Orthopaedic Bone Cements

Orthopaedic bone cements

Related titles: Bioceramics and their clinical applications (ISBN 978-1-84569-204-9) 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. Joint replacement technology: New developments (ISBN 978-1-84569-245-2) This book comprehensively reviews developments in joint replacement technology, covering the most pertinent materials science and engineering issues, specific joints, clinical trial results and sterilisation techniques. Part I discusses biomechanics, tribology, the chemical environment of the body and an overview of materials and engineering of joint replacement. The second part of the book reviews specific materials, bearing surfaces and bone cements, in addition to the failure mechanisms and lifetime prediction of joints. Part III discusses the biological environment and interaction of replacement joints. Specific joints are analysed in detail in Part IV. Biomedical polymers (ISBN 978-1-84569-070-0) This book reviews the structure, processing and properties of biomedical polymers. It discusses the various groups of biopolymers including natural polymers, synthetic biodegradable and non-biodegradable polymers. Chapters also review the application of biomedical polymers in such areas as scaffolds for tissue engineering, drug delivery systems and cell encapsulation. The book also considers the use of polymers in replacement heart valves and arteries, in joint replacement and in biosensor applications. Details of these and other Woodhead Publishing materials books as well as materials books from Maney Publishing can be obtained by: • •

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Orthopaedic bone cements Edited by Sanjukta Deb

Woodhead Publishing and Maney Publishing on behalf of The Institute of Materials, Minerals & Mining CRC Press Boca Raton Boston New York Washington, DC

Cambridge England

Woodhead Publishing Limited and Maney Publishing Limited on behalf of The Institute of Materials, Minerals & Mining Published by Woodhead Publishing Limited, Abington Hall, Granta Park Great Abington, Cambridge CB21 6AH, England www.woodheadpublishing.com Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2008, Woodhead Publishing Limited and CRC Press LLC 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 Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-376-3 (book) Woodhead Publishing ISBN 978-1-84569-517-0 (e-book) CRC Press ISBN 978-1-4200-9302-5 CRC Press order number WP9302 The publishers’ policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acid-free and elementary chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by SNP Best-set Typesetter Ltd., Hong Kong Printed by TJ International Limited, Padstow, Cornwall, England

Contents

Contributor contact details

xii

Part I Bone cements in medicine

1

1

3

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11 1.12 1.13 2

2.1 2.2 2.3 2.4 2.5

Bone disease M. Phillips and K. Joshi, King’s College London, UK Bone disease Osteoporosis Osteomalacia and rickets Paget’s disease (osteitis deformans) Malignancy Hyperparathyroidism Osteomyelitis Prosthesis-related infection Current state of the use of bone cement in the United Kingdom Other potentially useful applications of cements in bone Summary Acknowledgement References Hip replacements B. M. Wroblewski, P. D. Siney and P. A. Fleming, The John Charnley Research Institute, UK Introduction General principles Bone cements Fixation of components with cement Long-term results

3 8 15 18 20 25 26 28 30 33 34 34 34 41

41 42 43 44 44 v

vi

Contents

2.6 2.7 2.8

Long-term problems Future trends References

45 46 46

3

Knee replacements H. Pandit and B. H. van Duren, Nuffield Orthopaedic Centre, UK Relevant anatomy of the knee joint Conditions causing knee arthritis Clinical and radiological assessment of an arthritic knee Treatment options for osteoarthritis Indicators for total knee replacement Evolution of knee replacements Implant design rationales The cemented total knee replacement Future trends References

48

Vertebroplasty and kyphoplasty J. Yeh, St Bartholomew’s and The Royal London Hospitals, UK Introduction Vertebral compression fractures Kyphoplasty and vertebroplasty Clinical outcomes Clinical experiences with injectable bone cements Future trends References

74

3.1 3.2 3.3 3.4 3.5 3.6 3.7 3.8 3.9 3.10 4

4.1 4.2 4.3 4.4 4.5 4.6 4.7 5

5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8 5.9

Antibiotic-impregnated polymethylmethacrylate (PMMA) spacers in hip surgery K. Anagnostakos and J. Kelm, Universitätsklinikum des Saarlandes, Germany Introduction Construction of hip spacers Diagnosis of infection Consistency and function of antibiotic-loaded hip spacers Mechanical stability and behaviour of hip spacers Pathogenic organisms Antibiotic choice Antibiotic elution Clinical experience

48 49 52 54 56 57 58 63 70 70

74 74 77 78 81 86 87

92

92 94 95 96 97 98 98 100 101

Contents

5.10 5.11 5.12 5.13

Girdlestone or spacer? New techniques Conclusions References

6

Commercial aspects and delivery systems of bone cements R. Kowalski and R. Schmaehling, DePuy CMW, UK Introduction Commercial aspects: cemented versus cementless fixation Mixing and delivery systems of bone cements Regulatory aspects Future trends Conclusions References

6.1 6.2 6.3 6.4 6.5 6.6 6.7 7 7.1 7.2 7.3 7.4 7.5 7.6 7.7 7.8 7.9 7.10

Wear particles and osteolysis N. Patil and S. B. Goodman, Stanford Medical Center, USA Introduction Cellular cascade and mediators of osteolysis Morphology and bioreactivity of wear debris Osteolysis Cement debris Polyethylene wear debris Metallic wear debris Ceramic wear debris Future trends References

vii 107 108 108 109

113 113 114 121 135 136 137 137 140 140 141 143 144 146 150 152 153 154 155

Part II Materials

165

8

167

8.1 8.2 8.3 8.4 8.5 8.6 8.7 8.8 8.9

Acrylic bone cement: genesis and evolution S. Deb and G. Koller, King’s College London, UK Introduction Hip and knee joint Acrylic bone cement Regulatory perspectives Sterilization of bone cements Fluoride and other additives in bone cements Other applications of acrylic bone cement Conclusions References

167 169 170 175 176 178 179 180 181

viii

Contents

9

Poly(methylmethacrylate) bone cement: chemical composition and chemistry B. Vázquez Lasa, Institute of Polymer Science and Technology (CSIC) and CIBER-BBN, Spain Introduction Chemical composition Setting process Polymerisation reaction and kinetics Free radical studies on acrylic bone cements Curing parameters: standards Rheology of acrylic bone cements Low molecular weight residuals after curing Future trends References

9.1 9.2 9.3 9.4 9.5 9.6 9.7 9.8 9.9 9.10 10

10.1 10.2 10.3 10.4 10.5 10.6

Calcium phosphate bone cements M. P. Ginebra, Technical University of Catalonia (UPC), Spain Introduction Historical overview: calcium phosphate cements versus acrylic cements Chemistry of calcium phosphate cements Basic properties of calcium phosphate cements Applications of calcium phosphate cements: present and future perspectives References

183

183 184 186 187 190 191 193 195 197 200 206

206 207 208 213 221 224

Part III Properties of bone cement

231

11

233

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8

Mechanical properties of bone cements N. Dunne, Queen’s University Belfast, UK Introduction Nature and structure of polymethylmethacrylate bone cement Test standards Mechanical properties: short-term strength of polymethylmethacrylate bone cement Factors affecting the microstructure–mechanical properties relationship Modification of acrylic bone cements Summary References

233 234 236 239 249 256 256 257

Contents

12

12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10 12.11 12.12 13

13.1 13.2 13.3 13.4 13.5 13.6 13.7 13.8 13.9

Fracture toughness and fatigue characteristics of bone cements A. B. Lennon, Trinity College Dublin, Ireland Introduction Factors affecting fracture toughness and fatigue resistance of bone cement The effect of loading mode on fracture and fatigue The effect of porosity on fracture and fatigue The effect of inclusions on fracture and fatigue The effect of cement chemistry on fracture and fatigue Bone cement failure in joint replacements Failure at interfaces Residual stress and the initiation of damage Viscoelasticity, creep and creep–fatigue interaction Fracture and fatigue properties References Dynamic mechanical properties of bone cements S. N. Nazhat, McGill University, Canada; and J. V. Cauich Rodríguez, Centro de Investigación Cientifica de Yucatan A.C., Mexico A brief introduction to viscoelasticity in polymers Regions of viscoelasticity Theory of dynamic mechanical analysis (DMA) Material properties measured through dynamic mechanical analysis Applications of dynamic mechanical analysis in the characterisation of polymeric biomaterials Applications of dynamic mechanical analysis in the characterisation of bone cements Future trends Conclusions References

ix

265 265 265 266 272 274 275 277 279 281 282 283 285 296

296 297 299 301 302 303 307 308 308

Part IV Enhancing the properties of bone cements

311

14

313

14.1 14.2 14.3

Antibiotic-loaded bone cements S. Deb and G. Koller, King’s College London, UK Introduction Demands on acrylic bone cement systems Antibiotic-loaded bone cements

313 315 316

x

Contents

14.4

The effect of antibiotics on the mechanical properties of the bone cement Release of antibiotics from bone cements Other additives in bone cement Conclusions References

14.5 14.6 14.7 14.8 15

15.1 15.2

15.3 15.4 15.5 15.6 15.7 15.8 16

16.1 16.2 16.3 16.4 16.5 16.6 16.7 17 17.1 17.2

Modifications of bone cements J. San Román, B. Vázquez Lasa, M. R. Aguilar and L. F. Boesel, Institute of Polymer Science and Technology (CSIC) and CIBER-BBN, Spain Introduction Modulation of the hydrophilic/hydrophobic character of bone cements and the consequences on the properties and behaviour of such formulations Control of the fiexibility or stiffness of acrylic cement formulations Modification of formulations with bioactive and functionalised components with pharmacological activity Improvement and modulation of the radiopaque character New biohybrid composites for bone and cartilage regeneration Future directions in the design and development of cements with specific properties References Design of bioactive bone cement based on organic–inorganic hybrids T. Miyazaki, Kyushu Institute of Technology, Japan; and C. Ohtsuki, Nagoya University, Japan The need for bioactive bone cements How do materials exhibit bioactivity? Design of bioactive bone cements using bioactive ceramics Bioactive organic–inorganic hybrids Design of bioactive bone cements based on organic–inorganic hybrids Conclusions References Clinical aspects of calcium phosphate bone cements S. Larsson, Uppsala University Hospital, Sweden Introduction Material characteristics

318 321 327 328 328 332

332

334 340 343 346 349 350 351

358

358 359 364 365 369 372 372 377 377 378

Contents

17.3 17.4 17.5 17.6 17.7

Surgical technique Complications during surgery Clinical applications Future trends References

xi 382 383 386 396 397

Contributor contact details

(* = main contact)

Chapter 2

Editor

Professor B. M. Wroblewski*, P. D. Siney and P. A. Fleming The John Charnley Research Institute Wrightington Hospital Wigan UK E-mail: [email protected]

Dr Sanjukta Deb King’s College London Floor 17 Guy’s Tower GKT Dental Institute London Bridge London SE1 9RT UK E-mail: [email protected]

Chapter 1 Dr M. Phillips* Department of Orthopaedics King’s College Hospital King’s College London UK E-mail: [email protected]

Chapter 3 H. Pandit* and B. H. van Duren Nuffield Orthopaedic Centre Windmill Road Headington Oxford OX3 7LD UK E-mail: [email protected]

Chapter 4 Kush Joshi King’s College London Medical School King’s College London UK E-mail: [email protected]

xii

Dr John Yeh St Bartholomew’s and The Royal London Hospitals Queen Mary, University of London Department of Neurosurgery The Royal London Hospital Whitechapel London E1 1BB UK E-mail: [email protected]

Contributor contact details

Chapter 5

Chapter 8

Dr K. Anagnostakos and Dr J. Kelm Klinik für Orthopädie und Orthopädische Chirurgie Universitätsklinikum des Saarlandes Kirrbergerstr. 1 D-66421 Homburg/Saar Germany E-mail: [email protected]

Dr Sanjukta Deb* and Dr Garrit Koller King’s College London Floor 17 Guy’s Tower GKT Dental Institute London Bridge London SE1 9RT UK E-mail: [email protected]

xiii

Chapter 9 Chapter 6 Dr R. Kowalski* and R. Schmaehling DePuy CMW Cornford Road Marton Blackpool FY4 4QQ UK E-mail: [email protected]

Chapter 7 Dr Nilesh Patil* Visiting Assistant Professor Department of Orthopedic Surgery Stanford Medical Center 300 Pasteur Drive Stanford CA 94305 USA E-mail: [email protected] Stuart B. Goodman Robert L. and Mary Ellenburg Professor of Surgery Department of Orthopedic Surgery Stanford University 300 Pasteur Drive, R153 Stanford CA 94305 USA E-mail: [email protected]

Dr B. Vázquez Lasa Department of Biomaterials Institute of Polymer Science and Technology (CSIC) C/Juan de la Cierva, 3 28006 Madrid Spain E-mail: [email protected]

