of bioactive glasses to be combined with biopolymers to form improved composite constructs, which will exhibit the required architecture and microstructure (porosity content, size and orientation of the pores, amount of BG, size distribution of BG particles, type of polymer, surface topography) as well as suitable (time-dependent) mechanical properties. The scaling up of such constructs would also be a point of interest as most of the current studies are made in small-size materials. Larger scaffolds will have an influence on the nutrient diffusions of potential pre-seeded cells, but also on the specific surface area available, influencing the amount of released components. Overall, BG-based composite materials offer very promising properties that are of interest in bone tissue engineering and musculoskeletal tissue regeneration in general and future significant advances in the field are expected based on ongoing research efforts worldwide. This fact anticipates that development of composite scaffolds will remain a major area of application for bioactive glasses.

7.6

References

1. Hench L. L., Polak J. M. Third generation biomedical materials. Science 2002; 295(5557):1014–1017. 2. Hutmacher D. W., Schantz J. T., Lam C. X. F., Tan K. C., Lim T. C. State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective. J Tissue Eng Regen Med 2007; 1:245–260. 3. Zhang R. Y., Ma P. X. Poly(alpha-hydroxyl acids) hydroxyapatite porous composites for bone-tissue engineering. I. Preparation and morphology. Journal of Biomedical Materials Research 1999; 44(4):446–455.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

182 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

4. Agrawal C. M., Ray R. B. Biodegradable polymeric scaffolds for musculoskeletal tissue engineering. J Biomed Mater Res 2001; 55(2):141–150. 5. Carano R. A., and Filvaroff E. H. Angiogenesis and bone repair. Drug Discov Today 2003; 8:980. 6. Kanczler J. M., Oreffo R. O. Osteogenesis and angiogenesis: the potential for engineering bone. Eur Cell Mater 2008; 15:100–114. 7. Hench L. L. Bioceramics. J Am Ceram Soc 1998; 81:1705–1728. 8. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K. Bonding mechanisms at the interface of ceramic prosthetic materials. J Biomed Mater Res 1971; 5(6): 117–141. 9. Waltimo T., Mohn D., Paque F., Brunner T. J., Stark W. J., Imfeld T., Schaetzle M., Zehnder M. Fine-tuning of bioactive glass for root canal disinfection. J Dent Res 2009; 88(3):235–238. 10. Ducheyne P. Bioglass coatings and bioglass composites as implant materials. J Biomed Mater Res 1985; 19:273–291. 11. Kitsugi T., Nakamura T., Oka M., Senaha Y., Goto T., Shibuya T. Bone-bonding behavior of plasma-sprayed coatings of Bioglass®, AW-glass ceramic, and tricalcium phosphate on titanium alloy. J Biomed Mater Res 1996; 30(2):261–269. 12. Chen Q. Z., Thompson I. D., Boccaccini A. R. 45S5 Bioglass®-derived glass–ceramic scaffolds for bone tissue engineering. Biomaterials 2006; 27(11): 2414–2425. 13. Lai W., Garino J., Ducheyne P. Silicon excretion from bioactive glass implanted in rabbit bone. Biomaterials 2002, 23:213–217. 14. Zhang K., Wang Y., Hillmayer M. A., Francis L. F. Processing and properties of porous poly(L-lactide)/bioactive glass composites. Biomaterials 2004; 25:2489–2500. 15. Lu H., Zhang T., Wang X. P., Fang Q. F. Electrospun submicron bioactive glass fibers for bone tissue scaffold. J Mater Sci: Mater Med 2009; 20:793–798. 16. Hench L. L., Thompson I. D. Twenty-first century challenges for biomaterials. J Royal Soc Interface 2010, in press. 17. Xynos I. D., Hukkanen M. V. J., Batten J. J., Buttery L. D., Hench L. L., Polak J. M. Bioglass® 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: Implications and applications for bone tissue engineering. Calcified Tissue Int 2000; 67(4):321–329. 18. Bielby R. C., Christodoulou I. S., Pryce R. S., Radford W. J. P., Hench L. L. Polak J. M. Time- and concentration-dependent effects of dissolution products of 58S sol-gel bioactive glass on proliferation and differentiation of murine and human osteoblasts. Tissue Eng 2004; 10:1018–1026. 19. Bielby R. C., Pryce R. S., Hench L. L., Polak, J. M. Enhanced derivation of osteogenic cells from murine embryonic stem cells after treatment with ionic dissolution products of 58s bioactive sol-gel glass. Tissue Eng 2005; 11:479–488. 20. Xynos I. D., Edgar A. J., Buttery L. D. K., Hench L. L., Polak M. Gene expression profiling of human osteoblasts following treatment with the ionic products of Bioglass® 45S5 dissolution. J Biomed Mater Res, 2001; 55:151–157. 21. Day R. M., Boccaccini A. R., Shurey S., Roether J. A., Forbes A., Hench L. L., Gabe S. M. Assessment of polyglycolic acid mesh and bioactive glass for soft-tissue engineering scaffolds. Biomaterials 2004; 25:5857–5866. 22. Day R., Boccaccini A. R., Roether J. A., Surey S., Forbes A., Hench L. L., Gabe S. The effect of Bioglass® on epithelial cell and fibroblast proliferation and incorporation into a PGA matrix. Gastroenterology 2002; 122(4): Suppl. 1: T875.



183

23. Gorustovich A. A., Roether J. A., Boccaccini A. R., Effect of bioactive glasses on angiogenesis: In vitro and in vivo evidence. A Review. Tissue Eng Part B 2010; 16:199–207. 24. Vallet-Regi M., Salinas A. J., Roman J., Gil M. Effect of magnesium content on the in vitro bioactivity of CaO-MgO-SiO2-P2O5 sol-gel glasses. J Mater Chem 1999; 9(2):515–518. 25. Lao J., Nedelec J. M., Jallot E. New strontium-based bioactive glasses: physicochemical reactivity and delivering capability of biologically active dissolution products. J Mater Chem 2009; 19:2940–2949. 26. Varanasi V. G., Saiz E., Loomer P. M., Ancheta B., Uritani N., Ho S. P., Tomsia A. P., Marshall S. J., Marshall G. W. Enhanced osteocalcin expression by osteoblast-like cells (MC3T3-E1) exposed to bioactive coating glass (SiO2-CaO-P2O5-MgO-K2ONa2O system) ions. Acta Biomater 2009; 5:3536–3547. 27. Oki A., Parveen B., Hossain S., Adeniji S., Donahue H. Preparation and in vitro bioactivity of zinc containing sol-gel-derived bioglass materials. J Biomed Mater Res A 2004; 69A(2):216–221. 28. Liu X., Huang W., Fu H., Yao A., Wang D., Pan H., Lu W. W., Jiang X., Zhang X. Bioactive borosilicate glass scaffolds: in vitro degradation and bioactivity behaviours: J Mater Sci: Mater Med 2009; 20:1237–1243. 29. Rezwan K., Chen Q. Z., Blaker J. J., Boccaccini A. R. Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 2006; 27(18):3413–3431. 30. Guarino V., Causa F., Ambrosio L. Bioactive scaffolds for bone and ligament tissue. Expert Rev Medical Devices 2007; 4(3):405–418. 31. Boccaccini A. R., Maquet V. Bioresorbable and bioactive Polymer/Bioglass® composites with tailored pore structure for tissue engineering applications. Comp Sci Technol 2003; 63:2417–2429. 32. Lu H. H., El Amin S. F., Scott K. D., Laurencin C. T. Three-dimensional, bioactive, biodegradable, polymer-bioactive glass composite scaffolds with improved mechanical properties support collagen synthesis and mineralization of human osteoblast-like cells in vitro. J Biomed Mater Res A 2003; 64:465–474. 33. Roether J. A., Boccaccini A. R., Hench L. L., Maquet V., Gautier S., Jerome R. Development and in vitro characterisation of novel bioresorbable and bioactive composite materials based on polylactide foams and Bioglass for tissue engineering applications. Biomaterials 2002; 23:3871–3878. 34. Hutmacher D. W. Scaffolds in tissue engineering bone and cartilage’, Biomaterials 2001; 21:2529–2543. 35. Mano J. F., Sousa R. A., Boesel L. F., Neves N. M., Reis R. L. Bioinert, biodegradable and injectable polymeric matrix composites for hard tissue replacement: state of the art and recent developments. Comp Sci Tech 2004; 64(6):789–817. 36. Jagur-Grodzinski J., Polymers for tissue engineering, medical devices and regenerative medicine. Concise general review of recent studies. Polym Adv Technol 2007; 17: 305–418. 37. Maquet V., Martin D., Scholtes F. et al. Poly(D,L-lactide) foams modified by poly(ethylene oxide)-block-poly(D,L-lactide) copolymers and a a-FGF: in vitro and in vivo evaluation for spinal cord regeneration. Biomaterials 2001; 22:1137–1146. 38. Ruhe P. Q., Hedberg-Dirk E. L., Padron N. T., Spauwen P. H., Jansen J. A., Mikos A. G. Porous poly(DL-lactic-co-glycolic acid)/calcium phosphate cement composite for reconstruction of bone defects. Tissue Eng 2006; 12:789–800.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

184 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

39. Misra S. K., Vallappil S. P., Roy I., Boccaccini A. R. Polyhydroxalkanoate/inorganic phase composites for tissue engineering applications’, Biomacromolecules 2006; 7:2249–2253. 40. Cool S. M., Kenny B., Wu A., Nurcombe V., Trau M., Cassady A. I., Grondahl L. Poly(3-hydroxybutyrate-co-3-hydroxyvalerate) composite biomaterials for bone tissue regeneration: in vitro performance assessed by osteoblast proliferation, osteoclast adhesion and resorption, and macrophage proinflammatory response. J Biomed Mater Res A 2007; 82(3):599–610. 41. Jo J. H., Lee E. J., Shin D. S., Kim H. E., Kim H. W., Koh Y. H., Jang J. H. in vitro/in vivo biocompatibility and mechanical properties of bioactive glass nanofiber and poly(e-caprolactone) composite materials. J Biomed Mater Res B: Appl Biomaterials 2009; 91B(1):213–220. 42. Couto D. S., Hong Z., Mano J. F. Development of bioactive and biodegradable chitosan-based injectable systems containing bioactive glass nanoparticles. Acta Biomater 2009; 5(1):115–123. 43. Marelli B., Ghezzi C. E., Barralet J. E., Boccaccini A. R., Nazhat S. N., Threedimensional mineralization of dense nanofibrillar collagen-bioglass hybrid scaffolds. Biomacromolecules 2010, 11:1470–1479. 44. Ma P. X., Choi J. W. Biodegradable polymer scaffolds with well-defined interconnected spherical pore network. Tissue Eng 2001; 7(1):23–33. 45. Matthews F. L., Rawlings R. D., Composite Materials: Engineering and Science. London: Chapman and Hall 1994. 46. Misra S. K., Boccaccini A. R. Biodegradable and bioactive polymer/ceramic composite scaffolds. Chapter 4 in Boccaccini, A. R., Gough, J. E., eds., Tissue Engineering Using Ceramics and Polymers. Cambridge, UK: Woodhead Publishing/ CRC 2007. 47. Blaker J. J., Gough J. E., Maquet V., Notingher I., Boccaccini A. R. In vitro evaluation of novel bioactive composites based on Bioglass®-filled polylactide foams for bone tissue engineering scaffolds. J Biomed Mater Res 2003; 67A:1401–1411. 48. Brunner T. J., Grass R. N., Stark W. J. Glass and bioglass nanopowders by flame synthesis. Chem Commun 2006; 13:1384–1386. 49. Boccaccini A. R., Erol M., Stark W. J., Mohn D., Hong Z., Mano J. F. Polymer/ bioactive glass nanocomposites for biomedical applications. A review. Comp Sci Technol 2010; in press. 50. Kim H. W., Kim H. E., Knowles J. C. Production and potential of bioactive glass nanofibers as a next-generation biomaterial. Adv Func Mater 2006; 16(12):1529–1535. 51. Yang S., Leong K., Du Z., Chua C. The design of scaffolds for use in tissue engineering. Part I Traditional factors. Tissue Eng 2001; 7:679–689. 52. Liu X., Ma P. X. Polymeric scaffolds for bone tissue engineering. Annals of Biomedical Engineering 2004; 32:477–486. 53. Guan L. M., Davies J. E. Preparation and characterization of a highly macroporous biodegradable composite tissue engineering scaffold. Journal of Biomedical Materials Research Part A 2004; 71A(3):480–487. 54. Laurencin C. T., Lu H. H., Khan Y. Processing of polymer scaffolds: polymer-ceramic composite foams. In Atala A., Lanza R. P., eds, Methods of Tissue Engineering. California: Academic Press 2002, pp. 705–714. 55. Schugens C., Maquet V., Grandfils C., Jerome R., Teyssie P. Biodegradable and macroporous polylactide implants for cell transplantation: Polylactide macroporous biodegradable implants for cell transplantation. II. Preparation of polylactide foams by



56.

57.

58. 59. 60.

61.

62.

63.

64.

65.

66.

67.

68.

69.

70.

71.

185

liquid-liquid phase separation. Journal of Biomedical Materials Research 1996; 30:449–461. Zhang R. Y., Ma P. X. Poly(alpha-hydroxyl acids) hydroxyapatite porous composites for bone-tissue engineering. I. Preparation and morphology. Journal of Biomedical Materials Research 1999; 44(4):446–455. Taboas J. M., Maddox R. D., Krebsbach P. H., Hollister S. J. Indirect solid free form fabrication of local and global porous, biomimetic and composite 3D polymer-ceramic scaffolds. Biomaterials 2003; 24:181–194. Yang S., Leong K., Du Z., Chua C. The design of scaffolds for use in tissue engineering. Part II Rapid prototyping techniques. Tissue Eng 2002; 8:1–11. Ma P. X., Choi J. W. Biodegradable polymer scaffolds with well-defined interconnected spherical pore network. Tissue Eng 2001; 7(1):23–33. Blaker J. J., Maquet V., Jerome R., Boccaccini A. R., Nazhat S. N. Mechanical properties of highly porous PDLLA/Bioglass® composite foams as scaffolds for bone tissue engineering. Acta Biomaterialia 2005; 1(6):643–652. Hutmacher D. W., Cool S. Concepts of scaffold-based tissue engineering – the rationale to use solid free-form fabrication techniques. J Cellular Molec Med 2007; 11: 654–669. Xiong Z., Yan Y. N., Wang S. G., Zhang R. J., Zhang C. Fabrication of porous scaffolds for bone tissue engineering via low-temperature deposition. Scripta Mater 2002; 46:771–776. Misra S. K., Mohn D., Brunner T. J., Stark W. J., Philip S. E., Roy I., Salih V., Knowles J. C., Boccaccini A. R. Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB)/Bioglass® composites. Biomaterials 2008; 29(12):1750–1761. Hong Z., Reis R. L., Mano J. F. Preparation and in vitro characterization of scaffolds of poly(l-lactic acid) containing bioactive glass ceramic nanoparticles. Acta Biomaterialia 2008; 4:1297–1306. El-Kady A. M., Ali A. F., Farag M. M. Development, characterization, and in vitro bioactivity studies of sol-gel bioactive glass/poly(L-lactide) nanocomposite scaffolds. Mater Sci Eng C 2010; 30(1):120–131. Verrier S., Boccaccini A. R. Bioactive composite materials for bone tissue engineering scaffolds. Chapter 14 in J. Polak, S. Mantalaris, S. E. Harding, eds., Advances in Tissue Engineering. London, UK: Imperial College Press 2008, pp. 279–311. Yao J., Radin S., Reilly G., Leboy P. S., Ducheyne P. Solution-mediated effect of bioactive glass in poly(lactic-co-glycolic acid)-bioactive glass composites on osteogenesis of marrow stromal cells. J Biomed Mater Res A 2005; 75:794. Yao J., Radin S., Leboy S., Ducheyne P. The effect of bioactive glass content on synthesis and bioactivity of composite poly(lactic-co-glycolic acid)/bioactive glass substrate for tissue engineering. Biomaterials 2005; 26:1935. Day R. M., Maquet V., Boccaccini A. R., Jerome R., Forbes A. In vitro and in vivo analysis of macroporous biodegradable poly(D,L-lactide-co-glycolide) scaffolds containing bioactive glass. J Biomed Mater Res A 2005; 75:778. Keshaw H., Georgiou G., Blaker J. J., Forbes A., Knowles J. C., Day R. M. Assessment of polymer/bioactive glass-composite microporous spheres for tissue regeneration applications. Tissue Eng Part A 2009; 15:1451. Lu H. H., El Amin S. F., Scott K. D., Laurencin C. T. Three-dimensional, bioactive, biodegradable, polymer-bioactive glass composite scaffolds with improved mechanical properties support collagen synthesis and mineralization of human osteoblast-like cells in vitro. J Biomed Mater Res A 2003 64:465.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

186 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

72. Lu H. H., Tang A., Oh S. C., Spalazzi J. P., Dionisio K. Compositional effects on the formation of a calcium phosphate layer and the response of osteoblast-like cells on polymer-bioactive glass composites. Biomaterials 2005; 26:6323. 73. Roether J. A., Gough J. E., Boccaccini A. R., Hench L. L., Maquet V., Jerome R. Novel bioresorbable and bioactive composites based on bioactive glass and polylactide foams for bone tissue engineering. J Mater Sci: Mater Med 2002; 13:1207. 74. Ruuttila P., Niiranen H., Kellomaki M., Tormala P., Konttinen Y. T., and Hukkanen M. Characterization of human primary osteoblast response on bioactive glass (BaG 13-93)-coated poly-L,DL-lactide (SR-PLA70) surface in vitro. J Biomed Mater Res B: Appl Biomater 2006; 78:97–106. 75. Yang X. B., Webb D., Blaker J., Boccaccini A. R., Maquet V., Cooper C., Oreffo R. O. Evaluation of human bone marrow stromal cell growth on biodegradable polymer/ bioglass composites. Biochem Biophys Res Commun 2006; 342:1098–1104. 76. Tsigkou O., Hench L. L., Boccaccini A. R., Polak J. M., Stevens M. M. Enhanced differentiation and mineralization of human fetal osteoblasts on PDLLA containing Bioglass composite films in the absence of osteogenic supplements. J Biomed Mater Res A 2007; 80:837–842. 77. Verrier S., Blaker J. J., Maquet V., Hench L. L., Boccaccini A. R. PDLLA/Bioglass composites for soft-tissue and hard-tissue engineering: an in vitro cell biology assessment. Biomaterials 2004; 25:3013–3020. 78. Brauer D. S., Russel C., Vogt S., Weisser J., Schnabelrauch M. Degradable phosphate glass fiber reinforced polymer matrices: mechanical properties and cell response. J Mater Sci: Mater Med 2007. 79. Erdemli O., Captug O., Bilgili H., Orhan D., Tezcaner A., Keskin D. In vitro and in vivo evaluation of the effects of demineralized bone matrix or calcium sulfate addition to polycaprolactone-bioglass composites. J Mater Sci: Mater Med 2010; 21:295–298. 80. Misra S. K., Ansari T. I., Valappil S. P., Mohn D., Philip S. E., Stark W. J., Roy I., Knowles J. C., Salih V., Boccaccini A. R. Poly(3-hydroxybutyrate) multifunctional composite scaffolds for tissue engineering applications. Biomaterials 2010; 31: 2806–2812. 81. Misra S. K., Philip S. E., Chrzanowski W., Nazhat S. N., Roy I., Knowles J. C., Salih V., Boccaccini A. R. Incorporation of vitamin E in poly(3hydroxybutyrate)/Bioglass composite films: effect on surface properties and cell attachment. J R Soc Interface 2009; 6:401–408. 82. Bretcanu O., Misra S., Roy I., Renghini C., Fiori F., Boccaccini A. R., Salih V. In vitro biocompatibility of 45S5 Bioglass-derived glass-ceramic scaffolds coated with poly(3hydroxybutyrate). J Tissue Eng Regen Med 2009; 3:139–147. 83. Misra S. K., Ansari T., Mohn D., Valappil S. P., Brunner T. J., Stark W. J., Roy I., Knowles J. C., Sibbons P. D., Jones E. V., Boccaccini A. R., Salih V. Effect of nanoparticulate bioactive glass particles on bioactivity and cytocompatibility of poly(3-hydroxybutyrate) composites. J R Soc Interface 2010; 7:453–459. 84. Xu C., Wang Y., Yu X., Chen X., Li X., Yang X., Li S., Zhang X., Xiang A. P. Evaluation of human mesenchymal stem cells response to biomimetic bioglass-collagenhyaluronic acid-phosphatidylserine composite scaffolds for bone tissue engineering. J Biomed Mater Res A 2009; 88:264–272. 85. Day R. M., Maquet V., Boccaccini A. R., Jerome R., Forbes A. In vitro and in vivo analysis of macroporous biodegradable poly(D,L-lactide-co-glycolide) scaffolds containing bioactive glass. J Biomed Mater Res A 2005; 75:778.



187

86. Boccaccini A. R., Notingher I., Maquet V., Jerome R. Bioresorbable and bioactive composite materials based on polylactide foams filled with and coated by Bioglass particles for tissue engineering applications. J Mater Sci: Mater Med 2003; 14:443. 87. Boccaccini A. R., Blaker J. J. Bioactive composite materials for tissue engineering scaffolds. Expert Rev Med Devices 2005: 2:303. 88. Chen G. Q., Wu Q. The application of polyhydroxyalkanoates as tissue engineering materials. Biomaterials 2005; 26:6565–6571. 89. Misra S. K., Nazhat S. N., Valappil S. P., Moshrefi-Torbati M., Wood R. J., Roy I., Boccaccini A. R.T Fabrication and characterization of biodegradable poly(3hydroxybutyrate) composite containing bioglass. Biomacromolecules 2007; 8: 2112–2119. 90. Trueta J. The role of vessels in osteogenesis. Journal of Bone and Joint Surgery Series A 1963; 45B:402–408. 91. Kanczler J. M., Ginty P. J., Barry J. J., Clarke N. M., Howdle S. M., Shakesheff K. M., Oreffo R. O. The effect of mesenchymal populations and vascular endothelial growth factor delivered from biodegradable polymer scaffolds on bone formation. Biomaterials 2008; 29:1892. 92. Kanczler J. M., Ginty P. J., White L., Clarke N. M., Howdle S. M., Shakesheff K. M., Oreffo R. O. The effect of the delivery of vascular endothelial growth factor and bone morphogenic protein-2 to osteoprogenitor cell populations on bone formation. Biomaterials 2010; 31:1242. 93. Rocha F. G., Sundback C. A., Krebs N. J., Leach J. K., Mooney D. J., Ashley S. W., Vacanti J. P., and Whang E. E. The effect of sustained delivery of vascular endothelial growth factor on angiogenesis in tissue-engineered intestine. Biomaterials 2008; 29:2884. 94. Day R. M. Bioactive glass stimulates the secretion of angiogenic growth factors and angiogenesis in vitro. Tissue Eng 2005; 11:768. 95. Keshaw H., Forbes A., Day R. M. Release of angiogenic growth factors from cells encapsulated in alginate beads with bioactive glass. Biomaterials 2005; 26:4171. 96. Keshaw H., Georgiou G., Blaker J. J., Forbes A., Knowles J. C., Day R. M. Assessment of polymer/bioactive glass-composite microporous spheres for tissue regeneration applications. Tissue Eng Part A 2009; 15:1451. 97. Leach J. K., Kaigler D., Wang Z., Krebsbach P. H., Mooney D. J. Coating of VEGFreleasing scaffolds with bioactive glass for angiogenesis and bone regeneration. Biomaterials 2006; 27:3249. 98. Pyhalto T., Lapinsuo M., Patiala H., Niiranen H., Tormala P., Rokkanen P. Fixation of distal femoral osteotomies with self-reinforced poly(L/DL)lactide 70:30 and selfreinforced poly(L/DL)lactide 70:30/bioactive glass composite rods. An experimental study on rabbits. J Biomater Sci Polym Ed 2005; 16:725. 99. Ignatius A. A., Betz O., Augat P., Claes L. E. In vivo investigations on composites made of resorbable ceramics and poly(lactide) used as bone graft substitutes. J Biomed Mater Res 2001; 58:701. 100. Leu A., Stieger S. M., Dayton P., Ferrara K. W., Leach J. K. Angiogenic response to bioactive glass promotes bone healing in an irradiated calvarial defect. Tissue Eng Part A 2009; 15:877–887. 101. Leu A., Leach J. K. Proangiogenic potential of a collagen/bioactive glass substrate. Pharm Res 2008; 25:1222.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

188 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

102. Moosvi S. R., Day R. M. Bioactive glass modulation of intestinal epithelial cell restitution. Acta Biomater 2009; 5:76–82. 103. Helen W., Gough, J. E. In vitro studies of annulus fibrosus disc cell attachment, differentiation and matrix production on PDLLA/45S5 Bioglass composite films. Biomaterials 2006; 27(30):5220–5229. 104. Helen W., Merry C. L., Blaker J. J., Gough J. E. Three-dimensional culture of annulus fibrosus cells within PDLLA/Bioglass composite foam scaffolds: assessment of cell attachment, proliferation and extracellular matrix production. Biomaterials 2007; 28:2010–2020. 105. Helen W., Gough J. E. Cell viability, proliferation and extracellular matrix production of human annulus fibrosus cells cultured within PDLLA/Bioglass composite foam scaffolds in vitro. Acta Biomater 2008; 4:230–243. 106. Bosetti M., Zandardi L., Hench L. L., Cannas M. Type I collagen production by osteoblast like cells cultured in contact with different bioactive glasses. J Biomed Mater Res A 2003; 64:189–195.

7.7

Appendix: list of abbreviations

BG: Bioactive glass BMSC: Bone Marrow Stromal Cells ECM comp: Extra Cellular Matrix component (e.g. Type I collagen, hyaluronic acid, demineralized bone matrix) ECM: Extra Cellular Matrix OB: osteoblast P(3HB): poly(3-hydroxybutyrate) PCL: poly-caprolactone PDLLA: poly-(D/L-lactic) acid PGA: polyglycolic acid PHB: polyhydroxybutyrate PLA: polylactic acid PLGA: poly-(D/L-lactic-co-glycolic) acid PLLA: poly-(L-lactic) acid PS: phosphatidylserine VEGF: Vascular Endothelial Growth Factor


8 Use of bioactive glasses as bone substitutes in orthopaedics and traumatology J. HEIKKILÄ , Sports Clinic and Hospital Mehiläinen Turku, Finland

Abstract: Bioactive glass is one of many biomaterials suggested for bone replacement. Its advantages include non-toxicity, biocompatibility and bioactivity. Additionally, through a surface reaction the glass forms a Ca’P-layer after implantation within the body, through which the host bone grows together with the glass. Through continuous reaction and layer formation the glass is finally absorbed and replaced with bone. Fractures and cavitary bone tumors at metaphyseal bone regions are among the best clinical applications for bioactive glass. It has also been used in diaphyseal bone defects and fractures, and in spinal surgery with good clinical results. The possible clinical applications are reported in this chapter. There have been some good prospective randomized studies into using bioactive glass in metaphyseal bone areas; however, there is a need for more research into using bioactive glass within orthopaedics and traumatology. Key words: bioactive glass, bone substitutes, bone defect, fractures, metaphyseal area, clinical, orthopaedics and traumatology.

8.1

Introduction

The search for suitable bone substitute materials has been going on for more than 120 years [1]. The aim has been to find a material that can be used safely and effectively in place of autogenous bone, since the harvesting of autogenous bone from the iliac crest is always an invasive accessory operation causing pain, bleeding and morbidity for the patient. Furthermore, postoperative rehabilitation is usually more difficult. The operation time is lengthened and a second surgeon is usually needed for the harvesting procedure. In modern orthopaedics and traumatology there is a trend towards smaller incisions, arthroscopic procedures, and shorter operation times, which are all facets of minimal invasive surgery, i.e. surgery that disturbs the life of the patient as little as possible. In this respect, the use of bone substitutes is a necessity. Hospital directors want to ensure that operating theaters are used as effectively as possible: the use of bone substitutes offers interesting possibilities in this regard.

8.1.1 History of bone substitutes The first attempts to identify possible substitutes for autogenous bone were made by a Dutch scientist, Job van Meekeren, who tried to use fresh canine skull bone as a bone grafting material as early as 1668 [2]. It was not until 1880 that Maceven 189 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

190 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

first used fresh bone from a human cadaver to replace the damaged part of a bone [3]. Following on from this, there developed a widespread clinical use of allograft bone, and fresh frozen allograft bone is still in clinical use today [4]. In a way it can be regarded as the gold standard bone substitute material. The first artificial material used for this purpose was calcium sulphate, CaSO4: Dreesman was already using this to fill cavities in 1883 [5]. This material is surprisingly still on the market. In 1889 Senn used antiseptic decalcified bone [1], an idea that was further developed 75 years later by Urist [6], and in 1920 Albee introduced artificial tricalcium phosphate [7]. Deproteinized bovine bone was first used by Orell in 1934 [8]; he named it os purum, clean bone. In 1952 Bauermeister introduced a similar concept, Kieler bone, which was also made from deproteinized bovine bone. Demineralized bone was further studied by Urist in 1965 [9], and through these studies he discovered bone morphogeneic proteins in 1971 [10]: these proteins represent one future direction in the development of bone substitute materials. The idea that certain types of glass might be bioactive and could chemically bond with bone was first introduced by Hench in 1967; he then proved this to be true in 1971 [11]. Hench glass is still in clinical use today. Coralline hydroxylapatite was introduced in 1974 by Roy and Linnehan [12], and in 1977 Jarcho proved that dense hydroxylapatite could bond with bone [13]. At the same time the first attempt to use composite materials for the purpose was made by Mittelmeier and his colleagues, who combined demineralized bone matrix with collagen [1]. Gross and Strunz studied bioactive glass-ceramic in Germany [15], while Kokubowas was conducting similar investigations in Japan at the same time [16]. Osborn and Furlong further developed hydroxylapatite, and in 1991 they manufactured synthetic hydroxylapatite [17], which is used today for coating metal prosthesis and external fixation pins. Shortly afterwards, in 1992, Van Blitterswijk developed bioactive polymer in Leiden [18]. Meanwhile, in California, Constanz worked on the formation of coralline calcium carbonate in corals, and realized that it was possible to mimic coral formation in the laboratory. As a result, an injectable form of calcium phosphate was developed, which sets into fully hard dahllite 12 hours after injection into the body [19]. At Åbo Akademi University in Turku, Finland, Professor Kaj Karlsson developed a series of different glasses in the late 1980s, of which S53P4 bioactive glass seemed to have the most potential. Andersson published his thesis concerning these materials in 1991 [20]; it was the first thesis on bioactive glass to emerge from Turku, and since then several others have been published. The university project started by the Turku Biomaterials Group has now finally been commercialized, and the material is marketed by BonAlive Biomaterials Ltd.

8.1.2 Ideal bone substitute materials The ideal bone substitute material would be biomechanically stable, non-toxic and biocompatible. It would be replaced by healing bone, and would contain


Bioactive glasses as bone substitutes in orthopaedics

191

inductive factors to promote bone healing. Additionally, it must not induce any risks for the patient, should be easy to use, inexpensive, and readily available. With these prerequisites it is obvious that there is not a single material that would fulfill all these requirements. Implanting materials in the tissues can be regarded as forming a chronic wound in the tissue, and the material characteristics of the implant and implantation site regulate the histological and tissue response towards implantation. The response is also affected by biomechanical factors and the possible movement of the implants, which may affect the long-term results of the implanted material [21]. Bone defects can be divided into non-segmental, segmental, and articular joint defects [22]: each of these is best reconstructed with a specific type of bone substitute. There is no one material that is suitable for use in all of these defects.

8.1.3 History of bioactive glass in orthopaedics The concept of bioactive glass and its use in bone surgery was introduced in 1967 by Hench, who hypothesized that glass containing calcium and phosphorus would be biocompatible: he used the glass phase diagram as the basis for the choice of composition. In 1969 the US Army Medical Research and Development Command approved an experimental study plan with two aims. The first aim was to achieve a direct chemical bond between biomaterial and bone. The second was to achieve an understanding of the reactions occurring between bone and implanted materials. The first animal experiments were performed in 1969, with the first results published in 1971. The original hypothesis was proven to be correct [11], and in addition the first bone bonding material was designed. The development of these new materials started in various centers immediately after the first reports by the Florida group. A German group led by Gross and Strunz developed a bone-bonding glass-ceramic (GC), Ceravital®, and reported the first results in 1980 [15]. Ceravital® consists of a glass phase into which apatite crystals are embedded. It has a significantly higher bending strength compared with glasses [23]. An even stronger GC material, Cerabone®, was manufactured at Kyoto University in Japan by Professor Kokubo and his colleagues in 1982 [16, 24]. It has also been called AW G-C (apatite wollastonite glass-ceramic). It is crystallographically composed of apatite and wollastonite crystals embedded in a glassy matrix. Since its development, this material has been used clinically even in load-bearing situations [25]. Ilmaplant®-L1 [26] and Bioverit® are glass-ceramics developed in the former East Germany. The first has a composition resembling that of Ceravital® [15], but crystallographically it consists of apatite and wollastonite embedded in glass. The structure is similar to that of Cerabone® [23]. The second, Bioverit®, was developed at Friedrich Schiller University in Jena. It differs from the materials mentioned above, and includes remarkable amounts of fluoride and aluminum oxide. Crystallographically three phases can be detected: apatite, phlogopite and glass. Bioverit® has been claimed to be bioactive.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

192 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

At Åbo Akademi, Professor Kaj Karlsson developed his series of glasses, and selected those that were most interesting in terms of clinical applications. In his group, Dr. Andersson developed a mathematical analysis to evaluate glasses and he selected S53P4 as the glass that was both most suitable and most interesting for clinical use, and published this in his thesis in 1991 [27]. Since then, at least 40 theses on bioactive glass have been published by the Turku Biomaterials Group [20, 28–30].

8.2

Glass surface reactions

According to Hench [11], bone bonding can be presented as a complex series of reactions in the glass and on the glass surface (Fig. 8.1). The chemical reactions on the glass surface are based on leaching, dissolution, and precipitation as reviewed by Hench and Andersson [31] and Ducheyne et al. [32]. After implantation in simulated or in vivo body fluids the first rapid reaction on the glass surface is the exchange of Na+ or K+ with H+ or H3O+ from solution. The leaching is facilitated by the fact that alkaline and earth alkaline cations are not part of the network, but are only modifiers. This reaction produces an alkaline microenvironment in which the alkalinity of the solution results from the breaking of -Si-O-Si- bonds, mainly by hydroxyl ions. This dissolution occurs only locally at the glass surface and results in the formation of silanol (SiOH) groups at the glass–solution interface. The hydrated silica groups condense and repolymerize

8.1 Surface reactions and Si-rich and Ca,P-layer formation at the surface of bioactive glasses. First a layer of glass with the thickness of a couple of microns is dissolved (I). Secondly, Na, Ca, P and also Si are leached from the surface of the glass (II) and a Si-rich layer is formed through repolymerization (III). Ca and P from the solution and partly leached from the glass precipitate in the surface of the Si-rich layer (IV).



193

with the silanol groups forming a SiO2- rich layer on the surface. This precipitation is facilitated by the migration of Ca2+ and PO43− groups to the surface through SiO2-rich layer and there is further incorporation of Ca2+ and PO43− from the solution into this layer. This step occurs within minutes of the implantation of the reactive bioactive glasses (Fig. 8.1) [20]. The SiO2-rich layer increases in thickness due to diffusion-controlled alkali ion exchange and the final thickness of this layer is dependent on its composition. The amorphous CaO-P2O5-rich film grows by incorporation of soluble calcium phosphate from the supersaturated solution. The nucleation of a CaO-P2O5-rich film on top of the SiO2-rich layer and formation of bone-like carbonated calciumdeficient apatite on the glass surface has been observed by Fourier Transform Infrared spectroscopy (FTIR) within 10 hours of the implantation. Crystallization of the amorphous CaO-P2O5 film also occurs by incorporation of OH−, CO32− and F− anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer, probably through similar reactions as those that cause the calcium-deficient apatite dahllite to crystallize on the bone [31, 32]. The reactions described above result in a 100 to 120 µm thick layer, which is rich in SiO2 and a layer about 30 µm thick of hydroxycarbonate apatite (HCA) [33].

8.2.1 Tissue reactions on the surface of bioactive glass The biological processes involved in bone formation and bonding on the glass surface can be divided into six stages (Table 8.1). The SiO2-HCA layer is formed in a similar manner both in vitro and in vivo. However, in vitro a Ca,P layer is formed on the glass surface, whereas in vivo it is formed within the surface of the Si-rich layer. In vivo, biological compounds such as proteins and collagen adhere to the glass surface. This phase can occur even in the absence of cells or growth factors. Apatite is formed by crystallization around collagen which has been observed to be trapped in the growing HCA layer [34]. Bioactive glasses and in particular the formed HCA layer favor the chemotaxis of osteoblasts on the glass surface [21]. These cells attach and differentiate and commence their secretion, and more matrix is formed. Simultaneously and also causatively, collagen fibers and Table 8.1 The different stages of bond formation between glass surface and bone Stage 1.

Biological structures adsorbed in the SiO2-HCA layer

Stage 2.

Precipitation of Ca and P from the glass and solution between the biological structures, especially collagen

Stage 3.

Attachment of stem cells (and preosteoblasts) to the surface layer

Stage 4.

Differentiation of cells on the glass surface

Stage 5.

Matrix formation

Stage 6.

Crystallization of matrix


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

194 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

mucopolysaccharides agglomerate around and within the apatite crystals and bond with them (Table 8.1). The reactions in vivo produce an interface with a chemical connection between glass surface and bone. The phenomenon is called bonding osteogenesis: bone bonding. The chemical bonding is further strengthened by the adhesion of proteins and collagen to the Ca,P-layer. Hench and Paschall [35] published a paper on the adhesion of poly-L-alanine to a bioactive glass surface. Organic material and collagen have been observed in vitro by Wilson and Nolletti to be trapped within the growing HCA layer [34]. The same result has been shown to occur in vivo by various authors [36–38].

8.3

The bonding of bioactive glass and bone formation

In summary, bone repair after implantation of bioactive glass and the reactions that occur simultaneously on the surface of the glass can be explained in the following way. Bone minerals have a rapid turnover, and crystal formation occurs epitaxially on bone. The biological part of the matrix, consisting of collagens, non-collagenous proteins, and glycosaminoglycans, modulates crystal formation. Similarly, crystallization of Ca,P on the bioactive glass surface also occurs epitaxially; it is also likely that the same biological structures modulate the precipitation and are trapped between the crystals. The carbonated HA on the glass surface closely resembles the calcium deficient carbonated HA, dahlite found in bone. Taking these chemical and biological reactions into account, it is easy to understand the formation of the bond between bioactive glass and bone. A similar type of crystal formation at the surface of bioactive glass enables bone to bond biochemically with bioactive glass. Nevertheless, it should be noted that the bone formation cannot simply be a case of coalescence of two epitaxially-growing crystal fronts: the biological structures, collagen, proteins, and glycosaminoglycans affect bond formation, probably through similar processes to those that occur during bone crystal formation.

8.3.1 The bonding strength of bioactive glass The bonding between bone and bioactive glass has been verified with push-out tests, which have shown that the strength of the bond between bioactive glasses and bone ranges from 16 to 23 N/mm2. Using the same test method, the push-out strength when there is no contact between bone and implant (e.g. with inactive glass or metals) was 0.5 N/mm2, and for materials with physical contact from 2 to 4 N/mm2 [39] (Fig. 8.2). It is postulated that when the bond between the bone and bioactive glass forms, the silica-rich layer is the weakest part. However this is not supported by push-out tests, in which the fracture line occurs at random in the



195

8.2 Schematic presentation of bone response towards three types of implant materials: (a) a fibrous capsula encapsulates the biotolerant implant material; (b) intimate contact without bone bonding occurs at the interface between bioinert materials and bone; (c) intimate contact with chemical bone bonding between bioactive surface and bone. Note the gradual transformation between bone and implant material.

bone close to the glass, running a short distance into the reaction layer and within the glass [40]. It has been reported that the bond strength between bone and bioactive glass is higher than that of hydroxyapatite and bone [41, 42].

8.4

The biocompatibility of bioactive glasses

Bioactive glasses behave as biocompatible materials that are independent of the implantation site. In bone, they act mainly osteoconductively and there is no real evidence showing any osteoinductive capacity in bioactive glasses [43, 44]. The interaction at the interface cannot be predicted simply by theoretical analysis or calculation. Even the biocompatibility of a given material depends on the implantation site and loading [45]. Only bioactive materials are able to build up a chemical bond with the host bone: they do not create a chronic wound in the bone, while other implant materials cause an inflammatory reaction in accordance with their biocompatibility. Bone substitute material can be osteoinductive or osteoconductive. The two phenomena occur simultaneously in vivo and are not individual phenomena, but are affected by a multitude of factors besides the material itself. The stability of the implanted site is important, as are the biological implantation bed and the immunological reactions caused by the implant material [46]. Banked bone, cortical segments, hydroxylapatites and bioactive glasses and glass-ceramics are osteoconductive: they favor the proliferation of osteoblasts, but do not cause phenotype conversion of cells. With bioactive glasses, fibroblasts are transformed into a nonmitotic state while exposed to the surface of the glass [47].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

196 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

It has also been reported that calcium phosphates influence alkaline phosphatase activity as well as the incorporation of 3H-thymidine in the repairing bone. Furthermore, it has been suggested that there is likely to be a solution-mediated effect on cell proliferation and differentiation [48]. This activation probably occurs as a result of calcium-mediated events associated with intracellular dissolution of phagocytized material. Thus these types of materials do not cause bone formation at soft-tissue sites, but do promote bone formation when implanted in bone tissue. Hench has a similar view on the osteoconduction of bioactive glass, suggesting that bioactive glasses have an effect that differs from simple osteoconduction and is related to the soluble silicon released by the glass. In addition, the osteostimulatory effect of bioactive glass granules has been observed when filling rabbit bone defects with bioactive glass granules of the appropriate size [44].

8.5

The strength of bioactive glass

The clinical use of bioactive glasses and glass-ceramics has been limited by the structural weakness of the glass. In the case of metal prosthesis, there have been a number of technical problems in the coating process, such as crystallization of the coating material during the process and the limited thermal working range of glass, and these have been barrier to the formation of coatings that could function as the bulk glass. Bioactive glasses and glass-ceramics can be used clinically as space fillers or for regenerative purposes. They can protect other materials from corrosion within the body, replace or augment tissues, and replace functioning parts. Bioactive glasses resemble other glasses in terms of their mechanical strength. The modulus of elasticity is high, but due to the amorphous structure of glass, they are brittle and rigid, have low fracture toughness and are mechanically weak. Critical crack size is small, and the tensile bending strength is 40 to 60 MPa depending on the composition of the glass. When used as fillers, as buried implants or in unloaded applications, these mechanical properties are not harmful [24, 31]. The mechanical properties of bioactive glass together with those of cancellous and cortical bone are presented in Table 8.2.

Table 8.2 Mechanical strength of cancellous and cortical bone, and glass Cancellous bone

Cortical bone

Glass

2–12 MPa

100–230 MPa

785 MPa

Tensile strength

10–20 MPa

50–150 Mpa

35–175 MPa

Young’s modulus

0.05–0.5 MPa

7–30 MPa

70 000 MPa

Compressive strength

Pulling strength Poisson’s ratio

94 MPa 0.2–0.3


0.218


8.6

197

Bone formation

8.6.1 Histology Four main types of tissue can be found at the interface of implanted materials: soft tissue, osteoid, chondroid and bone [15]. Three types of tissue will develop at the interface between bone and bioactive glass: bone, fibrous, and bone marrow tissue. Only mild inflammatory responses have been observed around implanted bioactive glass with a reaction similar to that which occurs in empty defects after surgery. Some round cells and occasionally polynuclear cells can be seen early after implantation, but this reaction will subside rapidly. The reaction layer increases in thickness with time, and will be gradually replaced by bone.

8.6.2 Cartilage repair There is no evidence that cartilage tissue can grow on the glass surface. It seems, however, that a growing bone layer can support cartilage-like tissue, at least at the margins of osteochondral defects, as long as initial subchondral bone repair has occurred. Hyaline-like cells can be observed in the cartilage in the process of repair when bone is growing on the surface of bioactive implants [29]. For traumatic cartilage lesions and for osteochondritis dissecans, one potentially advantageous approach would be to reconstruct subchondral bone first in order to lay the foundations for cartilage healing. It is obvious that large osteochondral defects require a solid surface or a support on which cartilage cells are able to spread, allowing cartilage repair to occur [49–53].