Chapter 10 Professor Maria-Pau Ginebra Department of Materials Science and Metallurgical Engineering Technical University of Catalonia (UPC) Av. Diagonal 647 08028 Barcelona Spain E-mail: [email protected]

Chapter 11 Dr Nicholas Dunne School of Mechanical and Aerospace Engineering Queen’s University Belfast Ashby Building Stranmillis Road Belfast BT9 5AH UK E-mail: [email protected]

xiv

Contributor contact details

Chapter 12

Chapter 15

Dr A. B. Lennon Trinity Centre for Bioengineering Parsons Building School of Engineering Trinity College Dublin 2 Ireland E-mail: [email protected]

Dr J. San Román,* B. Vázquez Lasa, M. R. Aguilar and L. F. Boesel Department of Biomaterials Institute of Polymer Science and Technology (CSIC) C/Juan de la Cierva, 3 28006 Madrid Spain E-mail: [email protected]

Chapter 13 Dr Showan N. Nazhat* Department of Mining and Materials Engineering McGill University MH Wong Building 3610 University Street Montreal, Quebec Canada H3A 2B2 E-mail: [email protected] Dr Juan V. Cauich Rodríguez Centro de Investigación Cientifica de Yucatan A.C. Calle 43 # 130 Col. Chuburna De Hidalgo Merida, Yucatan Mexico E-mail: [email protected]

Chapter 14 Dr Sanjukta Deb and Dr Garrit Koller King’s College London Floor 17 Guy’s Tower GKT Dental Institute London Bridge London SE1 9RT UK E-mail: [email protected]

Chapter 16 Dr T. Miyazaki* Graduate School of Life Science and Systems Engineering Kyushu Institute of Technology 2-4 Hibikino, Wakamatsu-ku Kitakyushu 808-0196 Japan E-mail: [email protected] Professor C. Ohtsuki Department of Crystalline Materials Science Graduate School of Engineering Nagoya University Furo-cho, Chikusa-ku Nagoya 464-8603 Japan

Chapter 17 Professor Sune Larsson Department of Orthopedics Uppsala University Hospital S-751 85 Uppsala Sweden E-mail: sune.larsson@ortopedi. uu.se

Part I Bone cements in medicine

1 Bone disease M. P H I L L I P S and K. J O S H I, King’s College London, UK

Abstract: Bone diseases are conditions that result in the impairment of normal bone function and can make bones physically weaker by the deterioration of their structure. The most common result in bone disease is fracture of the affected bone but many other problems are encountered by patients. This chapter first discusses bone remodelling, resorption and formation processes in detail, and then gives a description of the many conditions that can result in impaired bone quality (e.g. osteoporosis, osteomalacia, etc.). Section 1.9 outlines the current state of the use of bone cements in the United Kingdom and this is followed by a description of other potential applications of cements as their formulations are improved and developed. Key words: bone disease, bone fracture, osteoporosis, bone cement.

1.1

Bone disease

Bone diseases are conditions that result in the impairment of normal bone function and can make bones physically weaker by the deterioration of their structure. Fracture of the affected bone is generally the main result of bone disease; however, fractures are just one of many problems encountered by patients. In defining bone diseases, which may be alleviated by cement reconstruction or reinforcement, the processes within bone may be classified by aetiology, and then by whether the lesions produced by the disease tend to be monofocal or polyfocal. Insertion of cement into the axial or appendicular skeleton is clearly an invasive procedure, and so treatment of polyfocal or diffuse generalised bone disease in the clinical setting would be directed at the areas of the skeleton most at risk of mechanical insufficiency. These rather obvious assertions give us a framework for which bone diseases may, currently and in the future, be amenable to tissue augmentation with cement. The common bone diseases that currently lead to mechanical insufficiency of the skeleton are listed in Table 1.1. These disease processes are further classified in the right-hand column into whether they are usually involved in only one bone, i.e. are monostotic (MO), or are involved in more than one bone (polystotic (PO)), or whether they can be both (MO/PO). Currently, very few conditions are treated by insertion of cement. The common usage of poly(methylmethacrylate) (PMMA) cement in 3

4

Orthopaedic bone cements

Table 1.1 Common bone diseases that currently lead to mechanical insufficiency of the skeleton Congenital

Osteogenesis imperfecta Vitamin D-resistant rickets Osteopetrosis Melorheostosis Fibrous dysplasia Haemophilic pseudotumour

PO PO PO

Metabolic

Renal osteodystrophy Diabetes mellitus (Charcot neuropathic joints) Gout Osteomalacia

PO MO MO PO

Endocrine

Hyperparathyroidism

MO/PO

Neoplastic

Benign (aneurysmal and simple bone cysts) enchondromata Giant cell tumours or Malignant secondary deposits and primary sarcomas

MO

MO/PO MO

MO/PO MO PO MO

Iatrogenic

Steroid-induced osteoporosis Cavitation produced by loose implants

PO PO MO

Traumatic

Stress fractures High-impact fractures with bone loss Fracture non-union

MO MO MO

Infective

Osteomyelitis (acute bacterial, tuberculous and other rarer organisms)

MO

Idiopathic

Paget’s disease

MO/PO

Degenerative

Osteoporosis Ganglion cysts of bone Osteoarthritis (to stabilise implants) Arthrosis (to fuse joints)

PO MO MO/PO MO

Vascular

Avascular necrosis of bone (most commonly hip, lunate, second metatarsal head) Intraosseous haemangioma

MO MO

orthopaedic surgical practice is illustrated in Table 1.2. Other, usually monostotic, cavitary defects are normally treated by insertion of cancellous bone graft, harvested from the patient’s iliac wing, in order to provide a matrix that will become incorporated. In contained defects, graft under load incorporates well in most cases; however, in uncontained defects, surgical measures are taken to contain the graft, such as using metallic mesh. The disadvantage of bone grafting is that the harvest site is often very painful postoperatively, and has a significant risk of developing complica-

Bone disease

5

Table 1.2 Current common usage of poly(methylmethacrylate) (PMMA) cement in orthopaedic surgical practice Malignant neoplasia Joint replacement Osteomyelitis

Osteoporosis

In pathological fractures with extensive bone loss, and patients with relatively short life expectancy Following infection of a joint prosthesis, antibioticloaded PMMA often used as temporary spacer May occasionally be used with antibiotic in bone at risk of fracture or with a view to later removal (as antibiotic carrier) To augment fixation, cement is occasionally inserted into medulla and once set, drilled to allow screws to hold more effectively In vertebral collapse, bone cement is inserted into the vertebra after balloon kyphoplasty

tions. The most common complications are haematoma or infection. Nerve injury can occur, often leaving an area of numbness down the lateral thigh. The donor site can lead to the propagation of a fracture, which can cause prolonged immobility, and in the most severe cases may require internal fixation. Cancellous bone has little structural strength, and so if structural properties are required, a cortico-cancellous graft is usually taken. Again, the iliac wing is the commonest site for harvest. Neighbouring bone can frequently be used for donor bone, and so the distal radius provides bone for scaphoid fracture grafting, the calcaneus provides for subtalar joint fusion, the fibula or tibia for ankle fusion, or the tibia for midfoot fusions. Cortico-cancellous grafts revascularise only very slowly, from the ends towards the centre, and the risk of fracture persists for several years for larger grafts. For this reason, vascularised grafts have been developed. The most frequently used free vascularised graft is the fibula. Donation of a fibula seems to cause only moderate morbidity, and it has a reasonable vascular pedicle for anastomosis. However, the operation is often lengthy and persistence of flow in the graft cannot be guaranteed. With these hazards of bone graft donation and the inherent properties of graft occasionally not being ideal (e.g. not strong enough, not enough volume), it is natural that orthopaedic surgeons should be keen to see progress in the field of biomaterials as applied to the reconstruction of bone defects. Other chapters in this book will deal with the properties of bone cements, and this chapter will outline the areas of clinical orthopaedics in which bone cements are being successfully used. Bone diseases and conditions that mimic bone disease where bone cements can be used are also discussed. The use of bone cements is primarily related to fixation of fractures; however, they can also be used in palliation and treatment of joint

6

Orthopaedic bone cements

infection as will be discussed. The main areas of where bone cement is currently used are as laid out in Table 1.2.

1.1.1 Bone remodelling Bone is a complex tissue that goes through constant remodelling throughout life, even once skeletal maturity has occurred and growth has been completed. It is composed of both mineral and organic phases that contribute two-thirds and one-third of the skeletal weight, respectively. By virtue of the design of these phases, as well as the fact that it is formed from a combination of dense compact bone and cancellous (trabecular) bone that is reinforced at points of stress, this specialised connective tissue is the principal mechanical support for the body. Further functions of bone also include it being the main store of calcium in the body as well as the site of haematopoiesis. The process of bone remodelling is a coupled process with localised removal of old bone (resorption), and replacement of the cavity that is created by the formation of new bone. This coupling mechanism ensures that an equivalent amount of bone is laid down following the previous resorption phase (Parfitt 1982). The two principal cells that are responsible for the turnover of bone matrix are osteoblasts and osteoclasts. Osteoblasts are derived from local mesenchymal stem cells. Their function in bone remodelling as well as in bone growth, is to synthesise bone matrix (osteoid). The bone matrix is subsequently mineralised, a process that is also regulated by osteoblasts. Osteoclasts develop from haemopoietic stem cells of the macrophage/monocyte lineage, and have the ability to remove this bone matrix when activated. (Walker 1973). At any one time, roughly 20% of the cancellous bone surface is undergoing bone remodelling. The process occurs in units, known as basic multicellular units (BMUs), which are small packets of cells that are geographically and chronologically separated from other packets of remodelling. This allows the turnover of bone on multiple surfaces (Frost 1991).

1.1.2 Bone resorption For bone resorption to occur, a cascade takes place that results in the removal of both the mineral and organic constituents of bone matrix, by osteoclasts. This process is also aided by osteoblasts. For bone resorption to take place, the first stage that is required is the recruitment and dissemination of osteoclast progenitors to bone. Through a mechanism involving cell-to-cell interaction with osteoblast stromal cells, the precursor cells proliferate and differentiate into osteoclasts. Osteoblasts remove the unmineralised osteoid layer by the production of a variety of proteolytic enzymes,

Bone disease

7

including matrix metalloproteinases (MMPs), collagenase and gelatinase (Meikle et al. 1992). This process allows osteoclasts access to the mineralised layer underneath. Through various cascades the osteoclasts then adhere to the mineralised surface and are subsequently activated by the effects of local factors in cells of the osteoblast lineage (Martin and Ng 1994). Activated osteoclasts resorb bone through enzymatic release and through the production of hydrogen ions. Hydrogen ions that are generated within the cell by the enzyme carbonic anhydrase II are responsible for the dissolution of the mineral (Laitala and Vaananen 1994). Once dissolution of the mineralised matrix has taken place, then proteolytic enzymes act upon the collagenous organic matrix, resulting in its degradation (Hill et al. 1994). Once resorption has taken place, the osteoclasts undergo apoptosis, which can be stimulated by transforming growth factor (TGF)-β (Roodman 1996). There are several mechanisms proposed as to why there is an arrest in osteoclastic activity. These include: • osteoclasts having a limited life span; • the accumulation of high concentrations of calcium in the resorption lacunae controlling osteoclast activity causing rapid cell retraction and, in the longer term, inhibiting enzyme release (Zaidi 1990); • as TGF-β and related peptides are released from the matrix during resorption, osteoclasts are as a result inhibited (Pfeilschifter et al. 1990a, 1990b). The period of quiescence seen in bone occurs after the maximum eroded depth has been reached and lasts for roughly 9 days. This stage sees the disappearance of osteoclasts and the appearance of macrophage-like cells on the bone surface. These macrophage-like cells are rich in collagenase and are thus able to remove residual matrix.

1.1.3 Bone formation In order for bone formation to occur, a complex cascade takes place that involves the proliferation of primitive mesenchymal cells, differentiation into osteoblast precursor cells (osteoprogenitor, pre-osteoblast), maturation of osteoblasts, formation of matrix and mineralisation. The resorption cavity that is created by the osteoclasts is converged on by osteoblasts to form osteoid. This begins to mineralise after 13 days. It takes between 4 and 6 months for a cavity to be filled. It takes between 4 and 6 months for an entire cycle of bone turnover. For cortical bone the annual rate of turnover is 4%, and for trabecular bone 25%. Many factors regulate bone remodelling, including systemic hormones as well as local factors. Therefore any change to the internal environment

8

Orthopaedic bone cements

Standing

1.1 Multiple metatarsal stress fractures (arrows) in an adult with mild osteogenesis imperfecta.

can change this delicate balance between bone formation and bone resorption. In childhood, there is bone growth, and the process of bone formation exceeds bone resorption, resulting in increasing bone mass. This reaches an optimum – the peak bone mass; in most people this is reached in their early to late twenties, dependent upon the ethnicity, genetic make-up and gender of each individual. Once the peak bone mass has been reached, there is a steady decline in the bone mass of individuals, thus making the elderly more susceptible to fractures. The main reason that the process of bone remodelling occurs throughout life is to adapt to mechanical stress as a consequence of physical exercise and other mechanical factors, but it also crucial for calcium homeostasis. Normally in adults there is a balance between the amount of bone formed by osteoblasts and bone that is resorbed by osteoclasts. However, in bone diseases such as osteoporosis, osteogenesis imperfecta (Fig. 1.1) and Paget’s disease this process is imbalanced and mechanical instability results. Malignant neoplasia and infection can also have a detrimental effect on the process of bone remodelling, which can lead to significant morbidity to patients due to the integrity of the bone being challenged.

1.2

Osteoporosis

Osteoporosis was first described in 1940 by Fuller Albright as a condition of impaired bone formation due to oestrogen deficiency (Albright 1989). Though a large proportion of those who suffer from osteoporosis are postmenopausal women, there are a number of sufferers who have different pathways leading to this disease process. Osteoporosis is a systemic skeletal

Bone disease

9

1.2 Left hip replaced due to osteoporotic fracture, with subsequent fracture below the stem fixed with plate. Right hip fixed after fracture, but with failure of fixation secondary to poor-quality bone stock.

disease that is associated with diminished bone mineral density and altered bone microarchitecture (Felsenberg and Boonen 2005). It can also be defined on the basis of bone mineral density (BMD), with a BMD that is less than 2.5 standard deviations below the mean for young people considered as indicating osteoporosis. This results in reduced bone strength (Fig. 1.2); as a consequence, the structural and material composition of bone can no longer adapt to the mechanical and physical demands of the skeleton, resulting in fragility fractures, which is often what sufferers of osteoporosis first present with (Seeman and Delmas 2006). Osteoporosis can be classified as either a primary disease, or a secondary disease. Primary osteoporosis is bone loss that is a consequence of the normal ageing process. Secondary osteoporosis is due to specific welldefined clinical disorders or due to medications, with bone loss above that which would be expected for age and natural menopause. Osteoporosis is diagnosed and its progression is investigated using dual energy X-ray absorptiometry (DXA), which measures BMD. The primary treatment is through the use of bisphosphonates to prevent fractures.

1.2.1 Epidemiology Fractures of the hip, vertebral body and the distal forearm have long been perceived as the typical fractures associated with osteoporosis. The systemic effect of osteoporosis, however, means that there is a heightened risk of sustaining any form of fracture, and once a patient sustains one form of

10

Orthopaedic bone cements

Table 1.3 Prevalence of osteoporotic fractures in the United Kingdom and United States Estimated annual incidence USA Type of fracture

UK

USA

Vertebral compression fractures Hip fractures Wrist fractures Fractures of other limbs

120 000 70 000 50 000

700 000 250 000 250 000 300 000

fracture, they are at greater risk of subsequent fractures (Kanis et al. 2004). Postmenopausal women are especially at risk of sustaining fractures, and the estimated risk for a 50-year-old woman to obtain an osteoporotic fracture in the remainder of her lifetime is 40–50%, compared with 13–22% for men (Melton et al. 1992). Furthermore, the annual incidence in the United States of fractures due to osteoporosis is estimated to be 1.5 million, of which 71% of fractures are experienced by women (Burton and Mendel 2003). Despite the already large numbers of fractures, the problems faced by health authorities are increasing as the elderly population is the fastest growing age group in the world. For example, the cost for hip fractures worldwide in 1990 was US $34 800 million. In 2050 this figure is estimated to rise to US $131 500 million (Johnell 1997) (Table 1.3). Although vertebral fractures are the fractures most commonly experienced (Ross 1997) by patients with osteoporosis, hip fractures are regarded as the most severe – with patients experiencing high morbidity and great loss of quality of life. Only 50% of patients who suffer from osteoporotic hip fractures actually regain prefecture status as judged by their ability to walk and the need for aids at home (Sernbo and Johnell 1993) (Fig. 1.2). Annually, 1% of women aged 65 years and 2.8% of women aged between 75 and 79 years, suffer from a new vertebral fracture (Johnell and Kanis 2005). Despite the high incidence of vertebral fractures, patients are often not seen. In fact, only a third of all vertebral deformities identified radiologically are seen by specialists, and less than 10% result in hospital admission (Sambrook and Cooper 2006). This is despite the many co-morbidities that are associated with these types of fractures, including pain which is common, as well as disability due to thoracic kyphosis that can eventually hinder the patient’s breathing. The importance of identifying osteoporosis early so that it can be treated is emphasised by the fact that only a quarter of vertebral fractures are as a result of falls. The remainder of these fractures occur as a consequence of

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routine daily activities such as bending or lifting light objects (Cooper et al. 1992). Furthermore, patients who suffer fragility fractures are at an increased risk of other types of fractures. For example, a vertebral fracture leads to a ten-fold increase in the risk of subsequent vertebral fractures (Melton et al. 1999). Similar trends are apparent for forearm fractures and hip fractures, with an increased likelihood of future fractures occurring in a similar area. It has been shown that there are increased changes in the microarchitecture of cancellous and cortical bone with resulting rapid bone loss after a fracture, and this would weaken the bone and reduce resistance to mechanical force (Crabtree et al. 2001; Mayhew et al. 2004). Fractures, however, are also seen at distant sites more often in patients with a previous history of fracture. Importantly, it has been shown that this increased risk of fracture is beyond the explanation of measured BMD.

1.2.2 Pathophysiology Patients with osteoporosis are at greater risk of fractures due to reduced bone strength as already discussed; however, these patients have an increased rate of falls due to co-morbidities that are common with the elderly population. This combination results in the high number of fractures seen. The development of osteoporosis occurs as a result of the following factors, in both its primary and secondary forms: • peak bone mass is less than optimal; • excessive bone resorption causes loss of bone mass and structural damage; • inadequate bone formation occurs in response to bone resorption. Peak bone mass acquisition occurs by the age of 30 years for the majority of the population. In general, males have larger bone mass than females; however, pre-puberty there is no gender-related difference in bone mass in any skeletal site (Glastre et al. 1990). The difference in bone mass that is present between males and females occurs after puberty. Cortical width increases by periosteal bone formation in boys, through androgen, growth hormone and insulin-like growth factor release. In girls, oestrogen inhibits periosteal apposition, resulting in narrower bones than boys, with cortical widening occurring through endocortical apposition which is limiting (Schoenau et al. 2000). Furthermore, males go through a longer period of bone mass gain, which further contributes to them having larger bone mass size and a greater cortical thickness than females (Seeman 1997). Peak bone mass is the maximum BMD that is achieved by an adult and is a major predictor of BMD in later life. It is determined by sex, ethnicity and body size, and varies within individuals according to the region of the bone. Studies in twins have shown that up to 80% of variance in peak bone

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Orthopaedic bone cements

mass is accounted for by genetic factors (Sambrook et al. 1994; Howard et al. 1998). Environmental factors such as nutrition, hormonal status, exercise, medications and illness account for the remainder of the variance. Low BMD before the onset of menopause or the natural ageing process predicts a higher rate of fracture over the next 10 years for individuals (Stewart et al. 2006). Therefore it is important that adequate amounts of calcium and vitamin D are taken, as well as the performance of weight-bearing exercises to ensure maximal BMD. Some bone loss begins to occur before the onset of menopause at the end of the third decade in both sexes at distinct skeletal sites such as the proximal femur and vertebrae (Riggs et al. 1986). The negative BMU balance is a consequence of an early reduction in bone formation within each individual BMU, and at this stage is not due to an increase in resorption by osteoclasts. As discussed previously, bone remodelling is a process that occurs throughout life, and occurs to enable adaption to physical stress, as well as repair to microdamage. This process becomes less efficient with age, and especially as a consequence of menopause in women, with bone material and structural properties degrading with time (Seeman 2002). An increasing amount of bone remodelling takes place with an increasingly negative bone balance in the many BMUs, with more bone removed than being replaced. The fall in BMD is as a consequence of delay in bone formation within the many resorption cavities. The cavities are eventually partially filled by bone formation, but due to the decreased rate that is seen, as well as the increased remodelling activity of osteoclasts, low BMD results with an increasingly worse bone quality, which greatly increases the risk of fracture (Felsenberg and Boonen 2005). In postmenopausal woman this is especially significant due to the role of oestrogen.

1.2.3 Role of oestrogen As shown above, postmenopausal woman are at greater risk of osteoporosis and as a consequence fractures due to bone loss. The reduced structural integrity of bone, as well as reduced BMD, is mainly as a result of oestrogen deficiency. It has been shown in morphological studies and through the measurement of biochemical markers that bone remodelling is accelerated at menopause, with markers of both resorption and bone formation increased (Ebeling et al. 1996). Bone loss occurs due to an imbalance between the activities of osteoblasts and osteoclasts. The coupled bone remodelling process is no longer matched in osteoporosis as would normally occur throughout most of life. The primary driving force seems to be increased bone resorption. There is also rapid and continuous bone loss for several years after menopause, which indicates that there is an impaired bone formation response, as would normally occur in response to resorption

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and mechanical loading. This therefore suggests that oestrogen is both anticatabolic and anabolic (Lee et al. 2003). There is an initial very rapid fall in the BMD of many of the BMUs, which increases the porosity of bone, due to the rapid rate of remodelling that occurs after menopause (Riggs et al. 1998). Oestrogen receptors are expressed on osteoblasts, osteoclasts and osteocytes (Weitzmann and Pacifici 2006). As well as this, these receptors are also present on bone marrow stromal cells, which are the precursors of osteoblasts. These cells provide physical support for osteoclasts, T-cells, β-cells and most other cells in human bone marrow. The deficiency of oestrogen results in T-cells releasing a variety of inflammatory cytokines, as a result of interaction between the immune system and bone cells. Interleukin-1 (IL-1) and tumour necrosis factor are two of the cytokines that are released. Both are osteoclastogenic pro-inflammatory which would be negatively regulated in the presence of oestrogen (Pacifici 1996; Cenci et al. 2003). This enhanced formation of osteoclasts, coupled with their prolonged lifespan, results in the progressive loss of trabecular bone. Other cytokines that are released include interleukin-7 (IL-7), which inhibits osteoblast differentiation and activity, and causes premature apoptosis of osteoblasts (Girasole et al. 1992; Weitzmann and Pacifici 2005). The increasing number of osteoclasts results in a net loss of bone. Furthermore, there is an increase in the number of unfilled cavities where bone resorption has taken place; these form stress risers, which are vulnerable sites that can easily perforate and result in microfractures. Often, complete loss of trabecular plates is seen in postmenopausal woman, leaving no template upon which new bone formation can take place (Khosla et al. 2006). Oestrogen deficiency is also observed in men, associated with the bone loss that is seen with age. Testosterone is only implicated in bone formation (Falahati-Nini et al. 2000). Although men do not undergo accelerated bone remodelling as females do, it has been shown that age-related decreases in bioavailable oestradiol below 40 pmol/l may be an important cause of bone loss in elderly men (Falahati-Nini et al. 2000). Bone remodelling rises moderately in life with men, unlike in women, with loss of primarily trabecular bone occurring in a linear fashion. Primarily a reduction in the volume of bone formed is implicated in the pathogenesis of osteoporosis in men, as opposed to an increase in the volume of bone removed, with the result that connectivity is better preserved.