8.6.3 Resistance to toxic effects Polymethylmethacrylate (PMMA) interrupts bone formation at the interface of bone and glass; however glass can withstand this effect better than hydroxylapatite. The surface reaction layer of a bioactive glass shows more resistance to the toxic effect of PMMA. One reason for this might be that the structure of HCA formed on the bioactive glass surface bears a closer resemblance to bone mineral than to synthetic HA [54]. In addition, the alkaline surface of bioactive glass [55] might be better able to buffer the toxic effects. The hydration layer of bioactive glass is one further possible explanation for the resistance of bioactive glass to the toxic effect of PMMA.

8.6.4 Bone formation pattern in defects Experimental defects filled with bioactive glass granules show rapid bone growth on the granules. The pattern of bone growth is centripetal starting from the margins of the defect and growing inwards. The granules did not disturb bone formation in experimental defects, and indeed the filler effect of bone together with bioactive


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

198 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

glass was better than that of bone alone. The centripetal growth pattern has also been observed in clinical use, when benign tumors and metaphyseal fracture defects have been filled with bioactive glass granules. Bone formation from the periphery of the defect [1, 56] is a feature common to all osteoconductive materials, but the combination of bone substitutes with bone inductors might enable more rapid healing of the defects as suggested by Miller et al. [56]. For bioactive glass, an osteoconductive growth pattern [58] and the time-related increase in the amount of bone at the interface has been reported [23, 42, 49]. The bone formation pattern at the interface of bioactive glass resembles intramembranous ossification, while enchondral ossification is absent. This finding leads to the conclusion that bioactive glass does not induce massive bone formation, but instead causes osteoblasts and osteocytes to spread along the glass surface, and that the material is mainly osteoconductive. This view has also is also shared by Hench and Paschall [35] and Ono et al. [57]. The mRNA measurements of type I and III collagen using Northern hybridization provide further evidence for this hypothesis [43]. Nevertheless, some indirect signs that bioactive glass also has bone inductive properties were observed during the experiments discussed. Similar observations have also been published by Ducheyne and Cuckler [60] and Schepers et al. [44].

8.6.5 Bone formation imaging methods In order to image the morphology of the interface various methods must be used. A combination of bulk glass, reaction layer, Ca,P-layer, crystallized Ca,P and mineralizing bone form a complex structure. Histological stains visualize the biological structures. Using toluidine blue and the von Kossa method, the reaction layer can also be stained in a particular way. Although the true nature of these stainings in the reaction layer remain unexplained, these staining characteristics are directly related to the bone bonding of the glass. The von Kossa method stains calcium, and it is likely that there is a layer of free calcium, deep in the reaction layer, that can be visualized using this method (Fig. 8.3). Toluidine blue stains the reaction layer differently and stains the acid part of the layer blue (Fig. 8.4). The staining intensity reflects the surface area of the structures, which has been shown to be important for bone bonding [61]. The formation of the reaction and bone bonding can be analysed by light microscopy and simple histochemical stainings. These histological methods are simpler and less expensive than electron microscopy (EM), scanning electron microscopy (SEM) or energy dispersive X-ray analysis (EDXA), and certainly simpler than push-out tests. Toluidine blue reveals that the interface has a double-layered staining character, and a similar staining pattern occurs when the von Kossa method is used: the former indicates the surface area of the reaction layer and the latter the calciumion distribution. Both are necessary for bonding. Van Gieson, a collagen stain,



199

8.3 Von Kossa staining of bioactive glass (bg) to bone (b) interface (x 75). Note the double-layered structure. Calcium phosphate at the outer part of the silica gel layer stains dark brown (long arrows); the base of the silica gel layer stains dark brown (short arrows).

8.4 Histology of bone (b) to bioactive glass (bg) interface. Toluidine blue (x 150) staining two parallel stripes (arrows) indicating high surface area and acid proteoglycans.

failed to stain the reaction layer, indicating that no collagen is present on the reaction layer side of the interface. The bone-bonding phenomenon of bioactive glass occurs owing to a set of complex reactions, which are interdependent. Consequently, the methods used to study the rate of bone bonding must also be manifold. SEM is not able to image the biological structures, but produces a highly magnified image of the glass, the reaction layer and mineralized bone. EDXA detects the elements at selected spots or linearly, and reflects the chemical


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

200 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

continuum between bone and the implant. Nevertheless, these methods cannot prove that bone bonding occurs: they simply suggest that it does. To definitively show and measure chemical bonding, push-out tests [20] must be combined with these methods.

8.7

Clinical use for benign bone tumors

For the treatment of benign bone tumors many different methods have been used: curettage with or without bone grafting, percutaneous sclerotherapy, corticosteroid injections, total or subtotal excision with or without reconstruction and irradiation. Some tumors disappear spontaneously, for example after fracture healing. Autograft has remained the gold standard filler material. It is osteoinductive without causing negative effects at the grafting site. However, sometimes the quantity of autogenous bone is insufficient, and especially in old or osteoporotic patients the density of bone is too low for packing in the large bone defects. It necessitates a second incision and further operations and can thus cause morbidity and complications. Among many other materials, bioactive glass granules have been used as bone substitute material. Benign bone tumors are contained lesions with intact bone boundaries, surrounded by sclerotic bone. The bone is normally sufficiently strong even after evacuating the tumor. Thus osteosynthesis or external support from outside the bone is not needed. For these reasons granules can be used in any benign bone tumor as filler material. The filler effect during the operation is good, and when granules are moistened with blood or saline implantation into the defect becomes easier [62]. It has been shown that the thickness of cortical bone at the implantation site shows a greater increase when bioactive glass granules are used than when autogenous bone is applied, highlighting the osteostimulative effect of bioactive glass [63]. The use of autogenous bone causes the volume of the cavity to diminish more rapidly than the use of bioactive glass. The growth pattern seems to start from the margins of the defect when osteoconductive bioactive glass is implanted and from the whole defect when autogenous bone is used. After 36 months there seems to be no difference between the two materials in terms of cavity volume, even in the case of large tumors, but there is a significant difference at 12 and 24 months [64]. Bioactive glass is well tolerated and does not put patients at higher risk of infections: no material-related adverse effects have been observed during or after clinical use [62–64]. The inflammatory reaction seems to be similar to that observed when autogenous bone is used. Furthermore, when examining blood samples no differences are observed between these materials [65]. Bioactive glass can be used with good results even in children, without disturbing the remodeling capacity of the bone even in cases where preoperative deformity is present [66]. This is true as long as the growth line is not disturbed either during the operations or by the tumor itself.



201

In certain cases repeat operations have been performed 1 or 2 years after the original operation. The bone between the granules seemed to be physically harder than the surrounding bone. Bioactive glass has also been used in repeat operations, in occurrences of recurrent tumors. The granules disappear gradually by surface reaction and dissolution and by osteoclastic activity in 1 to 4 years depending on the cavity size [64].

8.8

Bioactive glass and infection

Bioactive glass has also been shown to be bacteriostatic in experimental and clinical otorhinological use [67, 68]. Also in the treatment of fractures in the metaphyseal bone region, the granules have shown bacteriostatic properties. In certain cases with a superficial wound infection a revision of the wound has been necessary, leaving the granules open. The wound has healed well without deep infection, despite the direct contact between infected tissues and glass granules [62]. There is some evidence that bioactive glass is advantageous even in chronic osteomyelitic cases [69]. In a retrospective multicenter study, data from 11 patients has been collected and analysed. Postoperative, traumatic or bloodborne chronic osteomyelitis was treated in the spine or lower extremity using bioactive glass granules. The most common cultured pathogens were Staphylococcus aureus and gram-negative bacilli. The cavitary defect or the vicinity of the spinal implant was filled after revision with bioactive glass granules. The results of the surgical intervention and implantation were surprisingly good. It has been postulated that the result is related to the leaching of alkaline earth ions leading to a rapid increase in pH around the glass particles. Bearing the above-mentioned in mind it seems that the aims of the study started by Professor Hench and approved by the US Army Medical Research and Development Command during the Vietnam War seems to have been fulfilled. However, it must be borne in mind that these results are preliminary, and no prospective randomized studies exist concerning the use of bioactive glass in treating osteomyelitis; nor are there any studies dealing with the use of any other materials with the same aim.

8.9

Bioactive glass in cancellous bone and metaphyseal fractures

Fractures in the metaphyseal area often cause compression in the subchondral cancellous bone area, resulting in deformity in the joint line. This type of fracture includes for example distal radius and proximal tibia fractures. After reduction of the joint space a defect in the cancellous bone usually remains, which requires a bone graft. Osteosynthesis is also needed after reduction. In experimental studies [29] bioactive glass provided good support for the growing new bone. In clinical


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

202 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

study, with now more than 10 years of follow-up, results in the bioactive glass group were identical to those in the autogenous bone group [62]. No significant differences were observed between the two demographically identical groups with regard to clinical results, redepression of articular surface measured by plain films and computerized tomography (CT), valgus alignment, tibio-femoral (TF) angle, mechanical axis of lower limb (DMA) or subjective evaluation by the patient. Bioactive glass granules were incorporated into surrounding bone at three months and remodeled thereafter [62, 64]. According to this prospective randomized study it seems that bioactive glass can replace autograft and can be used as bone graft substitute material in metaphyseal fractures. During implantation of the granules meticulous care must be taken to make sure the granules do not enter intra-articular space.

8.10

Bioactive glass in diaphyseal bone fractures

A few studies exist that have examined the use of bioactive glass granules in the treatment of long bone fractures [70]; such as tibia or femoral shaft fractures. These fractures often need repositioning and rigid fixation. Sometimes a large amount of autogenous bone is needed to fill voids caused by these fractures and the use of bone substitute material is advantageous. There is some evidence that bioactive glass causes a thickening of cortical bone in the vicinity of the implantation site, which might help in diaphyseal bone repair. There is normally not as good a vascular supply in diaphyseal bone as there is in metaphyseal bone, especially after the soft tissue damage caused by the fracture. It is therefore to be expected that the reactions of implanted glass are slower, i.e. ossification and the resorption of the granules will take place more slowly.

8.11

Bioactive glass in spinal surgery

To treat instability or in fracture cases spinal fusion is needed to ensure stability between the vertebrae. This can be achieved using anterior or posterior or posterolateral interspinal fusion [71]. One feature common to all methods is the need to build bone in an extra-osseal location. Vascularity can be compromised, and this can negatively affect bone formation. There are experimental and clinical studies concerning bone formation using bioactive glass in posterolateral fusion, which show that a good fusion rate has been achieved. In a clinical study bioactive glass was examined in posterolatreal fusion of spinal fracture patients. The study was designed in such a manner that each patient served as its own control, while autogenous bone was used in the left posterolateral fusion and bioactive glass granules in the right. Similar fusion rate and bone formation was observed in both sides using computerized tomography as the analysis method. It is more difficult to create bone at an extra-osseal than at an intra-osseal implantation site: this is also true when autogenous bone is used. The remodeling



203

of bone tends to restore the original anatomical bone structures and to lead to resorption of the graft, unless the load is carried through the grafted area.

8.12

Arthroplasty

The use of bioactive glass as a coating for metal prosthesis does not appear to be technically possible. However, bioactive glass might be suitable for use in revision arthroplasties, together with autogenous or allogenous bone when filling defects adjacent to metal implants: bactriostatic property of glass is advantageous in this respect, although care must be taken to prevent granules from entering into the joint. No studies have been carried out in this regard.

8.13

Summary of applications in orthopaedics and traumatology

Repairing defects or fractures in metaphyseal areas are an ideal use for bioactive glass. There seems to be no contraindication for use in pediatric orthopaedic surgery or traumatology. In spinal surgery it can be used with good results in posterolateral spondylodesis. The results of the use of granules in the cages for interbody fusion have not been reported. In arthroplasty it is possible to use granules together with autogenous or allograft bone as an extender, though again, no clinical results have been reported. The applicability of bioactive glasses are presented in summary in Table 8.3.

Table 8.3 Clinical applications of bioactive glass granules in orthopaedics Metaphyseal fractures

+++

Diaphyseal fractures

+

Benign bone tumors

+++

Malignant bone tumors

–

Spinal surgery ALIF, PLIF

?

Posterolateral

++

Pediatric use Benign tumors

+++

Trauma

+++

Arthroplasty

+/−

Notes: +++, ++ – proven suitable; + – suitable, not proven; ? – not known; +/– – might be helpful as additive.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

204

Bioactive glasses

8.14

Future trends

The advantage of bioactive glass is the formation of hydroxylapatite within the body. The disadvantages are related to the long resorption time. For this reason bioactive glass is not suitable for clinical application where the filled area needs to be operated on, drilled or resected shortly after the primary operation. Resorption in large defects can take several years, and in smaller defects with diameter approximately 1 cm, up to a year. Any attempt to drill or resect through the implanted area too early will result in splittering and crushing of the formed bone by the harder fragments of unresorbed glass causing damage to the surrounding bone. Economic considerations are leading to an increase in the use of bone substitute materials instead of autogenous bone. The costs of the time in the operating theater are being calculated more accurately. In a situation whereby the cost of extra time in the operating theater is equivalent to the cost of bone substitute material, bone substitute materials will be favored: in a clinical situation these are preferable not only for the patient but also for the surgeon. There are currently a large number of synthetic bone substitute materials on the market. Allograft bone has so far been the most frequently used material. Local hospital bone banks are in the process of changing to more centralized bone banks. At the same time the use of bone substitute materials has increased [72]. The various materials are in competition, but at the same time the marketing of one material supports that of the others. It can reasonably be expected that certain materials will be used in a certain applications, i.e. one single material is not suitable for all applications. It is also to be expected that composite materials [73, 74] and materials embedded with growth factors and even prefabricated bone will appear in the near future [75].

8.15

References

1. Senn N. On the healing of aseptic bone cavities by implantation of antiseptic decalcified bone. J Med Sci (Am) 1889; 98(3):219–43. 2. van Meekeren J. Heel a geneeskonstige aanmerkingen. Amsterdam: Commelijn 1668. 3. Macewen W. Observations concerning transplantation on bone. Proc R Soc Lond 1881; 32:232–47. 4. Goldberg V. M., Stevenson S. Natural history of autografts and allografts. Clin Orthop 1987; 225:7–16. 5. Dreesman H. Read before the Surgical Section of the New York Academy of Medicine, March 13, 1893. 6. Urist M. R. Bone formation by autoinduction. Science 1965; 150:893–99. 7. Albee F. H. Studies in bone growth. Triple calcium phosphate as a stimulus to osteogenesis. Ann Surg 1920; 71:32–36. 8. Orell S. Surgical bone grafting with ‘os purum’, ‘os novum’ and ‘boiled bone’. J Bone Joint Surg 1937; 19:873–85.



205

9. Urist M. R., Strates B. S. Bone formation in implants of partially and wholly demineralized matrix. Clin Orthop 1970; 71:271–78. 10. Urist M. R., Strates B. S. Bone morphogenetic protein. J Dent Res 1971; 50:1392–406. 11. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K. Bonding mechanism at the interface of ceramic prosthetic materials. J Biomed Mat Res Symp 1971; No 2 (Part 1):117–43. 12. Roy D. M., Linnehan S. K. Hydroxyapatite formed from coral skeleton carbonate by hydrothermal exchange. Nature 1974; 247:220–22. 13. Jarcho M. Calcium phosphate ceramics as hard tissue prosthetics. Clin Orthop 1981; 19:152–54. 14. Mittelmeier H., Mittelmeier W. Moderne Entwicklung von Knochenersatzmaterialen. Heft zur Unfallheilkunde, pp. 69–84, Heft 216. Hrsg. A. H. Huggler, E. H. Kuner. Berlin 1991. 15. Gross U. M., Strunz V. Surface staining of sawed sections of undecalcified bone containing alloplastic implants. Stain Technol 1977; 52(4):217–19. 16. Kokubo T., Shigematsu M., Nagashime Y., Tashiro T., Nakamura T., Yamamuro T., Higashi S. Apatite- and wollastonite-containing glass-ceramic for prosthetic application. Kyoto University: Bull Inst Chem Res 1982; 60:260–68. 17. Furlong R. J., Osborn J.-F. Fixation of hip prosthesis by hydroxyapatite ceramic coatings. J Bone Joint Surg 1991; 73B:741–45 18. van Blitterswijk C. A., Bakker D., Leenders H., Brink J. V. D., Hesseling S. C., Bovell Y. P., Radder A. M., Sakkers R. J., Gaillard M. L., Heinze P. H., Beumer G. J. Interfacial reactions leading to bone-bonding with PEO/PBT copolymers (Polyactive®). In P. Ducheyne, T. Kokubo, C. A. van Blitterswijk (eds) Bone-Bonding Biomaterials. Holland: Reed Healthcare Communications 1992; pp. 13–30. 19. Constanz B. R., Ison I. C., Fulmer M. T., Poser R. D., Smith S. T., VanWagoner M., Ross J., Goldstein S. A., Jupiter J. B., Rosenthal D. I. Skeletal repair by in situ formation of the mineral phase of bone. Science 1995; 267:1796–99. 20. Andersson Ö. H. The bioactivity of silicate glass. Thesis. Åbo Akademi University, Turku, Finland 1990. 21. Hench L. L., Ethridge A. C. Biomaterials. An Interfacial Approach. New York: Academic Press, 1982; pp. 279–88. 22. Aebi M. Biologisher oder artifizieller Knochenersatz? Hefte zur Unfallheilkunde, Springler Verlag, Berlin Heft 1991; 216:1–9. 23. Gross U., Strunz V. Interface of various glasses and glass-ceramics of different solubility in the femur of the rat. J Biomed Mater Res 1980; 14:607–18. 24. Kokubo T., Ito S., Shigematsu M., Sakka S., Yamamuro T. Mechanical properties of a new type of apatite-containing glass-ceramic for prosthetic application. J Mat Sci 1985; 20:2001–4. 25. Kudo K., Miyashava M., Fujioka Y., Kamegai T., Nakano H., Seino Y., Ishikawa F., Shioyama T., Ishibashi K. Clinical application of dental implant with root of coated bioglass: short-time results. Oral Surg Oral Med Oral Pathol 1990; 70:18–23. 26. Wolf H., Lorenz T., Benser A., Schubert R., Bertoldt R., Reefschläger J., Berger G., Sauer R. In vitro detection and evaluation of pathophysiological effects of ions released from bioactive glass-ceramics. Biomaterials 3rd World congress. Summaries, abstract and programme, Kyoto, 198827; p. 69; Andersson Ö. H., Guizhi Liu, Karlsson K. H., Niemi L., Miettinen J., Juhanoja J. In vivo behaviour of glasses in the SiO2Na2O-CaO-P2O5-Al2O3-B2O3 system. J Mater Sci: Materials in Medicine 1 1990; 219–227.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

206 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

28. Kangasniemi I. K. Development of CaP-containing bioactive glass composites. Thesis, University of Leiden, 1993. 29. Heikkilä J. T. Bioactive glass in experimental and clinical orthopaedics. Thesis. Universtity of Turku, 1996. 30. Suominen E. Bioactive glass ceramics in reconstruction of bone defects. Studies with bioactive glasses, glass-ceramics, hydroxyapatite and their composites. Thesis, University of Turku, 1996. 31. Hench L. L., Andersson Ö. H. Bioactive glasses. In Hench L. L., Wilson J., eds, An Introduction to Bioceramics. Singapore: Reed Healthcare Communications 1993; pp. 41–62. 32. Ducheyne P., Bianco P., Radin S., Schepers E. Bioactive materials: mechanical and bioengineering conditions. In P. Ducheyne, C. van Blitterswijk, T. Kokubo, eds, Bone-Bonding Biomaterials. Leiderdorp, The Netherlands: Reed Healthcare Communications 1992; pp. 1–12. 34. Wilson J., Nolletti D. Bonding of soft tissues to bioglass. In P. Christel, A. Meunier, A. J. C. Lee, eds, Biological and Biomechanical Performance of Biomaterials. Amsterdam: Elsevier Science Publishers 1986; pp. 99–104. 35. Hench L. L., Paschall H. A. Direct chemical bond of bioactive glass-ceramic materials to bone and muscle. J Biomed Mater Res Symp 1973; 4:25–42. 36. Clark A. E., Pantano C. G., Hench L. L. Auger spectroscopic analysis of Bioglass corrosion films. J Amer Ceram Soc 1976; 59(1–2):37–39. 37. Wilson J., Pigott G. H., Schoen F. J., Hench L. L. Toxicology and biocompatibility of bioactive glasses. J Biomed Mater Res 1981; 15:805–17. 38. Andersson Ö. H., Vähätalo K., Yli-Urpo A., Happonen R.-P., Karlsson K. H. Shortterm reaction kinetics of bioactive glass in simulated body fluid and in subcutaneous tissue. In Ö. H. Andersson, A. Yli-Urpo, eds, Bioceramics 7. Butterworth-Heinemann Ltd 1994; pp. 67–72. 39. Andersson Ö. H., Karlsson K. H., Kangasniemi K. Calcium phosphate formation at the surface of bioactive glass in vivo. J Non-Cryst Solids 1990; 119:290–96 40. Andersson Ö. H., Karlsson K. H. Models for physical properties and bioactivity of phosphate opal glasses. Glastechnische Berichte 1988; 61:300–5. 41. Fujiu T., Ogino M. Difference of bond bonding behavior among surface active glasses and sintered apatite. J Biomed Mater Res 1984; 18:845–59. 42. Hench L. L. Bioactive ceramics. Ann NY Acad of Sci 1988; 523:54–71. 43. Virolainen P., Heikkilä J. T., Vuorio E., Aro H. T. Healing of cortical bone defects filled with autograft or with bioactive glass granules. J Biomed Mat Res. In press. 44. Schepers E., de Clercq M., Ducheyne P., Kempeneers R. Bioactive glass particulate material as a filler for bone lesions. J Oral Rehab 1991; 18:439–52. 45. Williams D. F. Titanium: epitome of biocompatibility or cause for concern. Editorial. J Bone Joint Surg (Br) 1994; 76:348–49. 46. Aebi M., Regazzoni P., Schwarzenbach O. Segmental bone grafting: comparison of different types of grafts in dogs. Int Orthop 1989; 13:1–11. 47. Seitz T. L., Noonan K. D., Hench L. L., Noonan N. E. Effect of fibronectin on the adhesion of an established cell line to a surface reactive biomaterial. J Biomed Mat Res 1982; 16(3):195–207. 48. Orly I., Gregoire M., Menenteau J., David M. Effects of synthetic calciumphosphates on 3H-thymidine incorporation and alkaline phosphatase activity of human fibroblast culture. J Biomed Mater Res 1989; 23:1433–40.



207

49. Chiroff R. T., White R. A., White E. W., Weber J. N., Roy D. The restoration of articular surfaces overlying replamineform porous biomaterials. J Biomed Mater Res 1977; 11:165–78. 50. Muckle D. S., Minns R. J. Biological response to woven carbon fibre pads on the knee. J Bone Joint Surg (Br) 1989; 71B:60–62. 51. Aston J. E., Bentley G. Repair of normal and arthritic articular surfaces by allografts of articular and growth plate cartilage. J Bone Joint Surg (Br) 1986; 68B:29–35. 52. Billings E. Jr, von Schroeder H. P., Mai M. T., Aratow M., Amiels D., Woo S. L. Y., Coutts R. D. Cartilage resurfacing of the rabbit knee. Acta Orthop Scand 1990; 61(3):201–6. 53. Dahlberg L., Kreicbergs A. Demineralized allogeneic bone matrix for cartilage repair. J Orthop Res 1991; 9(1):11–19. 54. Rehman I., Hench L. L., Bonfield W., Smith R. Analysis of surface layers on bioactive glasses. Biomaterials 1994; 15:865–70. 55. Hench L. L., Wilson J. Biocompatibility of silicates for medical use. Pp. 231–46. In Ciba Foundation Symposium, Silicon Biochemistry. Chichester: John Wiley & Sons 1986: 121. 56. Miller T. A., Ishida K., Kobayashi M., Wollmann J. S., Turk A. E., Holmes R. E. The induction of bone by an osteogenic protein and the conduction of bone by porous hydroxyapatite: a laboratory study in the rabbit. Plast Rec Surg 1991; 87(1): 87–95. 58. Andersson Ö. H., Liu G., Kangasniemi K., Juhanoja J. Evaluation of the acceptance of glass in bone. J Mater Sci: Mater Med 1992; 3:145–50. 59. Ono K., Yamamuro T., Nakamura T., Kokubo T. Quantitative study on osteoconduction of apatite-wollastonite containing glass ceramic granules, hydroxyapatite granules and alumina granules. Biomaterials 1990; 11:265–71. 60. Ducheyne P., Cucler J. Bioactive prosthetic coatings. Clin Orthop 1992; 276:102–14. 61. Walker M. M. An investigation into the bonding mechanism of bioglass. MSc Thesis University of Florida, 1977. 62. Heikkilä J. T., Kukkonen J., Aho A. J., Moisander S., Mattila K. Bioactive glass granules – a suitable bone substitute material in the operative treatment of depressed tibial plateau fractures. A prospective randomized one year follow-up study. Submitted to J Mater in Med: Mater Sci. 63. Lindfors N. C., Heikkilä J. T., Koski I., Mattila K., Aho A. J. Bioactive glass and autogenousbone as bone graft substitutes in benign bone tumors. J Biomed Mater Res B Appl Biomater 2009; Jul 90(1):131–6. 64. Lindfors N. C., Koski I., Heikkilä J. T., Mattila K., Aho A. J. A prospective randomized 14-year follow-up study of bioactive glass and autogenous bone as bone graft substitutes in benign bone tumors. J Biomed Mater Res B Appl Biomater 2010; Jul 94(1):157–64. 65. Lindfors N. C., Heikkilä J. T., Aho A. J. Long-term evaluation of blood silicon and ostecalcin in operatively treated patients with benign bone tumors using bioactive glass and autogenous bone. J Biomed Mater Res B Appl Biomater 2008; Oct 87(1): 73–6. 66. Lindfors N. C. Treatment of a recurrent aneurysmal bone cyst with bioactive glass in a child allows for good bone remodelling and growth. Bone 2009; Aug 45(2):398–400. Epub 2009, Apr 18. 67. Stoor P. Antibacterial effects of a bioactive glass paste on oral microorganisms. Acta Odontologica Scandinavica 1998; 56(3):161–165.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

208 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

68. Stoor, Patricia. Interactions between oral and nasal microorganisms and the bioactive glass S53P4 with special reference to nasal cavity surgery. Thesis, University of Turku, 2006. 69. Lindfors N. C., Hyvönen P., Nyyssönen M., Kirjavainen M., Kankare J., Gullichsen E., Salo J. Bioactive glass S53P4 as bone graft substitute in treatment of osteomyelitis. Bone 2010; Aug 47(2):212–18. 70. Sponer P., Urban K., Urbanová E., Karpas K., Mathew P. G. Behavior of bioactive glass-ceramic implanted into long bone defects: a scintigraphic study. J Pediatr Orthop B. 2010; Jan 19(1):102–7. 71. Acharya N. K., Kumar R. J., Varma H. K., Menon V. K. Hydroxyapatite-bioactive glass ceramic composite as stand-alone graft substitute for posterolateral fusion of lumbar spine: a prospective, matched, and controlled study. J Spinal Disord Tech 2008; Apr 21(2):106–11. 72. Nagata M. J., Furlaneto F. A., Moretti A. J., Bouquot J. E., Ahn C. W., Messora M. R., Fucini S. E., Garcia V. G., Bosco A. F. Bone healing in critical-size defects treated with new bioactive glass/calcium sulfate: a histologic and histometric study in rat calvaria. J Biomed Mater Res B Appl Biomater 2010; Nov 95(2):269–75. 73. Hautamäki M. P., Aho A. J., Alander P., Rekola J., Gunn J., Strandberg N., Vallittu P. Repair of bone segment defects with surface porous fiber-reinforced polymethyl methacrylate (PMMA) composite prosthesis. Acta Orthop Scand 2008; 79(4): 555–564. 74. Hautamäki M., Meretoja V. V., Mattila R. H., Aho A. J., Vallittu P. K. Osteoblast response to polymethyl methacrylate bioactive glass composite. J Mater Sci: Mater Med 2010; May 21(5):1685–92. 75. Xia W., Chang J. Bioactive glass scaffold with similar structure and mechanical properties of cancellous bone. J Biomed Mater Res B Appl Biomater 2010; Nov 95(2):449–55.


9 Bioactive glass S53P4 as a bone graft substitute in the treatment of osteomyelitis N. C. LINDFORS , Helsinki University Central Hospital, Finland

Abstract: Bioactive glasses (BAGs) are bone substitutes with bone-bonding, angiogenesis-promoting and antibacterial properties. The bioactive process that leads to bone bonding is described through reactions at the glass surface. This chapter describes the chemical reactions following implantation of the glass that lead to the formation of silanol (SiOH) groups at the glass surface and the subsequent formation of a CaO-P2O5 hydroxyapatite (HA) layer on top of the Si-rich layer. Finally, cell interactions with the HA layer are shown to initiate the bone-forming pathway. The chapter also explains and compares the antibacterial properties of different bioactive glasses, showing BAG-S53P4 to be the most effective against significant pathogens and bacteria. These findings are demonstrated in a multicentre trial involving eleven patients of osteomyelitis treated with BAG-S53P4 implants, which demonstrated the good grafting and antibacterial properties of S53P4 as an implant material and bone graft substitute. It has since been used successfully in further operations. Key words: bioactive glass, osteomyelitis, S53P4, bone subsitute, infection.

9.1

Introduction

Osteomyelitis is caused by infected micro-organisms and defines a destructive inflammatory process in bone that is often accompanied by bone destruction (Lazzarini et al., 2004). Osteomyelitis is heterogenous in its pathophysiology, clinical presentation, and management. It is often considered to be the most difficult-to-treat infectious disease. The infection can arise from a variety of aetiologies (Lew and Waldfogel, 2004). Most often it is caused by trauma, but any kind of bone or soft tissue surgery where pathogens can enter the bone may cause the infection. Haematogenous osteomyelitis has been found in children, as well as in elderly patients (Riise et al., 2008); and in diabetic patients, osteomyelitis may appear as a secondary manifestation due to vascular insufficiency and soft tissue infection (Haartemann-Heurtier and Senneville, 2008). The most common pathogens causing osteomyelitis are Staphylococcus aureus and Gram-negative bacilli (Parsson and Strauss, 2004). Acute osteomyelitis is characterized by oedema, locally decreased blood supply and pus formation. Untreated or due to treatment failure, the infection can progress to a more chronic phase, with formation of a sequestrum, a large area of 209 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

210 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

devascularized dead bone. In treatment of chronic osteomyelitis, debridement of the dead bone is mandatory. This procedure unfortunately, often results in a poorly vascularized large bone defect, a dead space. Bacterial infection can also cause local acidosis, leading to dissolution of bone matrix mineral (Konttinen et al., 2001). Many different methods have been used to treat osteomyelitis, including antibiotic-impregnated polymethyl methacrylate (PMMA) beads, free vascularized bone grafts, local muscle flaps, granulation formation according to the technique of Papineau and the Masquelet technique (Powerski et al., 2009) or bone reconstruction based on Ilizarov technology (Parsson and Strauss, 2004).

9.2

Bone grafts in the treatment of osteomyelitis

Bacterial colonization of implanted materials is promoted by the ability of the bacteria to produce protein-specific adhesins on their surfaces. This is followed by interactions with host protein components, such as fibrinogen, fibronectin and collagen. Bacteria communicate through hormone-like compounds in biofilms, making treatment with antimicrobial agents difficult (Lew and Waldfogel, 2004). Therefore, synthetic bone grafts are generally not recommended in treatment of osteomyelitis. Debridement in combination with the local administration of antibiotics, e.g. gentamicin-loaded PMMA beads, has become the method of choice in treating osteomyelitis. However, PMMA is known to provide a favourable environment for proliferation of bacteria (Boyd and Towler, 2005). In a long-term follow-up of 100 patients treated with gentamicin-PMMA beads, relapses occurred for 8.8% of patients with acute osteomyelitis and for 21.2% of patients with chronic osteomyelitis (Walenkamp et al., 1998). Patients with chronic osteomyelitis treated with biodegradable calcium sulphate tobramycin-impregnated pellets or calcium sulphate tobramycin– vancomycin-impregnated pellets have shown excellent osseous repair (Gitelis and Brebach, 2002). However, an increase in antibiotic-resistant bacteria, such as gentamicin- or methicilline-resistant Staphylococcus aureus, has been observed (Efstathopoulos et al., 2008). The prevention of bacterial proliferation due to ion release has been demonstrated for glass polyalkenoate cements (Wren et al., 2009). The effectiveness of a degradable and bioactive borate glass as a carrier for vancomycin has been compared with calcium sulphate in treatment of osteomyelitis in rabbits. At eight weeks, vancomycin-loaded borate glass was found to be effective in eradicating osteomyelitis caused by methicillin-resistant Staphylococcus aureus (MRSA), and the treated region was mostly reabsorbed and replaced with new bone. Treatment with pure borate glass was significantly less effective in eradicating MRSA (Zongping et al., 2009).


Bioactive glass S53P4 in the treatment of osteomyelitis

9.3

211

Antibacterial properties of bioactive glass S53P4

Bioactive glasses (BAGs) are synthetic biocompatible bone-bonding osteoconductive bone substitutes with documented antibacterial and angiogenesispromoting properties (Hench and Pashall, 1973; Hench and Wilson, 1984; Hench, 1988; Andersson et al., 1990; Andersson and Kangasniemi, 1991; Lindfors and Aho, 2003; Day, 2005; Leppäranta et al., 2008; Munukka et al., 2008). The bioactive process leading to bone bonding has been described as a sequence of reactions in the glass and at its surface. Implantation of the glass is followed by a rapid exchange of Na+ in the glass with H+ and H3O+ from the surrounding tissue, leading to the formation of silanol (SiOH) groups on the glass surface. After repolymerization, a SiO2-rich layer is formed. Due to migration of Ca2+ and PO43− groups to the surface and cystallization, a CaO-P2O5 hydroxyapatite (HA) layer is formed on top of the Si-rich layer. Finally, cell interactions with the HA layer subsequently initiate the bone forming pathway (Hench and Wilson, 1984; Andersson et al., 1990). The initial leaching of alkali and alkaline earth ions lead to a rapid increase in pH around the glass, which depends on the composition of the glass. BAG-S53P4 has in a simulated body fluid shown an increased pHmax value of 11.65 (Zhang et al., 2006). It has been suggested that the antibacterial properties observed for BAGs are caused by the high pH and the subsequent osmotic effect caused by dissolution of the glass (Stoor et al., 1998). This is confirmed by the observation that neutralization of a highly alkaline solution with BAG eliminates the antibacterial effect (Allan et al., 2001). BAG-S53P4 has presented effective bacterial growth-inhibiting properties in vitro, towards 17 anaerobic bacteria, as well as 29 clinically important aerobic bacteria. Comparing bactericidal effects of different BAGs, BAG-S53P4 has been shown to possess the fastest bacterial growth inhibitory effect (Leppäranta et al., 2008; Munukka et al., 2008). Previous studies on atrophic rhinitis often caused by Klebsiella ozaenae have shown that BAG-S53P4 does not favour adhesion or colonization of K. ozaenae on its surface. Neither can K. ozaenae form biofilms on BAG-S53P4 (Stoor et al., 1999).

9.4

Vascularization-promoting properties of bioactive glasses

Vascularization plays an important role in the bone tissue healing process, and therefore, vascularization of the poorly vascularized dead space of the bone cavity and the surrounding tissue is vital in treating osteomyelitis. The vascular endothelial growth factor (VEGF) has been successfully used in preclinical ischaemic tissue models to enhance and promote the development of collateral blood vessels (Banai et al., 1994; Pearlman et al., 1995; Day, 2005). BAG-45S5 Bioglass® has been shown to stimulate release of angiogenetic growth factors and to promote angiogenesis. An increase in tubule branching and the


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

212 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

formation of complex networks of interconnected tubules have been observed after addition of a fibroblast-conditioned medium, produced in the presence of 45S5 Bioglass® (Day, 2005). Soluble products of 45S5 Bioglass® inducing endothelial cell proliferation and up-regulation of VEGF production, observed for 45S5 Bioglass®, also indicate that 45S5 Bioglass® possesses a proangiogenic potential (Leu and Leach, 2008). A material-dependent angiogenetic response has also been demonstrated in an in vivo rat critical-size defect model. Significantly enhanced mitogenic stimulation of endothelial cells with an additive effect with VEGF release was observed in the presence of a BAG coating (Leach et al., 2006). Vascularization and new bone formation have been observed to be faster in defects filled with BAG-S53P4 than in hydroxyapatite-filled defects. Initial fibrous tissue formation related to a considerable amount of blood vessels has also been observed to be more rapid in BAG filled defects (Peltola et al., 2001).

9.5

Bioactive glass S53P4 in the treatment of osteomyelitis: a multicentre study

The aim of the study was to apply the experimentally known antibacterial properties of BAG-S53P4 to clinical practice, evaluating the operative outcome of using BAG-S53P4 as a bone graft substitute in treating osteomyelitis (Lindfors et al., 2010).

9.5.1 Patients and methods Eleven patients (nine males, two females) with radiologically diagnosed osteomyelitis participated in a multicentre study in Finland. Osteomyelitis was localized in the lower extremity in ten cases and in the spine in one case. Seven of the patients had sustained a fracture: in the distal tibia (three patients), in the calcaneus (two patients), in the distal fibula (one patient) and in the distal femur (one patient). Nine patients had undergone previous operative treatments, including revisions, osteotomies and artrodesis. Autologous bone grafts had been used in two patients and a bone substitute (Norian®) in one patient. Kanamycin granules had been used in one patient and Garamycin granules (Septocol®) in two patients. Antibiotic therapies had been given to all patients. One patient had been treated for osteomyelitis for 64 years, four patients for 7 to 16 years and six patients for ∼1 to 2 years. Between 2007 and 2009, all the patients were operated on due to chronic infection and verified osteomyelitis. In the operation, the infected bone and the soft tissue were removed, and the cavitary bone defects were filled with BAGS53P4 (BonAlive®, Bonalive Biomaterials Ltd., Finland). The whole cavitary defect was filled and, therefore, the amount of glass used depended on the size of the cavity. In four patients, muscle flaps were used as part of the treatment.



213

A patient with verified spondylitis was treated using a metal implant, which was covered with BAG-S53P4. The most common pathogen causing the infection was Staphylococcus aureus (six patients). The outcome of the treatment was evaluated by the surgeon as excellent (no complications), good (a small complication) or a temporary stable situation. Patients were seen at the outpatient departments at 1, 2, 3 to 4, and 6 to 15 months postoperatively. Five patients had a follow-up of 2 to 6 months and six patients of 8 to 15 months. Patient data were also obtained from hospital patient records until June 2010, resulting in a mean follow-up period of 27 months (range 13–41).

9.5.2 Results BAG-S53P4 was well tolerated. The use of BAG-S53P4 as a bone graft substitute resulted in a fast recovery, i.e. patients that had been treated with long-lasting therapies responded well to the treatment. Clinical outcome was good or excellent in nine of eleven patients. Postoperative complications needing treatment were seen in two patients. In one patient, vascular problems occurred in the muscle flap. In another patient, a postoperative complication was observed one month after treatment due to the fact that the evacuated cavity had not been properly filled with BAG. During arthroscopic revision it was observed that the empty part of the treated cavity was filled with a haematoma, which was considered to be the cause of the reinfection. According to the patients’ records, no relapses or other complications were observed. The preoperative and postoperative radiological appearance of the treated bone cavity in the distal tibia is shown in Fig. 9.1 (a)–(c).

9.6

Conclusions

Patients who suffer from osteomyelitis differ in the pattern of illness. The study showed that BAG-S53P4 could successfully be used as a bone substitute in treating osteomyelitis independent of aetiology, pathogens, localization or previous treatment of the infection. BAG-S53P4 resorbs slowly and is replaced by new bone in a process that takes many years. The slow constant resorption of BAG-S53P4 ensures that no dead space will be formed during the healing process as active bone formation takes place at the defect site. The healing process progresses from a fibrous tissue phase to bone formation with scattered fibrous tissue and bone obliteration maintaining BAG granules (Peltola et al., 2006). BAG-S53P4 was used with good results in an one-stage procedure in six patients, although pus was observed in some of the patients’ surrounding tissues.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

214 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

9.1 (a)–(c) Osteomyelitis caused by Staphylococcus aureus in distal tibia treated with BAG-S53P4 as bone graft substitute: (a) preoperative MRI showing osteomyelitis in tibia, (b) postoperative X-ray showing BAG-S53P4 in the treated bone cavity (arrow), and (c) X-ray at five months’ follow-up showing the treated region (arrow) (Pekka Hyvönen, Department of Orthopaedics and Traumatology, Oulu University Hospital, Oulu, Finland).



215

Using BAG-S53P4 as a bone substitute in a one-stage procedure, with no second operation required and no harvesting of AB from the iliac crest, makes BAGS53P4 a cost-effective, as well as a rapid method of treating osteomyelitis. Longer follow-ups are, however, needed to verify the long-term beneficial outcome of the treatment.

9.7

References

Allan I., Newman H., Wilson M. (2001), ‘Antibacterial activity of particulate Bioglass against supra- and subgingival bacteria’, Biomaterials, 22, 1683–1687. Andersson Ö. H., Karlsson K. H., Kangasniemi K. (1990), ‘Calcium phosphate formation at the surface of bioactive glass in vivo’, Journal of Non-Crystal Solids, 119, 290–296. Andersson Ö. H., Kangasniemi I. (1991), ‘Calcium phosphate formation at the surface of bioactive glass in vitro’, J Biomed Mater Res, 24, 1019–1030. Banai S., Jaklitsch M. T., Shou M., Lazarous D. F., Scheinowitz M., Biro S., Epstein S. E., Unger E. F. (1994), ‘Angiogenic-induced enhancement of collateral blood flow to ischemic myocardium by vascular endothelial growth factor in dogs’, Circulation, 89, 2183–2189. Boyd D., Towler M. R. (2005), ‘The processing, mechanical properties and bioactivity of zinc based glass ionomer cements’, J Mater Sci: Mater Med, 16, 843–850. Day R. M. (2005), ‘Bioactive glass stimulates the secretion of angiogenetic growth factors and angiogenesis in vitro’, Tissue Eng, 11, 768–777. Efstathopoulos N., Giamarellos-Bourboulis E., Kanellakopoulou K., Lazzarettos I., Giannoudis P., Frangia K., Magnissalis E., Papadaki M., Nikolaou V. S. (2008), ‘Treatment of experimental osteomyelitis by Methicillin Resistant Staphylococcus Aureus with bone cement system releasing grepafloxacin’, Injury, 39, 1384–1390. Gitelis S., Brebach G. T. (2002), ‘The treatment of chronic osteomyelitis with a biodegradable antibiotic-impregnated implant’, J Orthop Surg, 10, 53–60. Haartemann-Heurtier A., Senneville E. (2008), ‘Diabetic foot osteomyelitis’, Diabetes and Metabolism, 34, 87–95. Hench L. L., Paschall H. A. (1973), ‘Direct chemical bond of bioactive glass-ceramic materials to bone and muscle’, J Biomed Mater Res Symp, 4, 25–42. Hench L. L., Wilson J. (1984), ‘Surface active biomaterials’, Science, 226, 630–635. Hench L. L. (1988), ‘Bioactive ceramics’, Ann N Y Acad Sci, 523, 54–71. Konttinen Y. T., Tagaki M., Mandelin J., Lassus J., Salo J., Ainola M., Li T. F., Virtanen I., Liljeström M., Sakai H., Kobayashi Y., Sorsa T., Lappalainen R., Demulder A., Santavirta S. (2001), ‘Acid attack and cathepsin K in bone resorption around total hip replacement prosthesis’, J Bone Miner Res, 16, 1780–1786. Lazzarini L., Mader J. T., Cahoun J. H. (2004), ‘Osteomyelitis in long bones’, J Bone Joint Surg, 86-A, 2305–2318. Leach J. K., Kaigler D., Wang Z., Krebsbach P. H., Moonley D. J. (2006), ‘Coating of VEGF-releasing scaffolds with bioactive glass for angiogenesis and bone regeneration’, Biomaterials, 27, 3249–3255. Leppäranta O., Vaahtio M., Peltola T., Zhang D., Hupa L., Hupa M., Ylänen H., Salonen J. I., Viljanen M. K., Eerola E. (2008), ‘Antibacterial effect of bioactive glasses on clinically important anaerobic bacteria in vitro’, J Mater Sci: Mater Med, 19, 547–551.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