1.2.4 Secondary osteoporosis As previously defined, in a patient with osteoporosis where an underlying cause or factor can be found other than ageing or menopause, secondary osteoporosis is implicated. There are several causes for secondary osteoporosis, with primary treatment being treatment of the underlying condition

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Orthopaedic bone cements

or the removal of the contributing factor that is causing the osteoporosis. Often, patients may be on medications such as glucocorticosteroids that are required for long-term treatment for other co-morbidities, which can result in secondary osteoporosis. However, the treatment of these other conditions cannot be stopped, and in these situations lifestyle modifications such as increased calcium and vitamin D intake are required to prevent fractures. There are, however, many causes of secondary osteoporosis and an in-depth discussion is beyond the scope of this chapter. Secondary osteoporosis is common in both men and women, being the commonest form of osteoporosis that is found in men. Elderly patients and postmenopausal woman have secondary causes often in addition to their primary form. In fact, one-third of women with postmenopausal osteoporosis have identifiable secondary causes that further contribute to their bone loss (Stein and Shane 2003). In men, secondary causes of osteoporosis account for 65% of cases with bone loss leading to fractures. Glucocorticoid excess Glucocorticoid-induced osteoporosis is the commonest form of secondary osteoporosis (Bijlsma 1997). Bone loss is observed in 40–60% of patients Table 1.4 Common causes of secondary osteoporosis Endocrine disorders: • Hypogonadism • Hyperparathyroidism • Cushing’s disease • Thyrotoxicosis • Acromegaly • Adrenal insufficiency • Pregnancy-induced osteoporosis Malignancy and bone marrow disorders: • Carcinomatosis • Multiple myeloma • Leukaemia • Lymphoma • Anaemia • Gaucher’s disease Nutritional and lifestyle: • Vitamin K deficiency • Vitamin C deficiency – scurvy • Malnutrition • Smoking • Malabsorption Other causes: • Immobilisation • Reflex sympathetic dystrophy

Chronic disorders: • Rheumatoid arthritis • Ankylosing spondylosis • Chronic renal disease • Chronic pulmonary disease • Sarcoidosis • Systemic mastocytosis • Congestive heart failure • Multiple sclerosis Drug induced: • Long-term glucocorticoid treatment • Alcohol excess • Heparin • Anticonvulsants • Cyclosporin A • Methotrexate Genetic: • Osteogenesis imperfecta • Ehlers–Danlos syndrome • Marfan’s syndrome • Cystic fibrosis

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that are on long-term glucocorticoids, and pathological fractures have been observed in 16–67% of patients. Vertebrae and ribs are the most common site for fractures in these patients; however even the risk of hip fractures is doubled (Van Staa et al. 2000). Bone loss is mainly seen in the first 6–12 months of treatment, with the trabecular bone primarily affected (Reid 2000). Glucocorticoids affect bone remodelling by both reducing bone formation and increasing resorption. This occurs by decreasing osteoblastogenesis and osteoblast lifespan as well as by inducing osteocyte apoptosis. Reduction in gonadal and adrenal hormones and a negative calcium balance as a result of its reduced gastrointestinal absorption, coupled with increased urinary excretion, result in accelerated osteoclast maturation (Lane 2001). The end result of these changes is in an increased likelihood of bone fractures in these patients. Alcohol Although a moderate level of alcohol intake is associated with an increase in bone mass, excessive amounts of alcohol are associated with bone loss. This form of osteoporosis is common and is a commonly neglected cause for bone loss at all ages, and the likelihood of fractures is further increased due to a greater likelihood of falls. Osteoporosis is not restricted to patients with cirrhosis. Bone loss occurs in people who abuse alcohol, and the development of cirrhosis in these patients contributes to the severity of the bone disease. Spinal ostopenia can be observed in up to 50% of patients that present in hospital with alcoholism, and fractures of the ribs and vertebrae occur in 30% of this population (Israel et al. 1980; Lindsell et al. 1982). The mechanism by which alcohol causes bone loss is a combination of decreased calcium absorption, liver failure and a toxic effect on osteoblast function.

1.3

Osteomalacia and rickets

Bone-mineral homeostasis is extremely complex with many of the pathways still not fully understood. The role of the parathyroid hormone (PTH)– calcitonin–1,25(OH)2 vitamin D3 (calcitriol) axis is known to be critical in bone-mineral metabolism, and alterations in this pathway lead to abnormalities of the skeletal system (Rowe 2004). There are two main forms of vitamin D: vitamin D3 (cholecalciferol), which is formed after exposure to sunlight or ultraviolet light and vitamin D2 (ergocalciferol), which is obtained by irradiation of plants or plant materials or foods. Vitamin D3 is first hydroxylated in the liver and then subsequently hydroxylated in the kidney to form 1,25-dihydroxyvitamin D3

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Orthopaedic bone cements

(1,25(OH)2D). This is the metabolite that stimulates calcium absorption from the gut. This is also the metabolite that has an effect on the bone and kidneys, and as with the intestine, it stimulates calcium transport from the organs to the blood. The production of 1,25(OH)2D is stimulated by PTH. PTH is also regulated through this pathway, as there is a negative feedback loop through calcium as well as through 1,25(OH)2D that affects its production. Rickets and osteomalacia are disease processes that are a consequence of severe vitamin D deficiency. Both are different expressions of the same disease, with incomplete mineralisation of the osteoid matrix due to inadequate absorption and/or utilisation of calcium. The term osteomalacia is more descriptive of the general pathological process that occurs in both diseases, which is ‘bone softening’. The inadequate mineralisation of the matrix has a detrimental impact on the bone remodelling cycle, injuring the corticoendosteal tissue (Wharton and Bishop 2003). In this disease process, rickets is specifically related to children where there is defective bone growth. The impaired mineralisation of the osteoid matrix here occurs during growth and as a consequence this affects the epiphyseal growth plate and compromises both cortical and trabecular bone. In osteomalacia most surfaces of trabecular and cortical bone are covered with thick osteoid seams. There is greater secretion of PTH as there is no negative feedback from calcium and 1,25(OH)2D due to the deficiency of vitamin D, which results in greater resorption from bones to restore serum calcium levels, and as a consequence there is greater bone turnover. This process causes bone loss, which predominantly is from cortical bone, and may contribute to the pathogenesis of osteoporosis. Fractures can result from the increased bone turnover, as well as from the inadequate mineralisation process (Lips et al. 2001). The characteristic changes that result from the inability to calcify the intercellular matrix in the deepest layers of the physis are seen in rickets. The cellular part of the physis is thicker than normal, but the newly formed bone in the metaphysis is weak and may be indented and have a characteristic cup shape. Thin trabeculae surrounded by unusually wide uncalcified osteoid seams are characteristic of osteomalacia. In milder cases the bone may in fact look normal; however, the microarchitecture may be severely compromised and this can result in crush injuries and fractures occurring more easily. In more severe cases, long bone cortices are thinner and signs are present of new or old stress fractures. Causes of osteomalacia, or rickets, can be dietary deficiency of vitamin D, deficiency of vitamin D metabolites, intestinal malabsorption and renal disease. Other causes are calcium and/or phosphorus deficiency and primary bone mineralisation defects (hereditary hypophosphatasia) (Holick 2003).

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1.3.1 Risk factors Risk factors for vitamin D deficiency are pigmented skin in temperate climate, low sunshine exposure, premature and dysmature birth, obesity, malabsorption, advanced age (aged skin produces much less vitamin D than the skin in younger people), vegetarian diet, restrictive diets and being exclusively breastfed for long periods (Mughal 2002; Pettifor 2004). Vitamin D deficiency is high in elderly people compared with adults, especially those in residential homes and nursing homes, and those with hip fractures (Lips et al. 2001). This is most likely as a result of co-morbidities that result in the individuals being indoors for long periods of time, without any sun exposure. The prevalence of vitamin D deficiency in those who are institutionalised is 75% (Holick 1994). Severe vitamin D deficiency can also be seen in those with pigmented skin, particularly people of Afro-Caribbean descent, as highly pigmented skin makes ultraviolet light less efficacious (Holick 1994). Dietary rickets is now less common in the developed world as a consequence of improved diets. However, there is an increasing proportion of cases of rickets resulting from congenital abnormalities due to disorders within the pathway for vitamin D metabolism, referred to as ‘vitamin Dresistant rickets’, because vitamin D supplements fail to generate active vitamin D metabolites. The commonest form of this disorder is familial hypophosphataemic rickets, which is an X-linked disorder that has a dominant inheritance. The features begin to appear in early childhood and result in severe bony deformity, with children below normal height. Treatment for these children is with vitamin D and phosphate. Bony deformities often require bracing and osteotomy in current practice. The commonest form of osteomalacia in adults is due to malabsorption from the intestine, most commonly seen in coeliac disease, and occasionally in Crohn’s disease, or where extensive surgical resection of the small intestine has taken place. Renal and hepatic disorders can also result in osteomalacia. Long-term treatment for epilepsy may result in an induction of liver enzymes, which degrade vitamin D to inactive metabolites.

1.3.2 Clinical features Rickets Characteristic findings that are often seen include stunting of growth, deformities in the upper and lower limbs (genu varum or genu valgum), metaphyseal widening, palpable enlargement of the costochondral junctions (rachitic rosary), frontal prominence, horizontal depression along the lower border of the chest (Harrison’s groove), insufficient weight gain and distal tibial bowing. On X-ray examination it is possible to see thickening and widening

18

Orthopaedic bone cements

of the physes. Serum calcium and phosphate levels are diminished as is urinary output of calcium. The condition responds rapidly to vitamin D; however, those who are left untreated develop long-bone deformities which may later require corrective osteotomy. Osteomalacia The characteristic pathological feature seen in osteomalcia is the evidence of spontaneous incomplete fractures known as Looser’s zones that are often present in long bones or in the pelvis. Otherwise, symptoms are vague and have a slow rate of progression. Symptoms include backache, muscle weakness and bone pain, all of which may be present for many years. Often patients first present with a vertebral compression fracture or an ‘insufficiency fracture’ of the femur or tibia. Serum calcium and phosphate concentrations are often diminished, as well as values for 1,25(OH)2D. A bone biopsy is needed for diagnosis, demonstrating an increase in nonmineralised osteoid. Treatment with vitamin and calcium supplements is often effective.

1.4

Paget’s disease (osteitis deformans)

In Paget’s disease of the bone, enlargement and thickening of the bone are characteristic features. Despite this thickened appearance of the bone on the outside, the internal architecture of the bone in Paget’s disease is extremely brittle and very susceptible to fractures. In Paget’s disease there are rapidly alternating phases of bone resorption and formation occurring constantly. The primary cells that are implicated in this disease process are osteoclasts. It has been shown that there are increased numbers present in bone, and that they are uncharacteristically large compared with normal osteoclasts and are hypernucleated. These abnormal osteoclasts initiate an increase in bone resorption at affected skeletal sites. Part of this activity is ascribed to IL-6 produced by osteoblasts (Ishimi et al. 1990). As a consequence of this resorption, new bone formation occurs rapidly so as to fill the cavity that is created by the overactive osteoclasts. However, due to the rapid nature of the bone formation, architecturally the bone is disorganised and is structurally less stable than normal bone, making it more susceptible to deformity and fracture (Siris et al. 2006). Paget’s disease is most common in people in North America, Britain, Germany and Australia, affecting more than 3% of people over the age of 40. It affects men and women equally, although in this regard there are geographical differences. Patients with the disease present mostly at the age of 50; however, it is likely the disease process started well before that. This is because the commonest sites at which Paget’s disease presents itself are

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(a)

(b)

1.3 (a) Paget’s disease of the pelvis (‘mottled’ architecture); (b) Paget’s disease of the tibia (‘sabre tibia’) with site of stress fracture arrowed.

the pelvis (Fig. 1.3(a)) and the tibia (Fig. 1.3(b)), where the disease localises itself and often goes unnoticed for years until there is progression of the disease. In 15% of patients, there is monostatic disease. Despite certain bones being more favourable than others, Paget’s disease can affect any bone. Due to the disorganised remodelling of the bone as a result of the increased activity of osteoblasts and osteoclasts, there are many changes that are seen on X-ray. The usual changes seen are lytic and sclerotic alterations, characterised by a ‘cotton wool-like’ appearance on X-ray. The changes in the architecture of the bone are most prominently seen in the trabecula where there is a disorderly arrangement of densely calcified

20

Orthopaedic bone cements

tissue. In a cross-section of the tissue it can be seen that medullary spaces are obliterated so that bone has a sclerosed appearance on X-ray film. Bone also ends up encroaching on the spongiosa, which results in the nutrient blood vessels becoming engulfed. This leads to reduced vascularisation with susceptibility to infection. Other X-ray findings include: • • •

cortices becoming thickened, but are irregular; porous structure seen; sclerotic bone changes seen.

Most patients are asymptomatic throughout life, with 60–80% of patients being diagnosed by incidental X-ray findings. However, for those who have more widespread disease, depending on the location and disease progression, there are numerous symptoms that patients can present with. If the skull, for example, becomes involved and thickens, patients may present with deafness due to probable auditory nerve entrapment (Khetarpal and Schuknecht 1990). Kyphosis is also possible in patients, which can result in nerve compression or spinal cord compression. The most common symptoms for patients where there is a greater degree of disease progression are pain and fractures. Pain is a late feature of the disease, and is dull and constant in nature generally as a result of the thickening of the cortices against surrounding tissue. Fractures are common in symptomatic patients, and most commonly occur in the lower limbs due mechanical stresses being applied to the brittle bone architecture. Often these fatigue fractures have difficulty in healing, and deformities such as ‘pseudo-fractures’ result.

1.5

Malignancy

1.5.1 Metastases Malignant tumours that are found in bone are most commonly metastatic deposits is from other sites of primary cancer. The commonest source of deposits is from breast carcinoma. Carcinomas of the prostate, kidney, lung, thyroid, bladder and gastrointestinal tract are also frequent. The skeleton is the most common location where metastatic deposits from cancers are found, and they often produce the greatest morbidity in patients with metastatic disease. In fact, the most common cause of cancer-related pain is metastatic disease in the bone (Mercadante 1997). The red marrow that is present in the axial skeleton may assist the formation of bone metastases with its favourable circulation, cells and extracellular matrix, making it ideal for metastatic deposits. There is also venous blood from the breasts and pelvis that flows into the vertebral-venous plexus of vessels that extend from the pelvis throughout the epidural and

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perivertebral veins, which would explain the incidence of bone metastases in breast and prostate cancer. Specifically for breast cancer, it has been shown that subpopulations of breast cancer lines with a high bone metastatic potential express a variety of genes that confer high capacity to metastasise to bone (Kang et al. 2003). These genes include IL-11 (stimulates osteolysis), connective tissue growth factor (CTGF) (stimulates angiogenesis and osteoblast proliferation), CXCR4 (homing receptor expressed on the cancer cells), which binds the osteoblast-secreted ligand stromal-derived factor-1) and matrix metalloproteinase (MMP-1) (cleaves protein at the bone surface to make the site more palatable for osteoclastic bone resorption). It has been shown in vivo that the cumulative overexpression of the genes mentioned, as well as others that have not been mentioned that also promote bone metastasese, are critical to the process of metastasis to bone and not the individual expression of a few genes (Brown et al. 2007). As shown for breast cancer, similar pathways are believed to be present for other cancers, and make bone an ideal site for metastases; 70% of patents that die from breast or prostate cancer have bone metastases, as demonstrated in post-mortem examination. The incidences from thyroid, kidney, and lung cancer are also high, with around 40% of post-mortem studies showing the presence of bone metastases (Table 1.5). Cancer cells are able to thrive in the bone microenvironment due to the wealth of growth factors present that promote their engraftment and proliferation. Chemotaxis between these malignant cells and the normal bone microenvironment leads to the promotion and sustenance of abnormal cells. In general, the metastatic cells induce osteolysis with abnormal bone formation, so that bone remodelling is compromised. This ultimately results in the destruction of the bone microarchitecture, which is detrimental to the bone integrity and strength. This weakens the connective structure, decreasing the resistance of the skeleton to both compression and bending forces, and making the bone more susceptible to fractures (Kanis et al. 1991).

Table 1.5 Incidence bone metastases at post-mortem examination in different cancers (Galasko 1981) Primary tumour

Incidence of bone metastases (%)

Breast Prostate Thyroid Kidney Lung Gastrointestinal tract

73 68 42 35 36 5

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Orthopaedic bone cements

1.4 Multiple metastases causing fractures in a clavicle from a patient with myeloma.

The mechanism by which bone destruction occurs is through interaction between tumour cells and bone cells. However, the specific mechanism that leads to osteoclast activation differs between the malignancies. For example, in myeloma (Fig. 1.4), bone resorption is the predominant feature due to the secretion of osteoclast-activating substances that have low osteoblastic activity, and thus there is a purely osteolytic process leading to bone destruction. Other cancers such as lung cancer and renal cancer are similar to myeloma and make the patient more susceptible to pathological fractures. Prostate cancer cells act differently in that they produce osteoblast stimulatory factors, probably specific growth factors or acid phosphatase (Kanis et al. 1991). New bone is laid down directly on the trabecular bone surface before osteoclastic resorption can take place, resulting in sclerotic metastases that are less prone to fracture as a consequence of locally increased bone mass (Koutsilieris et al. 1987). Breast cancer metastases of the bone are associated with both osteolytic and osteoblastic bone lesions, with both lytic and sclerotic area present on a plain radiograph. Fractures do commonly occur, and usually occur through the lytic areas (Clavel 1991).

1.5.2 Clinical features On average, a patient with metastatic disease will experience a skeletalrelated event every 3–6 months. The symptoms that are experienced by the patient are very much dependent on the metastatic cancer. In general, osteolytic bone metastases present with bone pain, pathological fractures or hypercalcaemia, or more rarely with swelling over the area of metastatic lesion or neurological complaints. The most common presenting complaint in patients is pain, as many as two-thirds of patients with metastatic bone disease experience severe pain, particularly those who are in advanced stages (Coleman 2006). Pain develops gradually during a period of weeks and months, becoming progressively more severe.

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The common sites of metastatic involvement associated with pain are the base of the skull (in association with cranial nerve palsies and neuralgias), vertebral metastases (producing neck and back pain with or without neurological complications), and pelvic and femoral lesions (producing pain in the back and lower limbs, often associated with mechanical instability and incident pain). Symptoms such as fatigue, anorexia and constipation should be regarded with a high degree of suspicion, as these non-specific symptoms are suggestive of hypercalcaemia. Most commonly this occurs in patients with squamous cell lung cancer, breast and kidney cancers, and some haematological malignancies (myelomas in particular). In most cases the hypercalcaemia is as a result of bone destruction, and osteolytic metastases are present in 80% of cases. The consequence of the osteolytic changes seen in bone result in a reduced load-bearing capacity of the skeletal system and micro-fractures can result, which cause pain. Fractures resulting from these osteolytic changes occur most commonly in the ribs and vertebrae; while fractures of long bones and fractures due to the epidural extension of a tumour into the spine cause the greatest disability.

1.5.3 Diagnosis Bone metastases are often found by mere coincidence during an X-ray examination or as a result of a pathological fracture. Studies have shown that the risk of skeletal complications in both breast and prostate cancer is strongly related to the rate of bone resorption, and the use of the bone resorption marker n-telopeptide of type I collagen (NTX) was useful in the identification of patients at high risk of skeletal complications. This measurement can be done by doing monthly urinary NTX measurements, with highly elevated levels of NTX predictive of skeletal events (Bundred et al. 1996; Coleman 2002) Technetium-labelled radionuclide imaging is the most sensitive means of detecting silent metastatic deposits in bone.

1.5.4 Osteosarcoma Osteosarcomas (Fig. 1.5(a) and (b)) are the second most common primary malignant bone tumours, being accountable for approximately 15% of all biopsy-analysed primary bone tumours. There are a number of subtypes of osteosarcomas that are histologically identified, the most common being the conventional central high grade primary osteosarcoma of bone, which represents 90% of all cases (Schajowicz et al. 1995). Some 75% of all cases that present are in patients between the ages of 15 and 25 years, and it rarely presents in patients younger than the age of 6, or those older than 60 years.

24

Orthopaedic bone cements (a)

(b)

1.5 (a) Plain radiograph of osteosarcoma of the femur showing the soft tissue extent (A) and ‘Codman’s triangle’ (B) of elevated periosteum and new bone formation. (b) Same patient as panel (a), magnetic resonance imaging (MRI) showing full soft tissue (A) and bony (B) extent of tumour on MRI scan.

Osteosarcomas are highly malignant tumours, and generally arise from the bone and spread outwards to the periosteum involving surrounding soft tissues. This often leads to a ‘sunburst’ effect being visible on X-ray due to the outward bone growth seen from the cortex. The primary mechanism of bone destruction in these bone tumours is through osteolysis; however, on X-ray examination it can be shown that there is both an increase in bone resorption as well as in bone formation. Some 80–90% of all osteosarcomas occur in long tubular bones, with the femur, tibia and humerus affected in 85% of all cases.

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X-rays can be used in the diagnosis in the majority of cases; however, biopsy should be carried out regardless, as many of the atypical lesions can be confused for benign lesions. In 10% of all cases, pulmonary metastases are already present when the patient first presents, and due to the painful nature of osteosarcoma, the use of cements can be advocated to improve quality of life. Otherwise, long-term survival after surgical excision and chemotherapy is higher than 60% in patients.

1.5.5 Chondrosarcoma Chondrosarcomas can present as either benign chondromas that have undergone malignant change, or as a primary lesion. Benign lesions undergo malignant change most often when they are present on cartilage-capped exostoses of the pelvis and scapula. This type of tumour most commonly presents in patients in their forties and fifties, and may present either with a non-specific dull ache or, due to their slow-growing nature, as a pathological fracture if the lesion is of medullary origin. These tumours do not respond successfully to chemotherapy and radiotherapy; however, due to their slow-growing nature, they can be excised and be replaced with a prosthesis.

1.6

Hyperparathyroidism

Calcium homeostasis is maintained through the action of PTH, which targets cells in the bone and the kidney in order to maintain serum calcium levels. Its action is to increase serum calcium levels by increasing bone resorption, and by enhancing active reabsorption of calcium from the distal tubules in the kidneys. Furthermore, it enhances the absorption of calcium in the intestine by increasing the production of activated vitamin D. The secretion of PTH is regulated by a feedback mechanism, where large amounts of calcium in the blood inhibit its release.

1.6.1 Primary hyperparathyroidism Primary hyperparathyroidism is associated with specific skeletal disorders such as osteitis fibrosa cystica. This disorder includes subperiosteal resorption of the distal phalanges, tapering of the distal clavicles, brown tumours and the presence of bone cysts of the long bones. The pathophysiology of this disease process is as a consequence of the loss of normal feedback control of PTH by extracellular calcium. The major cause of this is due to parathyroid adenomas, where the parathyroid cells lose their sensitivity to calcium. In contrast, hyperplasia that results in primary hyperparathyroidism results in an increased number of cells leading to bone resorption. The

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increase in circulating levels of PTH increases bone turnover, with an increase in osteoclast-mediated bone resorption as well as osteoblast activity. The overall impact of this is a loss of both cortical and cancellous bone (Khosla et al. 1999). The increasing bone turnover has been shown to produce an increased incidence of vertebral, Colles’, rib and pelvic fractures. However, surgical cure by parathyroidectomy causes an increase in BMD levels.

1.6.2 Secondary hyperparathyroidism A decreased intake of calcium, impaired intestinal absorption of calcium due to ageing or disease, and deficiency in vitamin D can result in secondary hyperparathyroidism (Lips 2001). The active hormonal form, 1,25(OH)2 vitamin D3 (calcitriol) is important for the optimal intestinal absorption of calcium and phosphorus. It also has the role of exerting a tonic inhibitory effect on PTH synthesis. This therefore means that there are dual pathways that can result in secondary hyperparathyroidism (Lips 2001). Vitamin D deficiency and secondary hyperparathyroidism are extremely common in elderly patients that are admitted to hospital with fragility fractures, due to their role in accelerating bone loss and increasing fragility. Furthermore, they can also increase the risk of falls as a result of neuromuscular impairment (Gao et al. 2004).

1.7

Osteomyelitis

Osteomyelitis occurs in children in 85% of all cases, arising from unknown primary foci (nasopharynx) via direct inoculation, but can also be contiguous. The age group of the most commonly affected children is 2–5 years old. The most common infecting pathogen is Staphylococcus aureus; less often Streptococcus pyogenes or S. pneumoniae may be the organism responsible for infection. In young children, especially those between the ages of 2 and 3 years, it is not uncommon to see infection as a result of Haemophilus influenzae. When adults are affected by osteomyelitis, due to the general resistance of bone against infection, the patient may have other co-morbidities such as being debilitated, or being in an immunocompromised state. The primary cause for the infection is not found in 70% of cases, and may be due to a simple boil, an infected ulcer or a urinary tract infection (Concia et al. 2006).

1.7.1 Pathogenesis Bone is normally highly resistant to bacterial colonisation, and in effect if bone infection does take place, there generally are underlying causes.

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Osteomyelitis can therefore occur as a consequence of trauma, the presence of foreign bodies, prostheses or an inoculation of aggressive bacteria (Sugawara et al. 1998). The disease process is an inflammatory suppurative process of the bone marrow in which the endosteum and periosteum are actively involved. Transient bacteraemia can also cause osteomyelitis due to the rich capillary network and large venous channels in bones which can be favourable for the deposition of circulating micro-organisms and their subsequent growth. Initially, acute inflammation develops around the small capillaries and veins in the metaphysis causing a local focus of acute inflammation. This is characterised by hyperaemia, oedema, leukocyte infiltration and purulent transformation. In addition, by the second day of infection, pus can appear in the medulla which can force its way along the Volkmann canals to the surface, where it forms a subperiosteal abscess. The inflammatory response extends through the Haversian canals causing compression of the adjacent vessels and arteries, which results in the blood supply to the trabeculae being compromised. As a result, osteonecrosis takes place. Pieces of bone may separate as sequestra, which act as foreign-body irritants, causing persistent discharge through a sinus, until they escape or are removed. Larger sequestra remain entombed within the cavities of bone. Radiologically it is difficult to see to any definite signs of infection in the first 2 weeks. Specific signs of infection such as new subperiosteal bone growth and transcortical sinus tracts can eventually be seen. If infection spreads through the cortical bone to the periosteal layer, there is soft tissue swelling and periosteal elevation, which can cause an inflammatory response in surrounding tissues. If there is progression to a chronic infection, an enveloping layer of bone is formed beneath the periosteum, known as the involucrum, which is thickened and encloses the infected tissue and sequestra. There is also a tendency for the pus and necrotic debris that accumulate to drain from the inflammatory site to the surface of the skin through perforations (cloacae). These sinuses may persist for months to years.