216 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Leu A., Leach L. K. (2008), ‘Proangiogenic potential of a collagen/bioactive substrate’, Phar Res, 25, 1222–1229. Lew D. P., Waldfogel F. A. (2004), ‘Osteomyelitis’, Lancet, 364, 369–379. Lindfors N. C., Aho A. J. (2003), ‘Granule size and composition of bioactive glass affect osteoconduction in rabbit’, J Mater Sci: Mater Med, 14, 265–372. Lindfors N., Hyvönen P., Nyyssönen M., Kirjavainen M., Kankare J., Gullichsen E., Salo J. (2010), ‘Bioactive glass S53P4 as bone graft substitute in treatment of osteomyelitis’, Bone, 47, 212–218. Munukka E., Leppäranta O., Korkeamäki M., Vaahto M., Peltola T., Zhang D., Hupa L., Ylänen H., Salonen J. I., Viljanen M. K., Eerola E. (2008), ‘Bactericidal effects of bioactive glasses on clinically important aerobic bacteria’, J Mater Sci: Mater Med, 19, 27–32. Parsson B., Strauss E. (2004), ‘Surgical management of chronic osteomyelitis’, Am J Surg, 188, 57–66. Pearlman J. D., Hibberd M. G., Chuang M. L., Harada K., Lopez J. J., Gladstone S. R., Friedman M., Sellke F. W., Simons M. (1995), ‘Magnetic resonance mapping demonstrates benefits of VGEF-induced myocardial angiogenesis’, Nat Med 1, 1985–1989. Peltola M. J., Aitasalo K. M. J., Suonpää J. T. K., Yli-Urpo A., Laippala P. J. (2001), ‘In vivo model for frontal sinus and calvarian bone defect obliteration with bioactive glass S53P4 and hydroxyapatite’, J Biomed Mater Res Appl Biomater, 58, 261–269. Peltola M., Aitasalo K., Suonpää J., Varpula M., Yli-Urpo A. (2006), ‘Bioactive glass S53P4 in frontal sinus obliteration: A long-term clinical experience’, Head and Neck, 28, 834–841. Powerski M., Maier B., Frank J., Marzi I. (2009), ‘Treatment of severe osteitis after elastic intramedullary nailing of a radial bone shaft fracture by using cancellous bone graft in Masquelet technique in a 13-year-old adolescent girl’, J Pediatr Surg, 44, 17–19. Riise Ø. R., Kirhus E., Handeland K. S., Flatø B., Reiseter T., Cvancarova M., Nakstad B., Wathne K-O. (2008), ‘Childhood osteomyelitis-incidence and differentiation from other acute onset musculoskeletal features in a population-based study’, Pediatrics, 8, 45–55. Stoor P., Söderling E., Salonen J. I. (1998), ‘Antibacterial effects of a bioactive glass paste on oral microorganisms’, Acta Odontol Scand, 56, 161–165. Stoor P., Söderling E., Grenman R. (1999), ‘Interactions between the bioactive glass S53P4 and the atrophic rhinitis-associated microorganism Klebsiella ozaenae’, J Biomed Mater Res Appl Biomater, 48, 869–874. Walenkamp G. H., Kleijn L. L., de Leeuw M. (1998), ‘Osteomyelitis treated with gentamicin-PMMA beads: 100 patients followed for 1–12 years’, Acta Orthop Scand, 69, 518–522. Wren A. W., Boyd D., Thornton R., Cooney J. C., Towler M. R. (2009), ‘Antibacterial properties of a tri-sodium citrate modified glass polyalkenoate cement’, J Biomed Mater Res Appl Biomater, 90, 700–709. Zhang D., Munukka E., Leppäranta O., Hupa L., Ylänen H., Salonen J., Eerola E., Viljanen M. K., Hupa M. (2006), ‘Comparison of antibacterial effect of three bioactive glasses’, Key Eng Mat, 309–311; 345–348. Zongping Xie, Xin Liu, Weitao Jia, Changqing Zhang, Wenhai Huang, Jianqiang Wang (2009), ‘Treatment of osteomyelitis and repair of bone defect by degradable bioactive borate glass releasing vancomycin’, J Con Rel, 139, 118–126.


10 Bioactive glass for maxillofacial and dental repair M. J. PELTOLA and K. M. J. AITASALO , Turku University Hospital, Finland

Abstract: The use of biomaterials is increasing in treatments today and will increase further in the future. The clinical need for biomaterials in maxillofacial and oral cavity areas is both demanding and versatile. The chapter first discusses key facts and applications of traditional reconstruction materials, also offering comparative data. It then describes the properties and clinical applications of bioactive glass in maxillofacial and dental uses, and provides clinical findings. Key words: bioactive glass, clinical, maxillofacial, dental.

10.1

Introduction

This chapter reviews the use of bioactive glass both in maxillofacial and head and neck surgery, in skull bone reconstruction, and in dentistry, starting with the former. Bone reconstruction in the maxillofacial skeleton has been a surgical challenge for many decades. The maxillofacial area is a unique challenge to the surgeon because it is related to infection-sensitive structures such as the paranasal sinuses, upper respiratory tract and oral cavity. Reconstructions have been carried out with various materials including metals such as gold, silver, tantalium, stainless steel and titanium (Sanan and Haines, 1997; Chim and Schantz, 2005). The use of xenografts has been reviewed by Chim and Schantz (2005). For more than a century, there has been research to find a more suitable material to repair or replace bony segments of the musculoskeletal system (Damien and Parsons, 1991). Clinically used bone defect reconstruction materials are compared in Table 10.1, based both on the available literature and the authors’ clinical experience.

10.2

Current materials and requirements in maxillofacial reconstruction

According to many experts, the use of autogenic bone has been the gold standard in skull bone reconstruction. Cancellous bone grafts lead to more complete repair of bone defects than those using cortical bone. Autogenic bone has greater osteogenic capacity than alternatives such as allograft or xenograft (Damien and Parsons, 1991). Fat, muscle and bank bone have also been used in tissue augmentations and sinus obliterations in the head and neck area (Billings and May, 1989; Mann et al., 1989; Weber et al., 1999; Tessier et al., 2005). The use of bone, 217 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X


+

−

+

−

No additional costs

No prolongation of operation time

Possible to shape and mould during operation

Antibacterial properties

−

+

+

−

+

+

+/−

+/−

+

+

+

−

+

+

−

+/−

+

+/−

+/−

+

+

+

Materials of animal origin: Coral, collagen matrix, gelatin, bovine bone derivative

−

+/−

+

−

+/−

+/−

+

+/−

+

+

+

Acrylate, hydroxyapatite, ionomer cement, polyethylene, Proplast

Synthetic:

+

−

+

−

+/−

+

+

+

+

+

+

Bioactive glass

−

+

+

−

−

+/−

+

+/−

+

+

+

Plaster of Paris

Notes: + : Material has this property, − : Material does not have this property, +/− : Material has limitations with regard to this property.

+

No transmission of infection

+

+

No foreign-body reaction

No influence on follow-up

+

No donor-site morbidity

No toxicity

+

−

Easy to handle

+/−

Bank tissues: bone, lyophilized cartilage

Cancellous and cortical bone, fat

Available any time and amount

Properties

Allogenic:

Autogenic:

Bone defect reconstruction material

Table 10.1 A comparison of skull and maxillofacial bone defect reconstruction materials

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X


219

fat and muscle transplants requires that enough tissue is available in the patient’s donor site. However, the amount of such bone or other autogenic material is limited. Moreover, autogenic tissue transplantion will prolong the time of an operation because the surgeon has to perform another operation to obtain the transplant, increasing the risk of donor site morbidity (Ahlmann et al., 2002). Allogenic bone is a bone obtained from one individual and transferred to a genetically different individual, while xenograft is a transfer between different species. Allogenic bone demonstrates a lower osteogenic capacity than autogenic bone with new bone occurring at a slower rate (Mellonig et al., 1981; Oklund et al., 1986). Allograft also exhibits a higher resorption rate, a generally larger immunogenic response and less revascularization of the graft (Damien and Parsons, 1991). Whilst using these materials has the advantage of delivering a single stage reconstruction without associated donor site morbidity, it also carries the risk of biohazard transfer (Aho et al., 1998; Sailer et al., 1998). An ideal material for successful reconstruction of bone should promote bone repair as effectively as cancellous autogenic bone grafts. It should be available at any time and in any amount. It should be easy to handle. There should be no donor site morbidity. It should not cause foreign-body reactions on grafting or be toxic. It must not be a vehicle for the transmission of infectious diseases. The material should ideally be compatible with follow-up investigative techniques such as computer tomography and magnetic resonance imaging. The material should also be cost-effective, i.e. it should entail no additional cost nor prolong the operation time (Damien and Parsons, 1991; Weber et al., 1999; Aho et al., 1998).

10.3

Properties of bioactive glass

A bioactive material is one that elicits a specific biological response at the interface of the material, resulting in the formation of a bond between tissues and the material (Hench and Andersson, 1993). Bioactive glass (BAG) and ceramics are synthetic materials based on a SiO2-Na2O-CaO-P2O5-Al3O2-MgO-K2O structure. They are available as small granules or larger blocks, and have been shown to be biocompatible and non-toxic (Wilson et al., 1981; Gross and Strunz, 1985). The main chemical reactions on the surface of bioactive glass are dissolution, leaching and precipitation of ions. In studies comparing synthetic bioactive materials, BAG has been shown to produce more new bone over the same period than materials such as hydroxyapatite and tricalcium phosphate (Peltola et al., 2003; Cancian et al., 1999; Cancian et al., 2004). Furthermore, in studies using Fourier transform infra red spectroscopy (FTIR), the bone produced by BAG was shown to be closer to natural bone than the bone produced by hydroxyapatite (Peltola et al., 2003). It has also been shown that BAG has antimicrobial properties, an important characteristic for a successful clinical outcome (Hench and Andersson, 1993; Stoor and Grenman, 2004; Zhang et al., 2006; Zhang et al., 2007; Peltola et al., 2006; Munukka et al., 2008; Leppäranta et al., 2008).


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

220

Bioactive glasses

10.4

Clinical applications of bioactive glass in maxillofacial reconstruction

A special bioactive glass (BAG) S53P4 with the composition SiO2 53.0; CaO 20.0; Na2O 23.0; P2O5 4.0 wt% has been used in clinical frontal sinus obliteration and frontal bone reconstruction (Peltola et al., 2006) (Fig. 10.1, 10.2, 10.3), orbital floor reconstruction (Aitasalo et al., 2001) (Fig. 10.4), nasal septum perforation corrections (Stoor and Grenman, 2004) (Fig. 10.5) and canal wall down mastoidectomy (Della and Lee, 2006) (Fig. 10.6). The clinical experience of BAG is the longest in frontal sinus obliteration, orbital wall reconstructions and nasal septum reconstructions lasting up to 12 years (Peltola et al., 2006; Aitasalo et al., 2001; Stoor and Grenman, 2004). Amongst other applications, BAG has used in sinus lifts related to dental implantology (Turunen et al., 2004; Cordioli et al., 2001). As well as sinus lifts, there have been applications in other dental fields. In tympanomastoidectomy BAG and conchal cartilage showed an equivalent clinical outcome in reconstructing the posterior canal wall (Abramovich et al., 2008). Bioactive glass-ceramic middle ear implants in ossicular chain reconstructions also showed good tolerance after eight years (Reck et al., 1988). Long-term histotologic studies (Fig. 10.3, 10.4)

10.1 BAG granules, 0.5 to 0.8 mm in size, used in frontal sinus obliteration and mastoidal cavity filling.



221

10.2 Frontal sinus obliteration with BAG on a 3 mm thick CT scan, 5 years postoperatively. Complete obliteration of sinuses is seen without loss of volume (black arrow). Scanning parameters 140 kV and 94 mAs/3.0 s.

10.3 Histologic 20 µm thick section from BAG obliteration at 8 years postoperatively. Lamellar new bone formation (B) with scattered fibrous tissue (FT) between the glass granule remnants (BAG). (Masson–Goldner stain; magnification 10).


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

222 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

10.4 Micrograph from orbit-harvested BAG plate 3 years postoperatively. On surface of glass implant, green reaction layers and slight resorption of BAG (black arrow) are seen.

10.5 Photograph of BAG plate harvested in revision surgery from nasal septum perforation reconstruction at 5 years after primary surgery. Slight resorption of original margins of the plate is seen (white arrows).

10.6 Large defect in right fronto temporal area in skull after traffic accident. Tailor-made BAG–PMMA composite implant (white arrow) is manufactured utilizing rapid protyping technologies. Trunk of implant is perforated to enhance tissue ingrowth.



223

have shown successful new bone formation between BAG granules without harmful effects (Peltola et al., 2006; Aitasalo et al., 2001). In cranioplasty and maxillofacial reconstructions tailor-made BAG implants could be useful when large hard tissue defects need to be reconstructed, e.g. after tumor surgery. However, BAG is a brittle and rigid material and thus has limitations in shaping and bending for some clinical requirements. This limitation has led to the development of composite materials using BAG to make fuller use of its benefits. A composite material is a material including at least two component biomaterials. BAG and polymethyl metacrylate (PMMA) are widely used in surgical applications, for example in head and neck area reconstructions (Fig. 10.6) where a two-year follow-up study showed good aesthetic and functional outcomes (Peltola et al., 2009; Aitasalo et al., 2009). However, more long-term clinical research and follow up are needed to draw final conclusions about the clinical value of BAG–PMMA implants. Hybrid materials using combinations of cells and BAG are also a promising potential material for further study.

10.5

Clinical applications of bioactive glass in dentistry

The oral cavity, with its large spectrum of microbes and saliva secretion, is a demanding environment for all materials used in clinical dentistry. Apart from dental fillings, there is a need for biomaterials such as BAG in such areas of dentistry as periodontology (Mengel et al., 2006), implantology and prosthetics (Gatti et al., 2006), sinus lift (Tadjoedin et al., 2002) and hypersensitive dentin treatment (Lee et al., 2005). As noted earlier, BAG has both osteoconductive and antimicrobial properties that can be very useful in dental applications (Hench and Andersson, 1993; Wilson et al., 1981; Gross and Strunz, 1985; Zhang et al., 2006; Zhang et al., 2007). BAG mineralization effects are also promising properties in both restorative dentistry and in treatment of dentin hypersensitivity but need more clinical research. Research in clinical applications of BAG in such areas as dentistry has shown some promising results (Mengel et al., 2006; Gatti et al., 2006; Tadjoedin et al., 2002; Lee et al., 2005). BAG has been used in the treatment of intrabony defects in patients with generalized periodontitis (Mengel et al., 2006; Sculean et al., 2005), in dental extraction sites before dental implant placement (Gatti et al., 2006; Trisi et al., 2006) and after third molar extractions (Thorndson and Sexton, 2002). In such studies BAG has been seen to promote new bone formation in the jaw bone and the development of connective tissue in the periodontal area. The anti-gingivitis and anti-plaque effects of BAG powder have been studied in a placebo controlled study (Tai et al., 2006) which showed a reduction of gingival bleeding and oral plaque formation. The use of BAG has been investigated in the treatment of dentin hypersensitivity (Lee et al., 2005) and in tooth mineralization (Yli-Urpo et al., 2005), though the promising results in these studies need to be backed up by clinical results. As in other areas, BAG could in future be used as


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

224 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

part of a composite material, e.g. in tissue guiding membranes or scaffolds, and in endodontics where both osteoproductive and antimicrobial components are needed. Hybrid materials combining cells and BAG may offer promising solutions for demanding treatments, e.g. in implantology, periodontology and alveolar crest augmentation. However, more studies providing the results of long-term clinical performance, together with longer-term histological and tissue studies, are needed to demonstrate the reliability of these biomaterials for dental applications.

10.6

References

Abramovich S., Hannan S. A., Huins C. T., Georgalas C., McGuinness J., Vats A., Thompson I., ‘Prospective cohort comparison of bioactive glass implants and conchal cartilage in reconstruction of the posterior canal wall during tympanomastoidectomy’, Clin Otolaryngol 2008; 33:553–559. Ahlmann E., Patzakis M., Roidis N., Shepherd L., Holtom P., ‘Comparison of anterior and posterior iliac crest bone grafts in terms of harvest-site morbidity and functional outcomes’, J Bone Joint Surg Am 2002; 5:716–720. Aho A. J., Hirn M., Aro H. T., Heikkilä J. T., Meurman O., ‘Bone bank service in Finland. Experience of bacteriologic, serologic and clinical results of the Turku bone bank 1972–1995’, Acta Orthop Scand 1998: 69:559–565. Aitasalo K., Kinnunen I., Palmgren J., Varpula M., ‘Repair of orbital floor fractures with bioactive glass implants’, Journal Oral Maxillofac Surg 2001; 12:1390–1395. Aitasalo K., Peltola M., Vuorinen V., Vallittu P. ‘Novel composite implants in craniofacial reconstruction. Oral Presentation and abstract’, 9th European Skull Base Society Congress, Rotterdam, Netherlands, 15–18 April 2009. Billings E., May J., ‘Historical review and present status of freegraft autotransplantation in plastic and reconstructive surgery’, Plast Reconstr Surg 1989; 83:368–381. Cancian D. C., Hochuli-Vieira E., Marcantonio R. A., Marcantonio E. Jr, ‘Use of Biogran and Calcitite in bone defects: histologic studies in monkeys (Cebus apella)’, Int J Oral Maxillofac Implants 1999; 14:859–864. Cancian D. C., Hochuli-Vieira E., Marcantonio R. A., Garcia Júnior I. R., ‘Utilization of autogenous bone, bioactive glasses and calcium phosphate cement in surgical mandibular bone defects in Cebus apella monkeys’, Int J Oral Maxillofac Implants 2004; 1:73–79. Chim H., Schantz J. T., ‘New frontiers in calvarial reconstruction: integrating computerassisted design and tissue engineering in cranioplasty’, Plast Reconstr Surg 2005; 116:1726–1741. Cordioli G., Mazzocco C., Schepers E., Brugnolo E., Majzoub Z., ‘Maxillary sinus floor augmentation using bioactive glass granules and autogenous bone with simultaneous implant placement. Clinical and histological findings’, Clin Oral Implants Res 2001; 12:270–278. Damien C. J., Parsons J. R., ‘Bone graft and bone graft substitutes: a review of current technology and applications’, J Appl Biomater 1991: 2:187–208. Della Santina C. C., Lee S. C., ‘Ceravital reconstruction of canal wall down mastoidectomy: long-term results’, Arch Otolaryngol Head Neck Surg 2006; 6:617–623. Gatti A. M., Simonetti L. A., Monari E., Guidi S., Greenspan D., ‘Bone augmentation with bioactive glass in three cases of dental implant placement’, J Biomater Appl 2006; 4:325–329.



225

Gross U., Strunz V., ‘The interface of various glasses and glass-ceramics with bony implantation bed,’ J Biomed Mater Res 1985; 19:251–271. Hench L. L., Andersson Ö. H., ‘Bioactive glasses’, Bioceramics 1993; 1:41–62. Lee B. S., Tsai H. Y., Tsai Y. L., Lan W. H., Lin C. P., ‘In vitro study of DP-bioglass paste for treatment of dentin hypersensitivity’, Dent Mater J 2005; 4:562–569. Leppäranta O., Vaahtio M., Peltola T., Zhang D., Hupa L., Hupa M., Ylänen H., Salonen J. I., Viljanen M. K., Eerola E. Antibacterial effect of bioactive glasses on clinically important anaerobic bacteria in vitro’, J Mater Sci Mater Med 2008; 19:547–551. Mann W., Riechelmann H., Gilsbach J., ‘The state of the frontal sinus after craniotomy’, Acta Neurochir 1989; 100:101–103. Mellonig J. T., Bowers G. M., Bailey R. C., ‘Comparison of bone graft materials’, J Periodontol, 1981; 52:291–302. Mengel R., Schreiber D., Flores-de-Jacoby L., ‘Bioresorbable membrane and bioactive glass in the treatment of intrabony defects in patients with generalized aggressive periodontitis: results of a 5-year clinical and radiological study’, J Periododontol 2006; 10:1781–1787. Munukka E., Leppäranta O., Korkeamäki M., Vaahtio M., Peltola T., Zhang D., Hupa L., Ylänen H., Salonen J. I., Viljanen M. K., Eerola E., ‘Bactericidal effects of bioactive glasses on clinically important aerobic bacteria’, J Mater Sci Mater Med 2008; 19:27–32. Oklund S. A., Prolo D. J., Gutierrez R. V., King S. E., ‘Quantitative comparisons of healing in cranial fresh autografts and processed autografts, and allografts in canine skull defects’, Clin Orthop 1986: 205:269–291. Peltola M. J., Aitasalo K. M., Suonpää J. T., Yli-Urpo A., Laippala P. J., Forsback A. P., ‘Frontal sinus and skull bone defect obliteration with three synthetic bioactive materials. A comparative study’, J Biomed Mater Res (Appl Biomater) 2003; 1:364–372. Peltola M., Aitasalo K., Suopää J., Varpula M., Yli-Urpo A., ‘Bioactive glass S53P4 in frontal sinus obliteration: A long-term clinical experience’, Head Neck 2006; 9:834–841. Peltola M., Aitasalo K., Tirri T., Rekola J. ‘Biomateriaalit kallonalueen luupuutosten hoidossa. Suomen Lääkärilehti 2009; 9:815–820. In Finnish. (Biomaterials in Skull Bone Reconstructions. A review article’, Finnish Medical Journal 2009; 9:815–820. English summary). Reck R., Störkel S., Meyer A., ‘Bioactive glass-ceramics in middle ear surgery. An 8-year review’, Ann NY Acad Sci 1988; 523:100–106. Sailer H. F., Grätz W., Kalavrezos N. D., ‘Frontal sinus fractures: principles of treatment and long-term results after sinus obliteration with the use of lyophilized cartilage’, J Craniomaxillofac Surg 1998; 26:235–242. Sanan J., Haines S. J., ‘Repairing holes in the head: a history of cranioplasty’, Neurosurgery 1997; 40:588–603. Sculean A., Windisch P., Keglevich T., Gera I., ‘Clinical and histologic evaluation of an enamel matrix protein derivative combined with a bioactive glass for the treatment of intrabony periodontal defects in humans’, Int J Periodontics Restorative Dent 2005; 2:139–147. Stoor P., Grenman R., ‘Bioactive glass and turbinate flaps in the repair of nasal septum perforations’, Ann Otol Rhinol Laryngol 2004; 8:655–661. Tadjoedin E. S., de Lange G. L., Lyaruu D. M., Kuiper L., Burger E. H., ‘High concentrations of bioactive glass material (BioGran) vs. autogenous bone sinus floor elevation’, Clin Oral Implants Res 2002; 4:428–436.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

226 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Tai B. J., Bian Z., Jiang H., Greenspan D. C., Zhong J., Clark A. E., Du M. Q., ‘Antigingivitis effect of a dentifrice containing bioactive glass (NovaMin) particulate’, J Clin Periodontol 2006; 2:86–91. Tessier P., Kawamoto H., Posnick J., Raulo Y., Tulasne J. F., Wolfe S. A., ‘Complications of harvesting autogenous bone grafts: a group experience of 20 000 cases’, Plast Reconstr Surg 2005; 116:725–735. Thorndson R. R., Sexton S. B., ‘Grafting mandibular third molar extraction sites: a comparison of bioactive glass to nongrafted site’, Oral Surg Oral Med Oral Pathol Oral Radiol Endod 2002; 4:413–419. Trisi P., Rebaudi A., Calvari F., Lazzara R. J., ‘Sinus graft with biogran, autogenous bone, and PRP: a report of three cases with histology and micro-CT’, Int J Periodontics Restorative Dent 2006; 2:113–125. Turunen T., Peltola J., Yli-Urpo A., Happonen R. P., ‘Bioactive glass granules as a bone adjunctive material in maxillary sinus floor augmentation’, Clin Oral Implants Res 2004; 15:135–141. Weber R., Draf W., Kahle G., Kind M., ‘Obliteration of the frontal sinus – state of the art and reflections on new materials’, Rhinology 1999; 37:1–15. Wilson J., Pigott G. H., Schoen F. J., Hench L. L., ‘Toxicology and biocompatibility of bioglasses’, J Biomed Mater Res 1981; 15:805–817. Yli-Urpo H., Närhi M., Närhi T., ‘Compound changes and tooth mineralization effects of glass ionomer cements containing bioactive glass (S53P4), an in vivo study’, Biomaterials 2005; 30:5934–5941. Zhang D., Munukka E., Leppäranta O., Hupa L., Ylänen H. O., Salonen J. I., Eerola E., Viljanen M. K., Hupa M., ‘Comparison of antibacterial effect of three bioactive glasses’, Key Engineering Materials 2006; 309:345–348. Zhang D., Munukka E., Hupa L., Ylänen H. O., Viljanen M. K., Hupa M., ‘Factors controlling antibacterial properties of bioactive glass, Key Engineering Materials 2007; 330:173–176.