1.7.2 Clinical features Symptoms are generally variable; however, in the typical form, fever, pain and restricted movement of the affected limb are seen with local inflammation present or septicaemia, depending on the pathway that caused the infection. The infection of the bone can result in its functional loss, as well as the functional loss of the surrounding tissue (Gentry 1987). In adults, the suspicious features are backache and mild fever, as osteomyelitis is most commonly seen in the vertebrae. It may take weeks for signs to be apparent on X-rays, by which time the osteomyelitis has often

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spread to other vertebrae. To confirm clinical diagnosis, pus should be aspirated from the subperiosteal abscess or the adjacent joint. A smear of the aspirate should be examined for cells and organisms. In chronic osteomyelitis, the acute infection would have left sequestra surrounded by dense sclerotic bone. Bacteria may remain dormant for years but giving rise to frequent flare-ups. On X-ray examination bone rarefaction can be seen surrounded by dense sclerosis and cortical thickening, and within this area of bone there may be obvious sequestrum.

1.8

Prosthesis-related infection

With the increasing numbers of elderly in the population, the use of prostheses is becoming more widespread. In the United States alone there were more than 600 000 patients with prosthesic implants in 2006, a figure that is rising annually. Although there has been a decline in the number of infected prosthetic implants, from 5.9% in the 1970s to 1.2% in 2000 in Europe, osteomyelitis as a consequence of prosthetic implants is a major problem (Concia et al. 2006). This is due in particular to the increasing number of resistant strains of bacteria. Infection that is associated with prosthetic joints is typically due to the growth of micro-organisms in biofilms (Gristina 2004). Within biofilms, micro-organisms are enclosed in a polymeric matrix, protected from antimicrobial agents as well as the immune system of the host. This allows the development of organised, complex communities, with structural and functional heterogeneity, resembling multicellular organisms (Costerton et al. 1999). The most commonly cultured organisms are coagulase-negative staphylococci, occuring in 30–43% of all cases (Pandey et al. 2000). The process of attachment of such organisms to the surface of prosthetic implants is a two-step process, as is the case with Staphylococcus epidermidis. The primary attachment is by means of non-specific factors, such as surface tension, hydrophobicity and electrostatic interactions, or by specific adhesins. This is shortly followed by the formation of a biofilm, a process mediated by intercellular adhesins (Darouiche 2001). The process is very much governed by the type of prosthetic device used; in certain prostheses infection with S. epidermidis is not possible. This, however, is not the case with S. aureus, which accounts for 12–23% of all prosthetic infections (Pandey et al. 2000). S. aureus interacts with host proteins such as fibronectin and fibrinogen, and this enables it to cover the device immediately after implantation. Although on insertion implants will be exposed to blood due to the nature of the operation, once inserted – as is the case for all foreign bodies – they remain devoid of a microcirculation, which is crucial for host defence

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and for the delivery of antibiotics. This increased susceptibility to infection is also created by the mere presence of a subcutaneous foreign body, which reduces the minimal inoculum of Staphylococcus aureus that is required to cause infection by a factor of more than 100 000 (Zimmerli et al. 1982). A locally acquired granulocyte defect is also present due to the activation of neutrophils in response to the foreign body. Neutrophils release human neutrophil peptides that deactivate granulocytes (Kaplan et al. 1999). Even in the presence of antimicrobial prophylaxis, infection can occur due to a foreign body, since less than 100 colony-forming units can cause an infection (Zimmerli et al. 1982a).

1.8.1 Onset of infection according to organism Time of onset of infection Staphylococcus aureus and sometimes coagulase-negative staphylococci bacteria cause infection within a month of surgical implantation of a prosthetic implant. If the infection is of such early onset, the cause of infection is likely to be due to bacterial contamination at the time of surgery. Such infections have little involvement in the bone–prosthesis interface, meaning that they are located within the surgical access area and studies of the surrounding tissue can give an indication of the infecting pathogen. Patients present with an acute onset of joint pain, effusion, erythema and warmth at the implant site. Systemic effects such as fever are also commonly present. Cellulitis and the formation of a sinus tract with purulent discharge may also occur during infection. By diagnosing the infection early, there is a good prognosis for patients, and no replacement implant is required. Delayed infection (2–12 months) and late-onset infection (>12 months) can be caused by a number of pathogens. S. epidermidis is the one of note and causes both delayed and late-onset infection. Other organisms such as S. aureus (methicillin-resistant), Escherichia coli and anaerobes are only found in late-onset infection. Delayed-onset infection, as is the case with early-onset infection, is generally due to the presence of bacteria on surgical implantation. The onset of late infections, however, are primarily as a consequence of haematogenous seeding, the most frequent sources being skin, respiratory, dental and urinary tract infections (Maderazo et al. 1988). The infection spread is primarily at the bone–prosthesis interface, with possible soft tissue involvement due to secondary spread (Table 1.6). Patients’ symptoms are dependent on the infective pathogen, but often the patient will have subtle indicators to the presence of an infection. There may be implant loosening and persistent joint pain, which may make it difficult to distinguish from aseptic failure.

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Table 1.6 Infections in hip prostheses and their time of onset over a 16-year period Onset of infection

Percentage of patients with infection

Early infection Delayed infection Late infection

29 41 30

Diagnosis Infection can be detected by the use of plain radiographs if there are studies over time after implantation. Subperiosteal bone growth and transcortical sinus tracts are specific for infection. There may also be indications of loosening that can be indicative of an infection without any other radiological signs. Cultures of tissue surrounding the prosthesis are a reliable means by which to detect the pathogen. At least three tissue specimens should be sampled for culture (Atkins et al. 1998). In the case of low-grade infections, which may be difficult to detect, antimicrobial therapy should be stopped for 2 weeks prior to samples being taken.

1.9

Current state of the use of bone cement in the United Kingdom

1.9.1 Malignant neoplasia Most currently available bone cements are either non-resorbable (PMMA) or have insufficient strength in bending, torsion or axial loading (calcium salt pastes, e.g. Norian); surgeons are therefore reluctant to use cements in pathological fractures in bone that may heal after the fracture has been stabilised. Indeed, many pathological fractures occurring as a result of malignant tissue (primary or secondary tumour) will heal when the fracture is fixed (usually with an intramedullary nail). This is particularly true when the tumour tissue is sensitive to treatment (radio- or chemotherapy). Non-resorbable or slowly resorbable cements would help to stabilise the fracture, especially if there is significant bone loss, but would then act as a significant impediment to natural bone healing. Introduction of a cement into such a fracture would normally be considered when there was little chance of healing, either because the bone loss is extensive (more than about 5 cm), or because the patient’s life expectancy is so short that fracture healing cannot be expected and during the palliative phase of treatment prompt stability is more important than long-term considerations.

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1.6 Extensive bone loss around the left hip prosthesis caused by gradual aseptic loosening.

1.9.2 Filling of cavitary defects as a temporary spacer after removal of infected joint replacement When joint replacements become infected (around 2% of all joint replacements in the United Kingdom), eradication of infection is difficult. The gradual loosening that occurs with infection is accompanied by the formation of a soft inflammatory membrane around the implant, and resorption of bone occurs. This often creates large cavitary defects and bone loss which can be extensive (Jeanrot et al. 1999), as shown in Fig. 1.6. One-stage exchange of implants with broad spectrum antibiotic cover and thorough debridement and lavage can produce reasonable results. However, better long-term results, with lower rates of recurrent infection have been achieved by two-stage revision arthroplasty. This involves removing the infected joint replacement and old cement, debridement of the cavity, sending samples from the operative site for culture, washout of the tissue and implantation of a cement spacer to fill the cavities until a new prosthesis can be inserted. Commonly, the spacer used is PMMA (as is used for joint replacement), loaded with gentamicin or another broad spectrum antibiotic. Moulds are now available to manufacture prosthesis-shaped cement spacers, partly to allow some movement, but also to optimise the space preservation for later prosthesis insertion. If the organism is known in advance, the antibiotic of choice can be added to the cement.

1.9.3 Osteomyelitis Chronic bone infection produces a cavitary defect and this may contain devitalised bone, which acts as a potent focus for continued infection (the sequestrum). New bone forms around the margins of the original bone, under the periosteum, which has been elevated off the surface of the

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cortical bone by the inflammatory process; this is called the involucrum. Many sinuses may be present through which the pus will drain. The first stage of treatment is removal of any dead, infected bone, including the sequestrum if present. The bone is then washed out, and any membranous soft tissue, and sinuses are excised. The remaining bone may be of doubtful strength and some surgeons would wish to stabilise the bone with cement. This may be cement with high structural strength, such as PMMA – also serving as a carrier for antibiotic – which the surgeon would intend to remove at a later stage, or a cement that would be intended to resorb later, but also acting as an antibiotic carrier (such as calcium sulphate preparations) (Jacobs et al. 1990; Henry and Galloway 1995). Chronic osteomyelitis will often require long-term antibiotics, at least 6 weeks of high-dose targeted treatment, and even then it may recur later.

1.9.4 Osteoporosis Osteoporosis will become a major healthcare challenge over the coming decades, with large numbers of elderly patients suffering fragility fractures. Although primary osteoporosis is as a consequence of the normal ageing process, secondary osteoporosis has many identifiable risk factors. The main risk factors for both primary and secondary forms of osteoporosis include family history, low body-mass index, smoking, alcohol abuse, previous longterm steroid therapy and previous fracture after the age of 55. There are several other factors with lower prognostic influence. There are currently many initiatives directed at dealing with this problem ranging from improved awareness, screening and prophylactic treatments to improve fixation of fractures once they have occurred. The commonest fractures occur in the vertebrae, the distal radius and the hip. Distal radial fractures do not usually require internal fixation, although augmentation with cement may be useful in some cases (Cassidy et al. 2003). Intracapsular hip fractures are usually treated by joint replacement in the United Kingdom because of poor results with internal fixation. Extracapsular hip fractures (the intertrochanteric and subtrochanteric fractures) are internally fixed with a nail or dynamic hip screw. Fixation of hip fractures and fixation of fractures elsewhere in osteoporotic patients is frequently complicated by poor purchase of fixation. Devices are evolving to deal with this, but currently, the bone may be strengthened by the insertion of cement into the medulla to augment the hold of screws (Harrington 1975; Bartucci et al. 1985). The cement may be inserted first, allowed to set and then drilled, or the fixation device may be inserted before the cement has set. PMMA is used most commonly, but other, hard-setting cements (such as Norian) have been used in this context.

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There has been recent interest in treating vertebral collapse with vertebroplasty or kyphoplasty (Bartucci et al. 1985; Rao and Singrakhia 2003). Vertebroplasty involves the expansion and restoration of height of a vertebral fracture using an injection of cement. Kyphoplasty involves percutaneous insertion of a balloon on a catheter inserted into the crushed vertebra using image intensifier guidance. High-pressure inflation of the balloon then expands the vertebra, and cement is inserted to preserve the vertebral height. Good results are being reported from this technique.

1.10

Other potentially useful applications of cements in bone

Various case reports have been published demonstrating the usefulness of bone cement in unusual indications. Avascular necrosis of the hip (Fig. 1.7) has not so far been treated with cement to repair the defect in the subchondral bone, but animal studies creating similar subchondral defects have been promising (Welch et al. 2002). One report described the treatment of a case of an intraosseous gouty tophus with calcium phosphate cement (Morino et al. 2007) and also hydatid disease of bone (Tomak et al. 2001). A clinical series demonstrated that Norian and Novabone can be used in cranioplasty with good effect (Elshahat et al. 2004) and also in craniofacial reconstruction (Lew et al. 1997; Turk and Parhiscar 2000; Kubo et al. 2002; Losee et al. 2003).

1.7 Subchondral collapse in avascular necrosis of the head of the femur.

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A review by Wang et al. (2002) of the use of calcium phosphate cement in 32 heterogeneous clinical cases shows the range of potential uses, including treatment of fractures, bone cyst, fibrous dysplasia and enchondroma, and also in iliac bone harvesting; it has also been used in cases of bone tuberculosis, which involved femur, iliac, tibia, humerus, phalanx, fibula, calcaneus, talus and acetabulum. Good results were obtained in all cases. Calcaneal fractures have been particularly difficult to treat, due to the high degree of comminution, and the fragility of fixation. Bioresorbable cements have been proposed as a potential solution to this problem, although this technique has not been adopted clinically (Thordarson et al. 1999). Comminuted fractures in the metaphysis of long bones may also benefit from cement augmentation, but again this has not been widely adopted by surgeons in clinical practice (Frankenburg et al. 1998). An exciting future for fracture and bone defect repair may arise from the discovery that, at least in animals, the combination of recombinant human bone morphogenetic protein 2 (rhBMP2) and a calcium phosphate cement carrier (alpha BSM) injected into fractures and implanted into bone defects, produces an enhanced bone healing response (Seeherman et al. 2004, 2006). Clinical trials are awaited.