11 Bioactive glass and biodegradable polymer composites T. NIEMELÄ and M. KELLOMÄKI , Tampere University of Technology, Finland

Abstract: Bioactive ceramic materials have osteoconductive potential and excellent bone-bonding character, but they do not fulfill the required mechanical properties to replace bone tissue. Polymeric biomaterials have in many cases more applicable mechanical properties, but most of these polymers do not have properties that facilitate bone tissue healing. Both bioactivity and bone-bonding ability would be beneficial for bone tissue implants and thus the bioceramics have been combined with polymeric materials. In this chapter the composites of bioactive glasses and biodegradable polymers are reviewed. The chapter focuses on dense and load-bearing composites and reviews some recently reported studies of bioactive glass biodegradable polymer composites. Key words: load-bearing composites, bioactive glass biodegradable polymer composites, bioactive glass particles, bioactive glass fibers.

11.1

Introduction

Composites are developed by combining two or more individual materials on a scale larger than the atomic. The one advantage of composites is that the mechanical, biological and physiological properties can be tailored to the requirements of the applications better than with the individual homogeneous materials. Usually the composite consists of the continuous phase, called ‘matrix’, and dispersed phase, which can be fillers or fibers, for example. Most often the dispersed phase is stiffer than the matrix material and thus it is considered as a reinforcement component. However, the composite structure can also give some additional functionality to the biomedical composites, such as bioactivity, controlled drug release and the desired biodegradation profile. Bioactive glasses have osteoconductive potential and excellent bone bonding characteristics but at the same time they are ceramic in their nature. This means that they usually possess high compression strength but are also very brittle. These properties as such are not suitable for medical load-bearing applications and thus many research groups have started to study the possibility of combining the bioactive glasses together with a biodegradable polymer. In that way biodegradation and osteoconductivity can be combined and the properties of the composite material can be tailored to the demands of the applications. In this chapter the composites of bioactive glasses and biodegradable polymers are discussed. The chapter focuses on the dense and load-bearing composites 227 © Woodhead Publishing Limited, 2011

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

228 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

of different bioactive glasses and biodegradable polymers. The nature of the biodegradable polymers is defined and some of the commonly used manufacturing methods of the composites are briefly introduced. Bioactive glass can take different forms in composites, for example as particles, fibers or coatings. All these forms are discussed in more detail and several examples of the recently studied dense bioactive glass and biodegradable polymer composites are reviewed.

11.2

Biodegradable polymers

Biodegradable polymers are one group of polymeric materials. The molecular chains of the polymers can be broken down either through hydrolytic degradation or by enzymatic means. Interest in the use of the biodegradable polymers in biomedical applications has increased and current trends show that in the future the biodegradable polymers may replace the use of biostable biomaterials, step by step. The reasons why biodegradable polymers are preferred to biostable biomaterials in implant applications could be condensed into two major advantages of biodegradable polymers. First, once they have accomplished their function in the body they disappear and do not leave any marks of the residuals in the implantation site. Due to the total disappearance of the material there is no need for revision surgery, which is naturally more convenient for the patients. Secondly, because implants made from biodegradable polymers degrade gradually the stresses acted in the implantation site are transferred gradually to the healing tissue. Thus the stress shielding and weakening of the fixed tissue are prevented and a suitable remodeling rate of the healing tissue is enabled. Biodegradable polymers can degrade either by hydrolysis (without the enzyme catalysis) or by enzymatic mechanism. Hydrolysis is the main degradation mechanism of the biodegradable polymers, but depending on the polymer structure, they can also undergo at least partial enzymatic degradation. Hydrolytic degradation means the breakdown of the hydrolytically unstable polymer backbone in the presence of water. The water molecule penetrates the bulk of the polymer and randomly cleaves the chemical bonds. This occurs first in the amorphous region of the polymer. The breakdown converts long polymer chains into shorter ones and causes a decrease in the molecular weight of the polymer. At the same time the crystalline regions still keep the structure together and the physical and mechanical properties of the polymer remain unchanged. When the degradation proceeds further, the breakdown of the chemical bonds occurs also in the crystalline regions causing a reduction in the physical and mechanical properties of the polymer. At the last stage of the hydrolysis the fragments are metabolized by the enzymes in vivo. In enzymatic degradation, specific proteins, called enzymes, can cause the degradation of the polymer. The enzymatic degradation is a very complex mechanism and its occurrence is strongly influenced by the polymer composition. Most often,


Bioactive glass and biodegradable polymer composites

229

biodegradable polymers of natural origin (i.e. natural polymers) undergo the enzymatic degradation. Biodegradable polymers can be classified in several ways. One is the division into the groups of natural and synthetic polymers. Both can be used in biomedical applications, but the synthetic polymers have several advantages over the natural polymers as implant materials (Middleton and Tipton, 2000). Natural polymers are derived from renewable resources, which mean for example plants, animals and microorganisms, and thus they usually have an excellent biocompatibility and are naturally biodegradable. However, they are typically mechanically very weak and due to their rather complex chemical structures and thermal sensitivity, processing without solvents is difficult. The most common groups of biodegradable natural polymers are polysaccharides and proteins. In addition to the renewable resources biodegradable polymers can be manufactured synthetically from petrochemical resources. These are called synthetic biodegradable polymers and their properties can be modified in many ways, such as blending and copolymerization. The most important and widely studied biodegradable synthetic polymers are aliphatic polyesters, such as polylactides, polyglycolide and their copolymers.

11.3

Manufacturing of the composites

There are several methods of fabricating biodegradable polymer matrix composites. Most methods are applicable only to a specific kind of composite, and some are still at the development stage. Some methods are limited only to particulate reinforcements, whereas others are better suited to the handling of continuous fiber reinforcements. For some methods, the components themselves are used, whereas other methods use the preforms. Brief descriptions of some of the important manufacturing methods are given in the following sections and all the basic methods have variants named specifically.

11.3.1 Melt extrusion Melt extrusion is a widely used manufacturing method for producing continuous products having constant cross-sections (rods, sheets, pipes, fibers, etc.). It is a method suitable for thermoplastic polymers, which can be reshaped when heated. In the composite science, melt extrusion can also be used to mix and compound the matrix polymer and reinforcing elements to form pellets, which can be further used in other processes. The extruder consists mainly of the rotating screw in a heated barrel. At the beginning of the barrel is the feeding hopper into which the raw materials are fed. After feeding, the raw materials come into contact with the rotating screw, which is responsible for the moving and homogenizing of the polymer. Heating elements, placed over the barrel, soften and melt the polymer gradually as it is conveyed forward in the barrel. At the end of the barrel is the


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

230 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

heated die that has an orifice with the specific profile needed for the extrudate. The melted polymer paste is forced to run through the die and after that cooled to the final shape.

11.3.2 Self-reinforcing Self-reinforcing (SR) is an important method of manufacturing high-strength structures and it has been studied as a method of increasing the strength of biodegradable devices since the mid 1980s. In self-reinforcing, the polymer matrix is reinforced with oriented polymer fibers or fibrils of a similar chemical composition. This provides an excellent adhesion between the matrix and fibers resulting in high-strength devices. Studies show that the most effective way to create the biodegradable polymer from the self-reinforced structure is the mechanical deformation of non-reinforced material and especially the die drawing process. In this method the material is drawn through the heated die at a controlled temperature above the polymer’s glass transition temperature causing the orientation of the polymer chains (e.g. Törmälä, 1992).

11.3.3 Compression molding Compression molding is the process in which the material is pressed into a mold taking the shape of the mold cavity and becoming cured owing to heat and pressure applied to the material. It is suitable for both thermosets and thermoplastic polymers. The raw materials can be in the form of powder, granules or preforms. The technique is widely used to make flat laminates and simple shapes from fabric preforms, such as woven or random-oriented fiber mats.

11.3.4 Solvent casting In solvent casting, the matrix polymer is dissolved in a volatile solvent. Thus the solubility of the polymer is the most important prerequisite for the solvent casting technique. Other requirements are, for example, the formation of a stable solution with a reasonable minimum solid content and viscosity, and the possibility of removing the casting support. In the biomedical field the biocompatibility of the solvent and its residues is also important. After dissolving, the reinforcements, for example particles or short fibers, can be added and mixed into the solution. The final solution is cast to the mold, which can take for example the form of a plate (to form films) or of a three-dimensional structure (to form 3D scaffolds). The evaporation of the solvent creates the final structure of the product. Solvent casting can also be used to form porous structures. In that case soluble particles, for example salt, are added to the solution and finally dissolved out, leaving behind the porous structure. Selection of the porogens is dependent on solubility; porogens may not be soluble in the same solvents used to dissolve the polymer.



11.4

231

Bioactive glass particle composites

The idea of combining the biodegradable polymer with the bioactive ceramic is not new. The need of this kind of load-bearing structure started with the use of bone substitutes, which have similar physiological and mechanical characteristics to living bone. Bioactive ceramic materials have osteoconductive potential and excellent bone-bonding character, but they do not possess the required mechanical properties for replacing bone tissue. Polymeric biomaterials have more applicable mechanical properties, but most of these polymers do not have the right properties for facilitating bone tissue healing. Bioactivity and bone-bonding ability would be beneficial for bone tissue implants and thus the bioceramics were combined with polymeric materials. Hydroxyapatite (HA)-reinforced polyethylene (PE) composite (HAPEX™) was originally developed as a biomaterial for bone replacement on the basis of producing suitable mechanical compatibility. The HA was demonstrated to stiffen the polyethylene, and the polyethylene to toughen the composite. Additionally, as bone mineral resembles HA, natural bone will grow onto HA (Bonfield, 1993). HA/PE composite is biostable, and studies showed that its mechanical properties remained constant in physiological solution. Furthermore, HA/PE composite provides a favorable environment for human osteoblast-like cell attachment (Huang et al., 1997a). The results with HA/PE composite have given growth to the research and development on other bioactive composites using the same rationale. Different filler and matrix materials have been studied. In order to establish a stronger bond between the implant and the bone tissue, HA could be replaced by more bioactive bioceramic, such as bioactive glass. When comparing Bioglass®/ PE composite to the HA/PE composite in vitro, the faster formation of the bonelike apatite on the composite surface was noticed. This indicated higher bioactivity. However, mechanical properties of the Bioglass®/PE composite decreased during immersion in an aqueous environment (Huang et al., 1997b; Wang, 2003). If the higher level of mechanical properties is needed, polysulfone (PSU) may be a better choice for the matrix material than PE. Therefore, the HA/PSU composite has been developed as a hard tissue replacement material (Wang et al., 2001). Mechanical properties of HA/PE composite have also been increased using hydrostatic extrusion. During this method the polymer chains are aligned to a certain orientation, which leads to an increase in stiffness and strength along the direction of orientation. The resultant mechanical properties of the highly filled HA/PE composite were shown to be within the bounds for cortical bone and thus show promise also for major load-bearing applications (Ladizesky et al., 1997a). Another technique to improve the mechanical properties of HA/PE composite involves reinforcing the polymer matrix with high-performance polyethylene fibers (Ladizesky et al., 1997b, 1998). The reinforcement of the matrix polymer has also performed for biodegradable polymer matrices. The most studied reinforcing method is self-reinforcing


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

232 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

(Törmälä, 1992), which is already widely used for example to manufacture commercially available bioabsorbable orthopedic implants. Self-reinforcing creates an oriented, high-strength structure with reinforcing fibrous elements, which have the same chemical composition as the matrix polymer. Due to the excellent adhesion between the matrix polymer and fibers the mechanical properties are increased. Self-reinforcing has also performed successfully for bioactive glass particle/bioabsorbable polymer composites manufactured using single- or twin-screw extrusion (Fig. 11.1) (Kellomäki et al., 2000; Niiranen and Törmälä, 1999a, 1999b; Niiranen et al., 2001, 2004; Niemelä et al., 2005a, 2007, 2008). Self-reinforcing is reported to improve the initial mechanical properties and eliminate the brittle fracture behavior of the composite. Self-reinforcing also modified the composite structure. Bioactive glass filler particles initiated both interior and exterior pores, which are mainly not interconnected. At the composite surface, the initially polymer-covered bioactive glass particles were exposed as a result of self-reinforcing. This allows bioactive glass direct contact with the

11.1 Extruded composite rods containing poly-L-lactide-co-glycolide 80/20 as a matrix material and 40 wt% of bioactive glass 13–93 spheres as a filler material. The transparent rods on the left are self-reinforced from plain matrix polymer. The addition of bioactive glass rendered the rods opaque. The thicker rods on the right are non-reinforced rods. Self-reinforcing reduced the diameter of the rods and at the same time increased their mechanical properties and modified their composite structure. Scanning electron microscopy images presents the external porosity (a) and internal structure (b) of the composites formed during self-reinforcing. Previously unpublished data, Tampere University of Technology (TUT)/Department of Biomedical Engineering (BME).



233

surrounding environment and to enables it to react faster and thus there is no need for further mechanical processing to expose the bioactive glass (Niiranen and Törmälä, 1999b; Niemelä et al., 2005a). A formed porous surface is also thought to be beneficial for mechanical interlocking of bone tissue. Though die-drawing is usually used to prepare self-reinforced structures, they can also be prepared by very controlled compression molding of fibers. In compression molding, a temperature is chosen that will melt only the surface of the fibers allowing the matrix polymer to form and attach the fibers together. The processing parameters (temperature, pressure and time) need to be chosen and controlled carefully to retain the reinforcing effect of the polymer fibers (Hine et al., 1993). This has also been successfully applied in preparing bioceramic and bioactive glass containing composites (Bleach et al., 2001, 2002; Ellä et al., 2005; Huttunen et al., 2006; Kellomäki et al., 1997). Niemelä et al. (2005a) have reported the effects of the different bioactive glass filler contents on the initial mechanical properties and bioactivity of the selfreinforced bioactive glass poly-L/DL-lactide 70/30 composites. The composites were observed to become weaker and more brittle when large quantities of bioactive glass were added. The more bioactive glass was added, the more the mechanical properties decreased. However, the mechanical properties of the composites could be improved by self-reinforcing. Self-reinforcing also made the initially brittle composites ductile. It was noted that the mechanical properties were sufficient for small bone fracture fixations if the bioactive glass content was 20 to 30 wt%. Bioactive glass content also influences the bioactivity of the composites. The composites turn out to be bioactive if the bioactive glass content is 20 to 40 wt%. This involves open pores around the bioactive glass particles, because it has been shown that an open bioactive glass surface is required for the rapid surface reactions of this kind of composite. Composite surfaces containing more than 40 wt% of bioactive glass were not porous, due to difficulties in selfreinforcing, and thus they seemed not to be bioactive without additional machining of the surface. As a conclusion, Niemelä et al. (2005a) discovered self-reinforced bioactive glass containing polylactide composites to be a potential implant material for small bone fracture fixations. The addition of bioactive glass was seen in studies to affect the in vitro behavior of the self-reinforced bioactive glass poly-L/DL-lactide 70/30 composites (Niiranen et al., 2004; Niemelä et al., 2008). Although the initial mechanical properties of the self-reinforced bioactive glass-containing composites were lower than in the self-reinforced plain matrix polymer, the in vitro degradation rate was slower in terms of mechanical properties, mass loss and molecular weight loss of the samples and pH of the buffer solution. The slower degradation rate was thought to be due to the porous structure formed during the selfreinforcing process. The pores enabled the acidic degradation products of the polylactide matrix to diffuse out of the internal structure. This together with the neutralizing effect of the buffer solution due to the dissolution of the bioactive


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

234 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

glass diminished the autocatalytic degradation and thus slowed the overall degradation in vitro. The same effect has also been noticed for similar self-reinforced composites with other bioceramic filler particles, such as β-tricalcium phosphate (Niemelä et al., 2004; Niemelä, 2005b). Bioactive glass poly-L/DL-lactide 70/30 composites were also studied in vivo in the dorsal subcutaneous tissue of rats (Pyhältö et al., 2004). The results showed that the bioactive glass-containing composites were suitable for fixation of cancellous bone osteotomies in rats as long as the fixation technique was correct. The composite was also observed to elicit an osteostimulatory response at the tissue– implant interface after implantation. It was noticed that the degradation of the composites was more pronounced in vivo than in vitro (Niiranen et al., 2004). Some bioabsorbable polymers as such have shown osteoconductivity. One is Polyactive®, which is a block copolymer of polyethylene oxide (PEO) and polybutylene terephtalate (PBT). The bone-bonding ability has been noticed in certain PEO/PBT ratios (Radder et al., 1995). Kellomäki (2000) has investigated that the bone-bonding ability of Polyactive® could be improved by adding bioactive glass particles. The composite rods containing 0 to 23 wt% of bioactive glass and manufactured by extrusion turned out to exhibit excellent bioactivity. The rapid formation of apatite precipitation was noticed to spread over the whole matrix. The precipitation formation was detected even to the surface of the polymer close to the bioactive glass (‘halo-effect’). Due to the hydrogel feature the composites swelled strongly in vitro. This exposed more bioactive glass particles to direct contact with the surrounding fluids and thus further accelerated the surface reactions of the bioactive glass. This was noticed to be especially advantageous in the case of samples, like films, in which the bioactive glass particles are originally completely embedded in the polymer matrix (Kellomäki et al., 2000). An injectable composite material consisting of poly(ε-caprolactone-co-DLlactide) and bioactive glass has been produced for applications in orthopedics and in oral and maxillofacial surgery (Rich et al., 2002; Jaakkola et al., 2004). The samples for in vitro evaluation were manufactured by compounding and compression molding. Different particle sizes and amounts of bioactive glass were homogeneously incorporated in the matrix polymer to obtain either slower or accelerated bioactivity. The presence of bioactive glass was observed to affect the degradation rate of the composites in vitro. The reduction in molecular weight was more rapid the greater the amount of bioactive glass, and the smaller its particle size range. The presence of bioactive glass also affected the formation of the biologically active Ca-P layer. The higher the bioactive glass content and the surface/volume ratio was, the faster was the Ca-P formation in vitro. According to Rich et al. (2002) these materials have potential as implant materials in orthopedics and dentistry. The developed composite material was further studied by injecting as viscous liquid or moldable paste into the cancellous and cartilaginous subchondral bone defects of rabbits. The composites turned out osteoconductive



235

and easy to handle with a short setting time (Aho et al., 2004). Similar injectable composites consisting of calcium phosphate instead of bioactive glass have also been studied (Ekholm et al., 2003, 2006). Zhou et al. have manufactured bioactive glass poly-L-lactide composite membranes using a solvent evaporation technique (Zhou et al., 2007, 2009). The sol-gel-derived bioactive glass was homogeneously distributed in the composite achieving a 10 wt% bioactive glass content. The effect of the bioactive glass addition on the in vitro degradation of the composites was studied. The results showed that the bioactive glass addition reduced the overall degradation rate in terms of mechanical properties, mass loss and molecular weight loss. The reasons for that kind of behavior were thought to be the neutralizing effect caused by dissolution of bioactive glass and the interfaces between the matrix and filler, which facilitated the diffusion of the degradation products. The observations were similar to the self-reinforced polylactide composites containing 20 wt% of bioactive glass reported by Niemelä et al. (2008). In the solvent-evaporated composite the matrix polymer was detected to cover the bioactive glass particles on the composite surface and thus the bioactive glass has no immediate contact with the surroundings when immersed in the buffer solution. In spite of this, the rod-like hydroxyapatite crystals deposited on the surface after three days in vitro, and after 14 days the hydroxyapatite layer was formed (Zhou et al., 2007). The size of the filler particles influences the mechanical properties of the composites. The larger surface area of the bioactive glass leads to increased interface effects and this also contributes to improved bioactivity. Misra et al. (2008) have studied the differences in the structural, thermal, mechanical and biological behavior of poly(3-hydroxybutyrate)/bioactive glass composites resulting from the addition of nanoscale or microscale bioactive glass particles. The composite films (thickness 0.12 to 0.14 mm) were manufactured by the solvent-casting technique. The addition of nanoscale bioactive glass added the roughness and changed the morphology of the surface. The nanoscale bioactive glass particles on the surface were exposed, which was not detected in the case of microscale bioactive glass particles. The significant reinforcing effect of the nanoscale bioactive glass particles was also reported. During the immersion in the buffer solution, the nanoparticles containing composites were observed to absorb more water, to lose more mass and have higher bioactivity compared to microparticles containing samples. The nanoscale bioactive glass addition also considerably improved the total protein absorbtion. To conclude, the results suggested that the bioactive glass nanoparticle composite system might be well suited for bone tissue engineering (Misra et al., 2008).

11.5

Bioactive glass fiber composites

In non-medical applications, glass fiber-reinforced polymer composites have been used for a long time. However, bioactive glass fiber-reinforced polymer


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

236 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

composites, especially biodegradable polymer composites, have not been widely reported in biomedical applications. Marcolongo et al. have studied bioactive glass fiber reinforced polysulfones in vitro and in vivo. The results showed the formation of a calcium phosphate precipitation on the surface of glass fibers within the composite material in vitro. Even the formation of a calcium phosphate precipitation on the surface of polymer close to the glass fibers was detected (‘halo-effect’). In vivo, the bone tissue was seen to exhibit direct contact with the glass fibers and adjacent polymer matrix. This resulted in high interfacial bond strengths compared to plain polymer controls (Marcolongo et al., 1997, 1998). Also, composites containing fibers of bioceramics other than bioactive glass have been studied. Ahmed et al. (2009) have studied the in vitro retention of the mechanical properties and cytocompatibility of phosphate-based glass fiber/ polylactide composites and Charvet et al. (2000) the mechanical and fracture behavior of calcium phosphate fiber/polycarbonate composites. Melt-spun bioactive glass fibers as such have high strengths that depend largely on the diameter of the fibers and structural flaws. They are also brittle and very sensitive to all types of contamination. Even a slight abrasion on the surface of the bioactive glass was noticed to decrease the maximum strength drastically (Pirhonen et al., 2006a, 2006b). By coating the bioactive glass with the polymer, the abrasion can be reduced and even avoided. The coating also enables the further fabrication of the continuous bioactive glass fibers, such as those used in manufacturing woven and knitted textiles (Pirhonen and Törmälä, 2006c; Tukiainen et al., 2006a, 2006b). Pirhonen and Törmälä (2006c) have reported two possible methods (dipping the fibers and pulling them through a viscous solution) for coating the continuous bioactive glass (13-93) fibers with different biomedical polymers. Dipping was preferred for the bunch of thin fibers and a coating thickness of 2 to 5 µm was achieved. Pulling fibers through a viscous solution was a better method for thicker fibers. This method was suitable even for coating single fibers and the coating thickness achieved was 10 to 30 µm. The surface reactions of the coated bioactive glass fibers were studied in vitro in SBF. The formation of the calcium phosphate layer on the surface of the coated fibers was observed to be slightly delayed compared to the non-coated fibers. The handling and mechanical properties of the bioactive glass fibers were significantly improved in consequence of the polymeric coating. Before coating, thin fibers could not even be unwound from the roll without breakage. Both methods were shown to be suitable for coating of bioactive glass fibers but the adhesion between the coating and bioactive glass had to be improved (Pirhonen and Törmälä, 2006c). Tukiainen et al. (2006a) have successfully used coated bioactive glass fibers together with bioabsorbable polymer fibers to manufacture woven and knitted textile structures. Further, these textile structures have been used to manufacture porous bioactive/bioabsorbable load-bearing composites by compression molding (Fig. 11.2). In these structures polymer fibers function as a supportive, reinforcing or matrix forming material



237

11.2 Porous bioactive/bioabsorbable load-bearing composites containing both poly-L/D-lactide 96/4 and poly-L/DL-lactide 70/30 fibers and bioactive glass 1-98 fibers coated with poly-L-lactide-co-glycolide 50/50. Composites are manufactured by compression molding with different compression parameters. Samples size 10 mm × 3 mm × 30 mm. Previously unpublished data, TUT/BME.

and bioactive glass fibers provide a reinforcing effect and give osteoconductivity to completed material. The porosity of the structures affects the mechanical properties and was controlled by compression parameters. The initial mechanical properties obtained were thought to be sufficient for various load-bearing bone applications (Tukiainen et al., 2006b). A different manufacturing method of bioactive glass/biodegradable polymer composite was reported by Jiang et al. (2005). Preparation of continuous Bioglass® fiber/poly(ε-caprolactone) composite has been achieved by using a monomer transfer molding technique coupled with surface initiated polymerization generated by amine silane surface treatment. The method provided a route to processing implants of complex shapes and it was noted that the surface initiated polymerization improved the mechanical properties of the composites owing to the chemical bond formed between the matrix polymer and the surface of the reinforcement fibers. Improving the interfacial bonding is thought to be the key to the successful use of bioactive bioabsorbable polymer composites in the medical field. The amine silane treatment on the surface of the bioactive glass fibers also improved the strength retention of the composites. Surface-treated composites were seen still to have the properties in the range of bone after 6 weeks’ degradation (Jiang et al., 2005). Efforts have also been made to combine continuous bioactive glass fibers with bioabsorbable polymer by using an extruder with a crosshead nozzle (Paatola et al., 2001). Glass fiber reinforcing increased the initial properties of the composites, but after 6 weeks’ hydrolysis the reinforcing effect was totally


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

238 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

lost owing to the lack of adhesion between the bioactive glass and polymer matrix. However, the addition of bioactive glass was noted to generate the desired bioactivity to the composites. The fragility and poor abrasion resistance of the fibers have limited their use as continuous fibers and thus bioactive glass fibers have been mainly short (Pirhonen et al., 2001; Haltia et al., 2004) and directional fibers (Jukola et al., 2008; Huttunen et al., 2008). Composites containing 0 to 40 vol% chopped bioactive glass fibers with bioabsorbable polymer have been manufactured using piston injection molding (Pirhonen et al., 2001). The flexural properties of the rod-shaped samples were noticed to improve as their fiber content increased. In composites containing 30 to 40 vol% bioactive glass fibers, the flexural properties even reached the level of bone tissue. Jukola et al. (2008) have tried to enhance the mechanical properties of starch-poly-ε-caprolactone (SPCL) by combining the unidirectional bioactive glass fiber performs with the polymer sheets using compression molding. Several different structures were manufactured (bioactive glass content approximately 10 to 20 wt%) and compared with the non-reinforced polymer. It was observed that the initial mechanical properties of the bioactive glass fiber reinforced samples were at least 50% better than properties of the non-reinforced polymer. However, after two weeks’ in vitro period the mechanical properties of the reinforced samples decreased to the same level as non-reinforced samples, and lower in strength than bone tissue. Thus the studied composites were not found to be adequate for bone fracture fixation applications as such. Further development is needed. Huttunen et al. (2008) have used both bioactive glass fibers and bioabsorbable polymer fibers as reinforcement elements in bioabsorbable polymer matrix. The reinforced composites were manufactured by filament winding followed by compression molding. The studied hybrid composites were reported to be very strong, having the initial flexural modulus in the range of cortical bone. The degradation behavior of these hybrid composites is thought to be two-stepped, owing to the two different reinforcing elements. The bioactive glass fibers, which have high stiffness, give strong protection in the beginning of bone fracture healing during initial consolidation. After some weeks the bioactive glass fibers lose their reinforcing effect due to their degradation. However, the strength and modulus of the composite do not decrease drastically because of the bioabsorbable fibers, which still reinforce the matrix polymer and secure the final healing of the bone fracture.

11.6

Coatings

The one alternative use in the biomedical composites for bioceramics, such as bioactive glass, is to apply the bioceramic as a coating. In the field of load-bearing applications one of the main applications is the coating of the metallic implants with bioactive glass. In these cases the metallic implant will be responsible for the



239

strength of the device, and the bioactive glass coating provides an interfacial attachment to the bone tissue, and also protects the metal against corrosion (Hench and Andersson, 1993). In recent years, several coating techniques, such as enameling (Lopez-Esteban et al., 2003; Tomsia et al., 2005), plasma spraying (Schrooten et al., 1999), ion beam sputtering (Wang et al., 2002), pulsed laser deposition (Liste et al., 2004), laser cladding (Comesanã et al., 2010) and sol-gel technique (Durán et al., 2004; Fathi and Doost Mohammadi, 2008), have been studied to produce the bioactive glass coatings on the metallic implants. The major limiting factor in the many techniques is the poor adhesion achieved between the bioactive glass and metal. Attention should also be addressed to the fact that the coating process should not degrade the properties of the substrate or coating material (Lopez-Esteban et al., 2003). The techniques used in coating the metallic implants with bioactive glass involve high temperatures. Therefore, these techniques are not suitable for coating the polymeric materials, owing to the lower degradation temperature of the polymers. Thus the coating of the polymeric materials with the bioactive glass is studied less than the coating of the metals. Niiranen and Törmälä (1999b) have coated the high strength malleable bioabsorbable polymer plates with bioactive glass spheres (Fig. 11.3). The bioactive glass spheres were implanted on the other side of the plate preform by pressing between the metallic plates using the different pressing parameters. The pressing parameters were observed to influence the attachment of the bioactive glass spheres on to the surface of the plate. At low pressure and temperature the attachment did not occur, and under high pressure the bioactive glass spheres were crushed. Too high a temperature caused changes in the polymer structure, which, of course, was not desired. Applying pressure using the right

11.3 Scanning electron microscopy image of a poly-L/DL-lactide 70/30 plate coated with bioactive glass 13-93 spheres. Previously unpublished data, TUT/BME.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

240 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

parameters generated an osteoconductive, bioabsorbable, high strength plate, which is also malleable at room temperature. It was suggested that the composite plates under study could be used for guided bone regeneration and as fixation devices in the bone fractures (Ruuttila et al., 2006). The coating of biodegradable polymer sutures with bioactive glass has also been studied (Stamboulis et al., 2002; Boccaccini et al., 2003; Bretcanu et al., 2004). The sutures are fibers that have high mechanical strength owing to the fibrous structure, and are useful materials for fabricating three-dimensional scaffolds for tissue engineering applications. The hypothesis regarding coating biodegradable sutures is that their bioactivity, and thus bone-bonding ability, can be enhanced by coating with bioactive glass. Also, by adding the bioactive glass with biodegradable polymer, morphological changes, such as the degradation rate, can be controlled. Stamboulis et al. (2002) have made a preliminary experimental work of coating commercially available Polyglactin 910 (Vicryl®) sutures with bioactive glass (Bioglass®) powder. The coating method performed was a simple layer-pressing procedure in which the bioactive glass particles attached mechanically to the surface of the sutures. The achieved coating was not very uniform or homogeneous, but the results were encouraging as it was shown that the mechanical performance of the suture could be altered by coating with bioactive glass. Bioactive glass was noticed to act as a protective shield affecting both the extent and the rate of the degradation of the sutures. A novel method of coating Vicryl® sutures with Bioglass® based on a slurry-dipping technique was presented by Boccaccini et al. (2003). This method has been shown to have advantages over the dry powder pressing procedure. The bioactive glass coating was noted to reduce the initial tensile properties of the sutures. However, strength retention of the sutures in SBF was improved by bioactive glass coating. The above coating methods have also been successfully tested with other bioactive glasses (Bretcanu et al., 2004). The bioactive glass coatings on the biodegradable polymer substrates in the form of surgical meshes (Stamboulis and Hench, 2001; Stamboulis et al., 2002) and highly porous foams (Roether et al., 2002); biomaterials (Gough et al., 2003) were also studied. In addition to coating the polymer substrate with bioactive glass, the bioactive glass substrates coated with biodegradable polymer have been studied (Paatola et al., 2001; Pirhonen and Törmälä, 2006c; Tukiainen et al., 2006a; Bretcanu et al., 2009a, 2009b). The coating of the high-strength bioactive glass fibers has already been discussed in the previous section (Section 11.5: ‘Bioactive glass fiber composites’). Besides coating bioactive glass fibers, bioactive glass-based scaffolds with biodegradable polymer coating have also been studied recently. Bretcanu et al. (2009a) have recently coated Bioglass®-derived flat discs with different electrospun nanofibrous biodegradable polyesters. The aim was to mimic the structure of the natural extracellular matrix by incorporating a fibrous nanotopography on the surface of the bioactive glass substrate. It was found that all the samples were highly bioactive and promoted the hydroxyapatite crystals on



241

their surfaces, which should further enhance osteoblast cell attachment and proliferation. Bretcanu et al. (2009b) have also coated the highly porous 3D Bioglass®-derived scaffold with biodegradable poly(3-hydroxybutyrate). The coating was performed by immersing the scaffold into the P(3HB)-chloroform solution for a certain time and then dried at room temperature. The coating was observed to have a positive effect on the compression strength and structural integrity of the scaffold. Also the pH of the buffer solution after the cell culture was seen to decrease. This resulted in increasing cell proliferation.

11.7

Future trends

There is a clear need for materials that have specific osteoconductive properties, a degradation rate that it is possible to tailor, and that can be processed easily into medical products. Since no material alone can fulfill the request, composites provide a real opportunity. Composites offer features suitable to both load-bearing and non-load-bearing applications. However, it is very time-consuming to test all the parameters and options available by experimentation, and therefore one of the future trends is to apply design softwares and methods (like Taguchi) for the research and development of medical devices. Another future task remains in providing one solution with all the required properties: it still is very demanding to prepare highly load-bearing material with high porosity.

11.8

References

Ahmed I., Cronin P. S., Abou Neel E. A., Parsons A. J., Knowles J. C. and Rudd C. D. (2009). ‘Retention of mechanical properties and cytocompatibility of a phosphatebased glass fiber/polylactic acid composite’, J Biomed Mater Res Part B: Appl Biomater, 89B, 18–27. Aho A. J., Tirri T., Kukkonen J., Strandberg N., Rich J., Seppälä J. and Yli-Urpo A. (2004). ‘Injectable bioactive glass/biodegradable polymer composite for bone and cartilage reconstruction: Concept and experimental outcome with thermoplastic composites of poly(ε-caprolactone-co-D,L-lactide) and bioactive glass S53P4’, J Mater Sci: Mater Med, 15, 1165–1173. Bleach N. C., Tanner K. E., Kellomäki M. and Törmälä P. (2001). ‘Effect of filler type on the mechanical properties of self-reinforced polylactide – calcium phosphate composites’, J Mater Sci: Mater Med, 12, 911–915. Bleach N. C., Nazhat S. N., Tanner K. E., Kellomäki M. and Törmälä P. (2002). ‘Effect of filler content on mechanical and dynamic properties of particulate biphasic calcium phosphate–polylactide composites’, Biomaterials, 23, 1579–1585. Boccaccini A. R., Samboulis A. G., Rashid A. and Roether J. A. (2003). ‘Composite surgical sutures with bioactive glass coating’, J Biomed Mater Res Part B: Appl Biomater, 67B, 618–626. Bonfield W. (1993). ‘Design of bioactive ceramic-polymer composite’, in Hench L. L. and Wilson J., An Introduction to Bioceramics, World Scientific Publishing, Singapore, 299–303.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

242 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Bretcanu O., Verné E., Borello L. and Boccaccini A. R. (2004). ‘Bioactivity of degradable polymer sutures coated with bioactive glass’, J Mater Sci: Mater Med, 15, 893–899. Bretcanu O., Misra S. K., Yunos D. M., Boccaccini A. R., Roy I., Kowalczyk T., Blonski S. and Kowalewski T. A. (2009a). ‘Electrospun nanofibrous biodegradable polyester coatings on Bioglass®-based glass-ceramics for tissue engineering’, Mater Chem Phys, 118, 420–426. Bretcanu O., Misra S., Roy I., Renghini C., Fiori F., Boccaccini A. R. and Salih V. (2009b). ‘In vitro biocompatibility of 45S5 Bioglass® -derived glass-ceramic scaffolds coated with poly(3-hydroxybutyrate)’, J Tissue Eng Regen Med, 3, 139–148. Charvet J. L., Cordes J. A. and Alexander H. (2000). ‘Mechanical fracture behavior of a fiber-reinforced bioabsorbable material for orthopedic applications’, J Mater Sci: Mater Med, 11, 101–109. Comesanã R., Quintero F., Lusquiños F., Pascual M. J., Boutinguiza M., Durán A. and Pou J. (2010). ‘Laser cladding of bioactive glass coatings’, Acta Biomaterialia, 6, 953–961. Durán A., Conde A., Gómez Coedo A., Dorado T., Carcía C. and Ceré S. (2004). ‘Sol-gel coatings for protection and bioactivation of metals used in orthopaedic devices’, J Mater Chem, 14, 2282–2290. Ekholm M., Hietanen J., Tulamo R.-M., Muhonen J., Lindqvist C., Kellomäki M. and Suuronen R. (2003). ‘Tissue reactions of subcutaneously implanted mixture of ε-caprolactone-lactide copolymer and tricalcium phosphate. An electron microscopic evaluation in sheep’, J Mater Sci: Mater Med, 14, 913–918. Ekholm M., Hietanen J., Tulamo R.-M., Muhonen J., Lindqvist C., Kellomäki M. and Suuronen R. (2006). ‘The copolymer of ε-caprolactone-lactide and tricalcium phosphate does not enhance bone growth in mandibular defect of sheep’, J Mater Sci: Mater Med, 17, 139–145. Ellä V., Kellomäki M. and Törmälä P. (2005). ‘In vitro properties of PLLA screws and novel bioabsorbable implant with elastic nucleus to replace intervertebral disc’, J Mater Sci: Mater Med, 16, 655–662. Fathi M. H. and Doost Mohammadi A. (2008). ‘Preparation and characterization of sol-gel bioactive glass coating for improvement of biocompatibility of human body implant’, Mater Sci Eng A, 474, 128–133. Gough J. E., Arumugam M., Blaker J. and Boccaccini A. R. (2003). ‘Bioglass® coated poly(DL-lactide) foams for tissue engineering scaffolds’, Mat.-wiss. u. Werkstofftech, 34, 654–661. Haltia A.-M., Heino H. and Kellomäki M. (2004). ‘Bioabsorbable and bioactive composite structures from SiO2-glassfibers and polylactides’, Key Eng Mater, 254–256, 549–552. Hench L. L. and Andersson Ö. (1993). ‘Bioactive glass coatings’, in Hench L. L. and Wilson J., An Introduction to Bioceramics, World Scientific Publishing, Singapore, 239–259. Hine P. J., Ward I. M., Olley R. H. and Bassett D. C. (1993). ‘The hot compaction of high modulus melt-spun polyethylene fibers’, J Mater Sci, 28, 316–324. Huang J., Di Silvio L., Wang M., Tanner K. E. and Bonfield W. (1997a). ‘In vitro mechanical and biological assessment of hydroxyapatite-reinforced polyethylene composite’, J Mater Sci: Mater Med, 8, 775–779. Huang J., Di Silvio L., Wang M., Rehman I., Ohtsuki C. and Bonfield W. (1997b). ‘Evaluation of in vitro bioactivity and biocompatibility of Bioglass®-reinforced polyethylene composite’, J Mater Sci: Mater Med, 8, 809–813.



243

Huttunen M., Ashammakhi N., Törmälä P. and Kellomäki M. (2006). ‘Fibre reinforced bioresorbable composites for spinal surgery’, Acta Biomaterialia, 2, 575–587. Huttunen M., Törmälä P., Godinho P. and Kellomäki M. (2008). ‘Fiber-reinforced bioactive and bioabsorbable hybrid composite’, Biomed Mater, 3, 1–12. Jaakkola T., Rich J., Tirri T., Närhi T., Jokinen M., Seppälä J. and Yli-Urpo A. (2004). ‘In vitro Ca-P precipitation on biodegradable thermoplastic composite of poly(εcaprolactone-co-DL-lactide) and bioactive glass (S53P4)’, Biomaterials, 25, 575–581. Jiang G., Evans M. E., Jones I. A., Rudd C. D., Scotchford C. A. and Walker G. S. (2005). ‘Preparation of poly(ε-caprolactone)/continuous bioglass fibre composite using monomer transfer moulding for bone implant’, Biomaterials, 26, 2281–2288. Jukola H., Nikkola L., Gomes M. E., Chiellini F., Tukiainen M., Kellomäki M., Chiellini E., Reis R. L. and Ashammakhi N. (2008). ‘Development of bioactive glass fiber reinforced starch-polycaprolactone composite’, J Biomed Mater Res Part B: Appl Biomater, 87B, 197–203. Kellomäki M., Törmälä P., Bonfield W. and Tanner K. E. (1997). ‘Reinforced polylactide – hydroxyapatite composites’, 13th European Conference on Biomaterials, European Society of Biomaterials, Göteborg, Sweden, 4–7 September 1997. No. 90. Kellomäki M. (2000). ‘Bioabsorbable and bioactive polymers and composites for tissue engineering applications’, Dissertation, Tampere University of Technology, Tampere, Finland. Kellomäki M., Niiranen H., Puumanen K., Ashammakhi N., Waris T. and Törmälä P. (2000). ‘Bioabsorbable scaffolds for guided bone regeneration and generation’, Biomaterials, 21, 2495–2505. Ladizesky N. H., Ward I. M. and Bonfield W. (1997a). ‘Hydrostatic extrusion of polyethylene filled with hydroxyapatite’, Polym Adv Technol, 8, 496–504. Ladizesky N. H., Ward I. M. and Bonfield W. (1997b). ‘Hydroxyapatite/high-performance polyethylene fiber composite for high-load-bearing bone replacement materials’, J Appl Poly Sci, 65, 1865–1882. Ladizesky N. H., Pirhonen E. M., Appleyard D. B., Ward I. M. and Bonfield W. (1998). ‘Fibre reinforcement of ceramic/polymer composites for a major load-bearing bone substitute material’, Compos Sci Technol, 58, 419–434. Liste S., Serra J., González P., Borrajo J. P., Chiussi S., León B. and Pérez-Amor M. (2004), ‘The role of the reactive atmosphere in pulsed laser deposit of bioactive glass films’, Thin Solid Films, 453–454, 224–228. Lopez-Esteba S., Saiz E., Fujino S., Oku T., Suganuma K. and Tomsia A. P. (2003). ‘Bioactive glass coating for orthopedic metallic implants’, J Eur Ceram Soc, 23, 2921– 2930. Marcolongo M., Ducheyne P. and LaCourse W. C. (1997). ‘Surface reaction layer formation in vitro on a bioactive glass fiber/polymeric composite’, J Biomed Mater Res, 37, 440– 448. Marcolongo M., Ducheyne P., Garino J. and Schepers E. (1998). ‘Bioactive glass fiber/ polymeric composite bond to bone tissue’, J Biomed Mater Res, 39, 161–170. Middleton J. C. and Tipton A. J. (2000). ‘Synthetic biodegradable polymers as orthopedic devices’, Biomaterials, 21, 2335–2346. Misra S. K., Mohn D., Brunner T. J., Stark W. J., Philip S. E., Roy I., Salih V., Knowles J. C. and Boccaccini A. R. (2008). ‘Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB)/Bioglass® composites’, Biomaterials, 29, 1750–1761. Niemelä T., Kellomäki M. and Törmälä P. (2004). ‘In vitro decradation of osteoconductive poly-L/DL-lactide/β-TCP composites’, Key Eng Mater, 254–256, 509–512.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