1.11

Summary

Surgeons treating skeletal insufficiency utilise currently available cements to great advantage in metastatic and primary neoplasia, osteomyelitis, osteoporotic fracture fixation and for the filling of bone defects in various sites with a wide range of underlying pathology, including infected defects left by the removal of implants. Cements are used for structural strength and as carriers of antibiotics. The indications for the use of cements are broadening and cements will become more attractive to surgeons as cements of sufficient structural strength are formulated that are bioresorbable, encouraging normal bone to form as resorption occurs.

1.12

Acknowledgement

The authors would like to thank Masumi Tanaka, Queen Mary’s Hospital, Sidcup for assistance in preparing this chapter.

1.13

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melton, l. j., iii et al., ‘Vertebral fractures predict subsequent fractures,’ Osteoporos. Int. 10 (3): 214–221 (1999). mercadante, s., ‘Malignant bone pain: pathophysiology and treatment,’ Pain 69 (1–2): 1–18 (1997). morino, t. et al., ‘Intraosseous gouty tophus of the talus, treated by total curettage and calcium phosphate cement filling: a case report,’ Foot Ankle Int. 28 (1): 126– 128 (2007). mughal, z., ‘Rickets in childhood,’ Semin. Musculoskelet. Radiol. 6 (3): 183–190 (2002). pacifici, r., ‘Estrogen, cytokines, and pathogenesis of postmenopausal osteoporosis,’ J. Bone Miner. Res. 11 (8): 1043–1051 (1996). pandey, r. et al., ‘Histological and microbiological findings in non-infected and infected revision arthroplasty tissues. The OSIRIS Collaborative Study Group. Oxford Skeletal Infection Research and Intervention Service,’ Arch. Orthop. Trauma Surg. 120 (10): 570–574 (2000). parfitt, a. m., ‘The coupling of bone formation to bone resorption: a critical analysis of the concept and of its relevance to the pathogenesis of osteoporosis,’ Metab Bone Dis. Related Res. 4 (1): 1–6 (1982). pettifor, j. m., ‘Nutritional rickets: deficiency of vitamin D, calcium, or both?,’ Am. J. Clin. Nutr. 80 (6 Suppl): 1725S–1729S (2004). pfeilschifter, j. et al., ‘Characterization of the latent transforming growth factor beta complex in bone,’ J. Bone Miner. Res. 5 (1): 49–58 (1990a). pfeilschifter, j. et al., ‘Chemotactic response of osteoblastlike cells to transforming growth factor beta,’ J. Bone Miner. Res. 5 (8): 825–830 (1990b). rao, r. d. and m. d. singrakhia, ‘Painful osteoporotic vertebral fracture. Pathogenesis, evaluation, and roles of vertebroplasty and kyphoplasty in its management,’ J. Bone Joint Surg. Am. 85-A (10): 2010–2022 (2003). reid, i. r., ‘Glucocorticoid-induced osteoporosis,’ Baillières Best. Pract. Res. Clin. Endocrinol. Metab. 14 (2): 279–298 (2000). riggs, b. l. et al., ‘Rates of bone loss in the appendicular and axial skeletons of women. Evidence of substantial vertebral bone loss before menopause,’ J. Clin. Invest. 77 (5): 1487–1491 (1986). riggs, b. l. et al., ‘A unitary model for involutional osteoporosis: estrogen deficiency causes both type I and type II osteoporosis in postmenopausal women and contributes to bone loss in aging men,’ J. Bone Miner. Res. 13 (5): 763–773 (1998). roodman, g. d., ‘Advances in bone biology: the osteoclast,’ Endocr. Rev. 17 (4): 308–332 (1996). ross, p. d., ‘Clinical consequences of vertebral fractures,’ Am. J. Med. 103 (2A): 30S–42S (1997). rowe, p. s., ‘The wrickkened pathways of FGF23, MEPE and PHEX,’ Crit. Rev. Oral Biol. Med. 15 (5): 264–281 (2004). sambrook, p. and c. cooper, ‘Osteoporosis,’ Lancet 367 (9527): 2010–2018 (2006). sambrook, p. n. et al., ‘Genetics of osteoporosis,’ Br. J. Rheumatol. 33 (11): 1007–1011 (1994). schajowicz, f. et al., ‘The World Health Organization’s histologic classification of bone tumors. A commentary on the second edition,’ Cancer 75 (5): 1208–1214 (1995).

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schoenau, e. et al., ‘Influence of puberty on muscle area and cortical bone area of the forearm in boys and girls,’ J. Clin. Endocrinol. Metab. 85 (3): 1095–1098 (2000). seeherman, h. j. et al., ‘Recombinant human bone morphogenetic protein-2 delivered in an injectable calcium phosphate paste accelerates osteotomy-site healing in a nonhuman primate model,’ J. Bone Joint Surg. Am. 86-A (9): 1961–1972 (2004). seeherman, h. j. et al., ‘rhBMP-2 delivered in a calcium phosphate cement accelerates bridging of critical-sized defects in rabbit radii,’ J. Bone Joint Surg. Am. 88-A (7): 1553–1565 (2006). seeman, e., ‘From density to structure: growing up and growing old on the surfaces of bone,’ J. Bone Miner. Res. 12 (4): 509–521 (1997). seeman, e., ‘Pathogenesis of bone fragility in women and men,’ Lancet 359 (9320): 1841–1850 (2002). seeman, e. and p. d. delmas, ‘Bone quality – the material and structural basis of bone strength and fragility,’ N. Engl. J. Med. 354 (21): 2250–2261 (2006). sernbo, i. and o. johnell, ‘Consequences of a hip fracture: a prospective study over 1 year,’ Osteoporos. Int. 3 (3): 148–153 (1993). siris, e. s. et al., ‘Medical management of Paget’s disease of bone: indications for treatment and review of current therapies,’ J. Bone Miner. Res. 21 Suppl 2: 94–98 (2006). stein, e. and e. shane, ‘Secondary osteoporosis,’ Endocrinol. Metab. Clin. North Am. 32 (1): 115–134, vii (2003). stewart, a. et al., ‘Long-term fracture prediction by DXA and QUS: a 10-year prospective study,’ J. Bone Miner. Res. 21 (3): 413–418 (2006). sugawara, y. et al., ‘Rapid detection of human infections with fluorine-18 fluorodeoxyglucose and positron emission tomography: preliminary results,’ Eur. J. Nucl. Med. 25 (9): 1238–1243 (1998). thordarson, d. b. et al., ‘Superior compressive strength of a calcaneal fracture construct augmented with remodelable cancellous bone cement,’ J. Bone Joint Surg. Am. 81 (2): 239–246 (1999). tomak, y. et al., ‘Hydatid disease of the left femur: a case report,’ Bull. Hosp. Joint Dis. 60 (2): 89–93 (2001). turk, j. b. and a. parhiscar, ‘Bone source for craniomaxillofacial reconstruction,’ Facial. Plast. Surg. 16 (1): 7–14 (2000). van staa, t. p. et al., ‘Use of oral corticosteroids and risk of fractures,’ J. Bone Miner. Res. 15 (6): 993–1000 (2000). walker, d. g., ‘Osteopetrosis cured by temporary parabiosis,’ Science 180 (88): 875 (1973). wang, w. b. et al., ‘[Primary clinical study on self-setting calcium phosphate cement in bone defect repair of extremities],’ Zhongguo Xiu. Fu Chong. Jian. Wai Ke. Za Zhi. 16 (2): 100–102 (2002). weitzmann, m. n. and r. pacifici, ‘The role of T lymphocytes in bone metabolism,’ Immunol. Rev. 208: 154–168 (2005). weitzmann, m. n. and r. pacifici, ‘Estrogen deficiency and bone loss: an inflammatory tale,’ J. Clin. Invest. 116 (5): 1186–1194 (2006). welch, r. d. et al., ‘Subchondral defects in caprine femora augmented with in situ setting hydroxyapatite cement, polymethylmethacrylate, or autogenous bone

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graft: biomechanical and histomorphological analysis after two years,’ J. Orthop. Res. 20 (3): 464–472 (2002). wharton, b. and n. bishop, ‘Rickets,’ Lancet 362 (9393): 1389–1400 (2003). zaidi, m., ‘ “Calcium receptors” on eukaryotic cells with special reference to the osteoclast,’ Biosci. Rep. 10 (6): 493–507 (1990). zimmerli, w. et al., ‘Pathogenesis of foreign body infection: description and characteristics of an animal model,’ J. Infect. Dis. 146 (4): 487–497 (1982).

2 Hip replacements B. M. W R O B L E W S K I, P. D. S I N E Y and P. A. F L E M I N G, The John Charnley Research Institute, UK

Abstract: The Charnley low frictional torque arthroplasty (LFA) has now reached 45 years of continuous clinical success. Pain relief and excellent range of movements have been achieved and maintained over the follow-up time, but consumer demands for new products have stimulated the introduction of new materials and methods. The cemented Charnley LFA has withstood the test of time. The method offers a ‘customised’ prosthesis in every case but places the onus on the understanding and skill of the surgeon. Long-term results are results in young patients. The patient’s activity level is not a characteristic of any hip design or any method of component fixation, but a reflection of the patient selection. The immediate clinical success of hip replacement and the pressure to satisfy the increasing demand, especially in young patients, have highlighted the failure to understand the need for regular follow-up and the provision of ‘after-sales service’. The timing, indication and detailed operative findings at revision surgery have become important issues. Delays in revision surgery lead to progressive loss of bone and more complex technical problems. Regular follow-up after hip replacement with good-quality radiographs is essential. Cemented total hip arthroplasty continues to be an excellent method of treatment but, for young patients in particular, it is merely the beginning of the treatment. Key words: Charnley, long-term results, wear and loosening of the cup, strain shielding of the proximal femur, young patients.

2.1

Introduction

The success of total hip arthroplasty (THA) as a method of treating hip joints destroyed by arthritis has resulted in an increasing demand from an ageing population and has also extended the indications for this procedure. Consumer demand for new products stimulated the introduction of new materials and methods. Anecdotal, single-case, successes attract younger and less disabled patients. Immediate clinical success, short-term planning and pressure to satisfy the increasing demand, do not usually take into consideration the need for regular follow-up: the provision of ‘after-sales service’. 41

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The availability of a successful ‘replacement’ procedure has also had a negative effect on the systematic study of underlying hip pathologies. It is not readily appreciated that clinical success is primarily due to pain relief. Pain relief is due to the natural, symptomatic joint being replaced with a mechanical neuropathic spacer; the whole construct functioning within a foreign body bursa. Wroblewski et al. (1999), in their long-term follow-up paper, stated that ‘clinical results do not reflect the mechanical state of the arthroplasty’ and that good-quality radiographs offer more information. Wroblewski et al. (2007a) also state that the activity level achieved after the operation is not a characteristic of a particular design, nor the method of component fixation, provided the fixation is secure; but it is a reflection of patient selection for the operation, patients with multiple disabilities never feature as anecdotal, single-case successes.

2.2

General principles

The method of treatment can be suitably reviewed under three main headings: design, materials and surgical technique.

2.2.1 Design The aim must be a set of components that allow a range of movements comparable with the natural joint, while ensuring stability without constraint. Any design must also take advantage of the best possible arrangements and combinations with respect to friction, frictional torque, wear and load transfer from the implant to the skeleton.

2.2.2 Materials Any combination of materials must be compatible with respect to each other and biocompatible with the tissues, not only in the solid but especially in their particulate form. They must also be fatigue and wear resistant.

2.2.3 Surgical technique Exposure must give an unimpeded, circumferential view of the acetabulum and the access to the medullary canal, in the area of the piriformis fossa, while preserving the integrity of the neighbouring structures as well as the all-important abductor muscle mass. Exposure of good-quality cancellous bone, pulse lavage, drying, containment, pressurisation of cement and correct timing of component insertion, are integral parts of the cementing technique.

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2.3

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Bone cements

The characteristics of the various bone cements are well documented by Kühn (2000). Two aspects are of clinical importance: shrinkage on polymerisation and release of antibiotics. Shrinkage on polymerisation is the characteristic of the monomer. Our findings suggest that shrinkage occurs after setting/polymerisation has apparently taken place. The changes observed indicate that shrinkage would be away from both bone and the prosthesis. Under operative conditions the effect of the components acting as a ‘heat sink’ and intrusion into cancellous bone resisting the shrinkage, are yet to be quantified. Measured volume reduction is 25% (unpublished data), close to the 21% reported by Kühn (2000). Elution from polymerised cement is purely a surface phenomenon (Fig. 2.1). Wroblewski (1977), in his experimental evaluation, states ‘the volume released is proportional to the surface area of the cement mass’. Elution of gentamicin is neither continuous nor complete. Wroblewski et al. (1986) also noted that ‘on average 22% of the antibiotic is released within days of the operation’. Some of the remaining 78% will be released if exposure of the new cement surface occurs, as in cases of component loosening or cement fragmentation at a revision. This may affect samples taken for bacteriology.