244 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Niemelä T., Niiranen H., Kellomäki M. and Törmälä P. (2005a). ‘Self-reinforced composites of bioabsorbable polymer and bioactive glass with different bioactive glass contents. Part I: initial mechanical properties and bioactivity’, Acta Biomaterialia, 1, 235–242. Niemelä T. (2005b). ‘Effect of β-tricalcium phosphate addition on the in vitro degradation of self-reinforced poly-L,D-lactide’, Polym Degrad Stab, 89, 492–500. Niemelä T. and Kellomäki M. (2007). ‘Three composites of bioactive glass and PLA-copolymers: Mass loss and water absorption in vitro’, Key Eng Mater, 330–332, 431–434. Niemelä T., Niiranen H. and Kellomäki M. (2008). ‘Self-reinforced composites of bioabsorbable polymer and bioactive glass with different bioactive glass contents. Part II: in vitro degradation’, Acta Biomaterialia, 4, 156–164. Niiranen H. and Törmälä P. (1999a). ‘Self-reinforced bioactive glass-bioabsorbable polymer composites’, in Neenan T., Marcolongo M. and Valentini R. F., Biomedical Materials – Drug Delivery, Implant and Tissue Engineering, vol. 500, 267–272. Niiranen H. and Törmälä P. (1999b). ‘Bioabsorbable polymer plates coated with bioactive glass speheres’, J Mater Sci: Mater Med, 10, 707–710. Niiranen H., Pyhältö T., Rokkanen P., Paatola T. and Törmälä P. (2001). ‘Bioactive glass 13-93/P(L/DL)LA composites in vitro and in vivo’, Key Eng Mater, 192–195, 721–724. Niiranen H., Pyhältö T., Rokkanen P., Kellomäki M. and Törmälä P. (2004). ‘In vitro and in vivo behavior of self-reinforced bioabsorbable polymer and self-reinforced bioabsorbable polyer/bioactive glass composites’, J Biomed Mater Res, 69A, 699–708. Paatola T., Pirhonen E. and Törmälä P. (2001). ‘Coating of bioactive glass (13-93) fibers with bioabsorbable polymer’, Key Eng Mater, 192–195, 717–720. Pirhonen E., Grandi G. and Törmälä P. (2001). ‘Bioactive glass fiber/polylactide composite’, Key Eng Mater, 192–195, 725–728. Pirhonen E., Niiranen H., Niemelä T., Brink M. and Törmälä P. (2006a). ‘Manufacturing, mechanical characterization, and in vitro performance of bioactive glass 13-93 fibers’, J Biomed Mater Res Part B: Appl Biomater, 77B, 227–233. Pirhonen E., Moimas L. and Brink M. (2006b). ‘Mechanical properties of bioactive glass 9-93 fibers’, Acta Biomaterialia, 2, 103–107. Pirhonen E. and Törmälä P. (2006c). ‘Coating of bioactive glass 13-93 fibers with biomedical polymers’, J Mater Sci, 41, 2031–2036. Pyhältö T., Lapinsuo M., Pätiälä H., Rokkanen P., Niiranen H. and Törmälä P. (2004). ‘Fixation of distal femoral osteotomies with self-reinforced poly(L/DL)lactide 70:30/ bioactive glass composite rods. An experimental study on rats’, J Mater Sci: Mater Med, 15, 275–281. Radder A. M., Van Loon J. A., Puppels G. J. and Van Blitterswijk C. A. (1995). ‘Degradation and calcification of PEO/PBT copolymer series’, J Mater Sci: Mater Med, 6, 510–517. Rich J., Jaakkola T., Tirri T., Närhi T., Yli-Urpo A. and Seppälä J. (2002). ‘In vitro evaluation of poly-ε-caprolactone-co-DL-lactide)/bioactive glass composites’, Biomaterials, 23, 2143–2150. Roether J. A., Boccaccini A. R., Hench L. L., Maquet V., Gautier S. and Jérôme R. (2002). ‘Development and in vitro characterization of novel bioresorbable and bioactive composite materials based on polylactide foams and Bioglass® for tissue engineering applications’, Biomaterials, 23, 3871–3878



245

Ruuttila P., Niiranen H., Kellomäki M., Törmälä P., Konttinen Y. T. and Hukkanen M. (2006). ‘Characterization of human primary osteoblast response on bioactive glass (BaG 13-93) coated poly-L,DL-lactide (SR-PLA70) surface in vitro’, J Biomed Mater Res Part B: Appl Biomater, 78B, 97–104. Schrooten J., Van Oosterwyck H., Vander Sloten J. and Helsen J. A. (1999). ‘Adhesion of new bioactive glass coating’, J Biomed Mater Res, 44, 243–25. Stamboulis A. and Hench L. L. (2001). ‘Bioresorbable polymers: Their potential as scaffolds for Bioglass® composites’, Key Eng Mater, 192–195, 729–732. Stamboulis A., Hench L. L. and Boccaccini A. R. (2002). ‘Mechanical properties of the biodegradable polymer sutures coated with bioactive glass’, J Mater Sci: Mater Med, 13, 843–848. Tomsia A. P., Saiz E., Song J. and Bertozzi C. R. (2005). ‘Biomimetic bonelike composites and novel bioactive glass coatings’, Adv Eng Mater, 7, 999–1004. Törmälä P. (1992). ‘Biodegradable self-reinforced composite materials; manufacturing structure and mechanical properties’, Clin Mater, 10, 29–34. Tukiainen M., Arstila H., Hupa L. and Kellomäki M. (2006a), ‘Composite structures of bioactive glass/biodegradable polymer hybrid yarns’, 20th European Conference on Biomaterials, European Society of Biomaterials, Nantes, France, 27 September– 1 October. Tukiainen M., Suokas E., Arstila H., Hupa L. and Kellomäki M. (2006b). ‘A porous bioactive and biodegradable load-bearing composite manufactured using hybrid yarns of bioactive glass and biodegradable polymer’, 10th Annual Seminar and Meeting, ‘Ceramics, cells and tissues’, Faenza, Italy. Wang C. X., Chen Z. Q. and Wang M. (2002). ‘Fabrication and characterization of bioactive glass coatings produced by the ion beam sputter deposition technique’, J Mater Sci: Mater Med, 13, 247–251. Wang M., Yue C. Y. and Chua B. (2001). ‘Production and evaluation of hydroxyapatite reinforced polysulfone for tissue replacement’, J Mater Sci: Mater Med, 12, 821–826. Wang M. (2003). ‘Developing bioactive composite materials for tissue replacement’, Biomaterials, 24, 2133–2151. Zhou Z., Ruan J., Zou J., Zhou Z. and Shen X. (2007). ‘Bioactivity of bioresorbable composite based on bioactive glass and poly-L-lactide’, Trans Nonferrous Met Soc China, 17, 394–399. Zhou Z., Yi Q., Liu X., Liu L. and Liu Q. (2009). ‘In vitro degradation behaviors of polyL-lactide/bioactive glass composite materials in phosphate-buffered solution’, Trans Nonferrous Met Soc China, 63, 575–586.


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

12 Bioactive glasses for wound healing M. SHAH MOHAMMADI, C. STÄHLI and S. N. NAZHAT, McGill University, Canada

Abstract: Bioactive and soluble glasses have the potential for use in woundhealing applications. Metal oxides may be incorporated into either silicate- or phosphate-based glasses in order to controllably release antimicrobial or angiogenic ions. This chapter discusses the effect of various ions on the chemical and biological properties of these types of glasses. Key words: silicate-based glass, phosphate-based glass, antimicrobial, angiogenesis, wound healing.

12.1

Introduction

Wound healing can represent a major challenge in medicine and is often mentioned in relation to surgical sutures, oral infections or chronic wounds such as nonhealing diabetic ulcers. The potential of some metallic cations as wound-healing agents has been widely studied and may prove useful as an alternative to traditional antibiotic treatments, which are increasingly associated with bacterial resistance. Bioactive glasses can be doped with various metallic oxides to provide an interesting strategy of controllably delivering these metal ions in wound-healing applications. The dissolution products of bioactive and soluble glasses can affect different phases of wound healing. In this chapter, the use of the antibacterial or antimicrobial properties of several metallic cations, which aim to prevent infection and thereby to shorten the inflammatory phase of wound healing, will be presented, and the potential of bioactive glasses to stimulate angiogenesis (which allows the re-establishment of blood supply in an injured area) will be discussed.

12.2

Silicate-based versus phosphate-based bioactive glasses

Silicate-based glasses (SGs) are an interesting class of bioactive material developed by Hench et al. (1971) for biomedical applications [1–4]. Bioglass® is a commercially available bioactive glass which is based on 45S5 composition that corresponds to 45SiO2-24.5CaO-24.5Na2O-6P2O5 (wt%). Over the last three decades, bioactive SGs have generated significant interest for bone tissue regeneration applications. When exposed to physiological fluid in vivo, these glasses form a surface hydroxycarbonate apatite (HCA) layer that has the ability 246 © Woodhead Publishing Limited, 2011


247

to create a direct bond with bone through interactions with collagen synthesized by osteoblasts (bone forming cells) [5]. The biological behaviour of glasses depends on the relative proportion of bridging oxygen bonds to non-bridging bonds in the phases of the materials [6]. The bone-bonding mechanism has been extensively researched by Hench and co-workers [6]. In summary, three general processes occur when bioactive SGs are immersed in an aqueous solution: leaching, dissolution and precipitation. Leaching occurs due to the release of alkali or alkaline earth elements, usually by cation exchange with H+ or H3O+. Since these cations are not part of the glass network, and only modify the network by forming non-bridging oxygen bonds, ion exchange occurs easily. Network dissolution occurs by the breaking of -Si-O-Si-O-Si- bonds through the action of hydroxyl (OH–) ions. This occurs locally and releases silica into the solution. The hydrated silica (SiOH) formed on the glass results in a silica-rich gel layer formation. In the precipitation reaction, phosphate and calcium ions released from the glass along with those from the solution form a calcium-phosphate-rich (CaP) layer on the surface. The nucleation and formation of an apatite layer that is considered to be the main factor for the bioactivity of Bioglass® is due to the solubility of the phosphate species [7]. Although SGs have had great success in many clinical applications, questions have been raised as to their long-term degradation. In addition, more rapid solubility is required in wound-healing applications. The limitations associated with SGs have led to continual research for new materials for bone defect repair. Soluble phosphate-based glasses (PGs) are an example of one of these materials: they provide a diverse range of solubility and can be predicted and controlled by altering the glass composition [8–10]. In recent decades numerous PG formulations of binary, ternary and quaternary compositions have been developed. Different compositions have been investigated for biomaterials and tissue engineering applications [11–16]. The 3D network structure of the SG is a SiO4 tetrahedron, owing to the strong affinity of silicon towards oxygen [17]. Phosphates are common in nature because phosphorous also has an affinity towards oxygen. They also have a tetrahedral unit; however, the PO4 unit is quite different from that of SiO4 (Fig. 12.1).

12.1 Tetrahedral unit of (a) silicate and (b) phosphate. Reproduced from [17].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

248 1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

Bioactive glasses

Phosphorous has a charge of 5+ while silicon has a charge of 4+ and in the case of P2O5, a terminal double bond forms, since the oxygen atoms that are not shared between phosphate tetrahedral share their two unpaired electrons with the P5+. Such terminal oxygen limits the connectivity of PGs and decreases their interatomic forces and rigidity. In addition, PGs contain fewer cross-links while having a higher number of terminal oxygen atoms when they are mixed with metal oxides resulting in more flexibility of PO43− tetrahedra [18]. Therefore, the range of glass formation is wider in PGs compared to the SG system [17], which allows for more antimicrobial metal oxides, such as Ga2O3, Ag2O, and CuO, being incorporated into the glass structure. Phosphate tetrahedra can be classified by the number of oxygen atoms that are shared with other phosphate tetrahedra, referred to as bridging oxygen atoms (BOs). This classification leads to phosphate tetrahedra labelled with Qi where i is the number of BOs and ranges between 0 and 3. Figure 12.2 shows the various Q species. While the three-dimensional vitreous P2O5 has Q3 tetrahedra, the addition of modifying oxides results in depolymerization of the network through P-O-P bond cleavage. Kirkpatrik and Brow proposed the depolymerization model predicting that the dominant Qi varies based on Q3→ Q2→ Q1→ Q0 as the amount of modifying oxides increases [17]. The aqueous dissolution mechanisms and the stability of resultant anionic species of PGs and SGs are different. While silicate species can be repolymerized, phosphate chains are stable in solution: once dissolved they form new structures without any resemblance to the original glass structure [20]. In order to determine

12.2 Representation of PO4 tetrahedra. Reproduced from [19].



249

the dissolution mechanisms of these glasses, the types of reactions that can occur between the glasses and water should be considered, which involve [21]: acid/ base reactions, which can aid glass dissolution by disrupting the ionic interactions between chains, hydration of the entire phosphate chains; and hydrolysis reactions, which result in the cleavage of P-O-P bonds and lead to the ultimate destruction of the phosphate network to produce orthophosphate. Given that pure P2O5 is chemically unstable due to the hydrolysis of the P-O-P bond, the addition of glass-modifying metal oxides improves its stability by forming P-O-M+ bonds that are generally more stable towards atmospheric hydrolysis [22]. A ternary glass system based on 45P2O5-xCaO-(55-x)Na2O, where x was between 8 and 40 mol%, was developed by Franks et al. [10]. An inverse relationship between CaO content and the solubility rate was observed which was linear over time for glass containing up to 20 mol% CaO. It was suggested that Ca2+ and its interaction with the glass network is a dominant factor in the solubility rate of these formulations. In high CaO-containing glasses, an ion exchange process accompanied by a gradual breakdown of the glass network are two suggested responsible processes for dissolution. The potential application of this glass system to bone regeneration was studied by Salih et al. [23]. They suggested that greater amounts of Ca2+ are released with low dissolution rate glass, which has an essential role in cell activation mechanisms affecting cell growth. However, a sharp change in pH associated with high release rates of Na+ and PO43− may have an adverse effect on cells in highly soluble glasses. In order to increase the durability of PGs, a study by Ahmed et al. [13] incorporated Fe2O3 through partial substitution of Na2O leading to the evolution of 50P2O5-(30, 40 or 45)CaO-Na2O-xFe2O3 (x was between 1 and 5 mol%). Fe2O3 addition up to 5 mol% resulted in a significant decrease in the solubility rate by one order of magnitude and an increase in the glass transition temperature, indicative of greater cross-linking. Fibres of 50P2O5-30CaO-15Na2O-5Fe2O3 were also shown to allow adhesion and proliferation of myoblasts and the formation of myotubes in vitro, however, the dissolution rate was very low. Abou Neel et al. [24] showed that the overall surface energy of the glass decreased with increasing Fe2O3 leading to a significant decrease in the dissolution rate which could be due to the formation of more hydration-resistant P-O-Fe bonds. PG fibres with the composition of 50P2O5-30Ca(15-x)Na2O-5Fe2O3x-SiO2 have been studied by Patel et al. [8]. It was found that the substitution of 5 mol% NaO2 for SiO2 led to a more rapid dissolution of the glass fibres and resulted in 60% mass loss on day 4. The effect of adding other compounds such as TiO2 and MgO on the solubility, cell attachment, viability, and proliferation have also been studied [10, 25]. It has been suggested that TiO2 could decrease the solubility rate of PG, probably due to the formation of a TiO5 or TiO4 structural unit and the strong Ti-O-P bonds. Also, the addition of upto 5 mol% TiO2 supported the cell attachment and maintained high cell viability [11].


1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

1 2 3 4 5 6 7 8 9 10 1 2 3 4 5 6 7 8 9 20 1 2 3 4 5 6 7 8 9 30 1 2 3 4 5 6 7 8 9 40 1 2 43X

250

Bioactive glasses

12.3

Antibacterial properties of bioactive glasses

Based on the unstable nature of these glasses in an aqueous environment, they may be particularly interesting for achieving a controlled release of certain antimicrobial ions. Since it is important to prevent infection and resulting inflammation in wound-healing procedures, these glasses may provide interesting properties for wound-healing applications. Many of the diseases, such as airway infections in cystic fibrosis patients, chronic wound and sinus infections caused by P. aeruginosa, are associated with biofilm formation [26–29]. Biofilm formation occurs due to microbial surface attachment, cell proliferation, matrix production and detachment [30]. Biofilmassociated bacteria cause a decreased susceptibility to antibiotics [31], disinfectants [32] and clearance by host defences [28, 33]. It has been found that Ga3+ ions can hinder P. aeruginosa growth and biofilm formation in vitro by decreasing bacterial Fe uptake and interfering with Fe signalling via the transcriptional regulator Iron Starvation Sigma Factor (pvdS) [34]. Ga2O3-doped PGs can be used as a unique system for the delivery of gallium ions in a controlled manner [34]. Since ions incorporated into the glass structures are not a separate phase, the overall dissolution rate of the glass would indicate their release rate. Vallapil et al. [35, 36] has shown that Ga2O3-doped PGs hold promise as antimicrobial agents, and could provide some advantages over conventional therapeutic agents. They showed that the net bactericidal effect was due to Ga3+, and a concentration as low as 1 mol% Ga2O3 was adequate to provide a potent antibacterial effect. Other metal ions, such as silver and copper, have been incorporated into bioactive and soluble glasses for potential uses in wound-dressing applications to prevent infections [37]. As metallic silver reacts with moisture on the skin surface or with wound fluids, silver ions are released that damage bacterial RNA and DNA, hence inhibiting replication. For this reason, silver-containing materials provide interesting properties for wound repair applications [38–40]. Silver has been shown to aid healing in sterile skin wounds in rat models by reducing the inflammatory and granulation tissue phases of healing and inducing epidermal repair [40]. Continuous silver-release products have a bactericidal action, and manage wound exudates and odour [38, 39]. Incorporation of Ag2O into bioactive glass compositions to minimize the risk of microbial contamination through the leaching of Ag+ ions that have potential antimicrobial activity has been reported [41–43]. A bioactive SG composition doped with Ag2O was shown to be bacteriostatic to elicit rapid bactericidal reaction [43]. It was also confirmed that 3 wt% Ag2O incorporation conferred antimicrobial properties to the glass without compromising the glass bioactivity [43]. It has also been suggested that surgical sutures combined with bioactive phases should result in practical bioactive composite materials with an extensive range of applications in wound healing, augmentation devices and tissue engineering scaffolds [44–47]. Blaker et al. [48] developed novel silver-doped bioactive glass powder (AgBG)-coated



251

12.3 Surface morphology of (a) Vicryl® and (b) Mersilk® sutures coated with AgBG particles by slurry dipping. Average particle size is

Bioactive Glasses: Materials, Properties and Applications (Woodhead Publishing in Materials)

Building Decorative Materials (Woodhead Publishing in Materials)

Mineral Wool: Production and Properties (Woodhead Publishing in Materials)

Materials in Sports Equipment, Volume 2 (Woodhead Publishing in Materials)

Electroceramics: Materials, Properties, Applications

Thermal Barrier Coatings (Woodhead Publishing in Materials)

Hydrometallurgy: Principles and Applications (Woodhead Publishing in Materials)

Bioactive Materials in Medicine: Design and applications

Biomedical Composites (Woodhead Publishing in Materials)

MIG Welding Guide (Woodhead Publishing in Materials)

Electroceramics: Materials, Properties, Applications

Electroceramics: Materials, Properties, Applications

High Temperature Superconductors (Woodhead Publishing in Materials)

Sustainability of Construction Materials (Woodhead Publishing in Materials)

Sintering of Advanced Materials (Woodhead Publishing in Materials)

Glasses and Amorphous Materials

Polymer Carbon Nanotube Composites: Preparation, Properties and Applications (Woodhead Publishing in Materials)

Nanolithography and Patterning Techniques in Microelectronics (Woodhead Publishing in Materials)

Welding and Joining of Magnesium Alloys (Woodhead Publishing in Materials)

Biomaterials, Artificial Organs and Tissue Engineering (Woodhead Publishing in Materials)

Corrosion Control in the Aerospace Industry (Woodhead Publishing in Materials)

Ceramic Materials: Processes, Properties and Applications

Friction in Textile Materials (Woodhead Publishing in Textiles)

NANOSTRUCTURED MATERIALS Processing Properties and Potential Applications

Superplasticity and Grain Boundaries in Ultrafine-Grained Materials (Woodhead Publishing in Materials)

Ion Exchange Materials: Properties and Applications

Monolithic Materials: Preparation, Properties and Applications

Photorefractive Optics: Materials, Properties, and Applications

Photorefractive Optics: Materials, Properties, and Applications

Ceramic Materials: Processes, Properties, and Applications

Corrosion of Magnesium Alloys (Woodhead Publishing in Materials)

Bioactive Glasses: Materials, Properties and Applications (Woodhead Publishing in Materials)

Building Decorative Materials (Woodhead Publishing in Materials)

Mineral Wool: Production and Properties (Woodhead Publishing in Materials)

Materials in Sports Equipment, Volume 2 (Woodhead Publishing in Materials)

Electroceramics: Materials, Properties, Applications

Thermal Barrier Coatings (Woodhead Publishing in Materials)

Hydrometallurgy: Principles and Applications (Woodhead Publishing in Materials)

Bioactive Materials in Medicine: Design and applications

Biomedical Composites (Woodhead Publishing in Materials)

MIG Welding Guide (Woodhead Publishing in Materials)

Electroceramics: Materials, Properties, Applications

Electroceramics: Materials, Properties, Applications

High Temperature Superconductors (Woodhead Publishing in Materials)

Sustainability of Construction Materials (Woodhead Publishing in Materials)

Sintering of Advanced Materials (Woodhead Publishing in Materials)

Glasses and Amorphous Materials

Polymer Carbon Nanotube Composites: Preparation, Properties and Applications (Woodhead Publishing in Materials)

Nanolithography and Patterning Techniques in Microelectronics (Woodhead Publishing in Materials)

Welding and Joining of Magnesium Alloys (Woodhead Publishing in Materials)

Biomaterials, Artificial Organs and Tissue Engineering (Woodhead Publishing in Materials)

Corrosion Control in the Aerospace Industry (Woodhead Publishing in Materials)

Ceramic Materials: Processes, Properties and Applications

Friction in Textile Materials (Woodhead Publishing in Textiles)

NANOSTRUCTURED MATERIALS Processing Properties and Potential Applications

Superplasticity and Grain Boundaries in Ultrafine-Grained Materials (Woodhead Publishing in Materials)

Ion Exchange Materials: Properties and Applications

Monolithic Materials: Preparation, Properties and Applications

Photorefractive Optics: Materials, Properties, and Applications

Photorefractive Optics: Materials, Properties, and Applications

Ceramic Materials: Processes, Properties, and Applications

Corrosion of Magnesium Alloys (Woodhead Publishing in Materials)

Recommend Documents