2.1 A mixture of bone cement and gentian violet immersed in commercial bleach for 2 years. Note the thin outer bleached layer, some extensions of the process into the cement irregularities and the persistence of the dye within the cement mass.

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2.4

Fixation of components with cement

Charnley (1963) in his lecture at the Société Internationale de Chirurgie Orthopédique et de Traumatologie (SICOT) said ‘The crux of the operation lies in the use of cement. By means of cement the load of the body weight is distributed over a large area of bone.’ This statement summarises succinctly all aspects of the technique. Charnley (1964) also emphasised that ‘Acrylic cement does not adhere to bone like glue, it merely forms an accurate cast of the interior of the bone so that load is transmitted evenly over all parts of the interface between cement and the cancellous bone’. This is the only method that offers immediately a ‘customised’ prosthesis in every case but also places the onus on the understanding and skill of the surgeon to achieve that aim. That skill is acquired by practice and is not a commodity that can be sold at a profit. This would probably explain the proliferation of ‘cementless’ designs. The terminology does not offer an indication of the method of component fixation except by stating what it is not, i.e. not with cement. It also hints that use of cement may be detrimental. Ultimately such designs become saleable commodities and take the responsibility for the quality of component fixation away from the surgeon and place it on the patient’s skeleton.

2.5

Long-term results

The Charnley low frictional torque arthroplasty (LFA) has now reached 45 years of continuous clinical application and must, therefore, be accepted as the gold standard (Wroblewski and Siney 1993).

2.5.1 Clinical results Using the Charnley LFA, pain relief is achieved and maintained; 96% of patients have either a totally pain-free hip or have no more than an occasional discomfort. Activity level, although improved and maintained, clearly depends on the underlying pathology and any factors that may become significant with time. Excellent, functional range of movements is achieved in 78% of patients and is maintained over the follow-up time.

2.5.2 Survivorship analysis Survivorship analysis used routinely in many hip registers continues to be the standard method of presenting long-term results with revision as the end point; each revision documented under a single indication (Soderman 2000). Our practice is to document all findings at revision (apart from infection which is a single finding). This gives a higher number of findings than

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Table 2.1 Survivorship analysis of 22 066 LFAs with a follow-up to 36 years

Reason for revision

Percentage success rate at 36 years’ follow-up

95% confidence limits

Infection Dislocation Fractured stem Loose stem Loose cup

91.4 98.0 88.6 72.5 48.6

100–67.9 100–85.9 100–62.4 100–39.2 79.1–18.0

hips revised and offers more valuable information (Wroblewski et al. 2007b). Survivorship analysis of the Charnley LFA, with a follow-up to 36 years is shown in Table 2.1.

2.6

Long-term problems

Two main long-term problems have been identified and addressed: (a) wear and loosening of the ultra-high molecular weight polyethylene (UHMWPE) cup and (b) strain shielding of the proximal femur. Charnley and Halley (1975) anticipated wear and loosening of the UHMWPE cup. They stated that ‘more than 5 mm wear might cause impingement of the neck of the prosthesis against the inner rim of the socket and cause loosening of the cement–bone bond in the acetabulum.’ The exponential correlation between wear and cup loosening has been documented in more recent studies (Wroblewski 1985a, Worblewski and Siney 1992, Worblewski et al. 2002). Issac et al. (1992) and Goldsmith et al. (2001), in their clinical and tribological studies, have identified factors affecting wear. The sequence of events of progressive penetration of the cup is restriction of movements, impingement of the neck of the stem on the rim of the cup and cup loosening. Reducing the diameter of the neck of the Charnley stem from 12.5 mm to 10 mm has delayed the impingement by the equivalent of 2 mm cup penetration. Clinical experience over the past 23 years (Wroblewski et al. 2004) supports the findings of the original theoretical model that used acrylic casts and shadowgraph techniques to measure real and radiographic wear (Wroblewski 1985b). Alumina ceramic/cross-linked polyethylene combinations, in the Charnley LFA, have now reached 20 years of clinical application (Wroblewski et al. 2005). The mean penetration rate is 0.02 mm/year with total penetration to date not exceeding 0.41 mm. Strain shielding of the proximal femur has been identified as being due to distal stem support not allowing proximal load transfer. The continuing developmental design of the stem, the C-Stem (DePuy International,

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Orthopaedic bone cements

Leeds, UK) takes advantage of the common engineering principle: that of male (stem) and female (cement) tapers engaging under load. A number of factors essential for the concept to function have been detailed (Wroblewski et al. 2001). These are: a polished stem with a continuous triple taper; absence of distal stem support; and attention to the anatomical calcar. Clinical experience over the past 13 years has shown 100% survivorship – the end-point being stem loosening or revision – with 22% showing improvement in the radiographic appearance of the quality of the proximal femur.

2.7

Future trends

The cemented Charnley LFA has withstood the test of time. As with any THA, clinical results may not reflect the mechanical state, hence follow-up by good-quality radiographs is essential. Postal or telephone methods of collecting information are of limited value. Progressive loosening is an indication for revision, irrespective of the clinical result. This aspect must be understood and accepted by the patient, the surgeon and the system funding such a method of treatment. This must be agreed before the operation. Delays in revision surgery lead to progressive loss of bone stock and more complex technical problems to be tackled at revision. Long-term results are results in young patients. Increasing follow-up identifies ever-younger patients at the time of the operation. Now approaching 40 years’ follow-up, the mean age of patients at the time of the primary operation was 35 years. These are the results that have already been achieved with the Charnley LFA. Any newly proposed method or materials claiming to be suitable for young patients has the impossible task of providing data to support the statement that it will: accept patients with a mean age of 35 years and show a functioning arthroplasty when the patients reach the age of 75 years. This will not be achieved by a single generation of surgeons. Even then, continuity of design, materials and technique must be maintained. THA in general, and the Charnley LFA in particular, continue to be an excellent method of treatment, but for a young patient it merely heralds the beginning of treatment.

2.8

References

charnley j. 1963. Low friction arthroplasty of the hip in rheumatoid arthritis. SICOT Congress, Vienna, pp. 168–70. charnley, j. 1964. The bonding of prostheses to bone by cement. J Bone Joint Surg Br, 46, 518–29. charnley, j. & halley, d. k. 1975. Rate of wear in total hip replacement. Clin Orthop Relat Res, Oct (112), 170–9.

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goldsmith, a. a., dowson, d., wroblewski, b. m., siney, p. d., fleming, p. a. & lane, j. m. 2001. The effect of activity levels of total hip arthroplasty patients on socket penetration. J Arthroplasty, 16 (5), 620–7. isaac, g. h., wroblewski, b. m., atkinson, j. r. & dowson, d. 1992. A tribological study of retrieved hip prostheses. Clin Orthop Relat Res, Mar (276), 115–25. kühn k. d. 2000. Bone Cements. Up to Date Comparison of Physical and Chemical Properties of Commercial Materials. Springer, Berlin. soderman p. 2000. On the validity of the results from the Swedish National Total Hip Arthroplasty Register. Acta Orthop Scand, 71 (suppl 296), 4–33. wroblewski, b. m. 1977. Leaching out from acrylic bone cement. Experimental evaluation. Clin Orthop Relat Res, May (124), 311–2. wroblewski, b. m. 1985a. Charnley low-friction arthroplasty in patients under the age of 40 years. In Sevastik J. & Goldie, I. (eds), The Young Patient with Degenerative Hip Disease. Almquist and Wiksell, Stockholm, pp. 197–201. wroblewski, b. m. 1985b. Direction and rate of socket wear in Charnley low-friction arthroplasty. J Bone Joint Surg Br, 67 (5), 757–61. wroblewski, b. m., esser, m. & srigley, d. w. 1986. Release of gentamicin from bone cement. An ex-vivo study. Acta Orthop Scand, 57 (5), 413–4. wroblewski, b. m. & siney, p. d. 1992. Charnley low-friction arthroplasty in the young patient. Clin Orthop Relat Res, Dec (285), 45–7. wroblewski, b. m. & siney, p. d. 1993 Charnley low-friction arthroplasty of the hip. Long-term results. Clin Orthop Relat Res, Jul (292), 191–201. wroblewski, b. m., fleming, p. a. & siney, p. d. 1999. Charnley low-frictional torque arthroplasty of the hip. 20-to-30 year results. J Bone Joint Surg Br, 81 (3), 427–30. wroblewski, b. m., siney, p. d. & fleming, p. a. 2001. Triple taper polished cemented stem in total hip arthroplasty: rationale for the design, surgical technique, and 7 years of clinical experience. J Arthroplasty, 16 (Suppl 1), 37–41. wroblewski, b. m., siney, p. d. & fleming, p. a. 2002. Charnley low-frictional torque arthroplasty in patients under the age of 51 years. Follow-up to 33 years. J Bone Joint Surg Br, 84 (4), 540–3. wroblewski, b. m., siney, p. d. & fleming, p. a. 2004. Reduced diameter of the neck and its effect on the incidence of aseptic cup loosening in the Charnley LFA. J Bone Joint Surg Br, 87 (Suppl 1), 43. wroblewski, b. m., siney, p. d. & fleming, p. a. 2005. Low-friction arthroplasty of the hip using alumina ceramic and cross-linked polyethylene. A 17-year follow-up report. J Bone Joint Surg Br, 87 (9), 1220–1. wroblewski, b. m., siney, p. d. & fleming, p. a. 2007a. Charnley low-frictional torque arthroplasty in young rheumatoid and juvenile rheumatoid arthritis: 292 hips followed for an average of 15 years. Acta Orthop, 78 (2), 206–10. wroblewski, b. m., siney, p. d. & fleming, p. a. 2007b. Charnley low-friction arthroplasty: Survival patterns to 38 years. J Bone Joint Surg Br, 89, 1015–8.

3 Knee replacements H. PA N D I T and B. H. VA N D U R E N, Nuffield Orthopaedic Centre, UK

Abstract: This chapter aims to provide the reader with an overview of cemented total knee replacement (TKR). To understand the various aspects of TKR, such as the indications for surgery, design rationale and outcome, one needs to have a background knowledge of the anatomy and biomechanics of the knee joint. The first two sections are devoted to these aspects and will cover the management of a patient presenting with an arthritic knee. Other treatment modalities are available for the treatment of osteoarthritis and these will be explained in brief. This is followed by a description of the development of TKR before presenting the cemented TKR. Key words: osteoarthritis, total knee replacement, unicompartmental knee replacement, cemented total knee replacement.

3.1

Relevant anatomy of the knee joint

The knee is the largest synovial joint in the human body and is the articulation between the principal bones of the lower extremity (femur and tibia) and the patella. The patella is an oval-shaped sesamoid bone that is part of the quadriceps and patella tendon; it articulates anteriorly on the femur and as part of the patello-femoral joint. The lower end of the femur features two prominences called the lateral and medial condyles. These condyles articulate with the patella and the articular surface of the tibia and together form the tibio-femoral joint. The medial condyle is the inside prominence, resting on the medial meniscus, while the outside, or lateral, condyle rests on the lateral meniscus. The contact/articular surfaces of the bone are cushioned from the stresses placed on them, by a layer of hyaline cartilage (Fig. 3.1). The joint is stabilised by four major ligaments. The ligaments ensure the mechanical stability of the joint, guide joint motion and prevent excessive laxity. The medial collateral ligament runs along the medial side of the femur down to the tibia and the lateral collateral ligament runs along the lateral side of the femur to the head of the fibula to provide stability in the coronal plane. The anterior and posterior cruciate ligaments (ACL, PCL) connect the inner surfaces of the distal femur with the top of the tibia 48

Knee replacements

49

Quadriceps tendon Femur Patella Lateral condyle Articular cartilage Lateral meniscus Lateral collateral ligament Fibula

Medial collateral ligament Posterior cruciate ligament Anterior cruciate ligament Patella tendon (ligament)

Tibia

3.1 Line drawing of human knee joint.

providing restraint to the relative anterior–posterior movement of the tibia and femur. The joint is enclosed by the joint capsule; a tough, protective tissue lined with a synovial membrane. The membrane secretes a viscous fluid called synovium, which not only provides the joint with nutrients, but also introduces a form of lubrication between the articular surfaces.

3.2

Conditions causing knee arthritis

Symptomatic end-stage arthritis of the knee is the most common indication for total knee replacement (TKR). Irregularity and loss of articular cartilage can be caused by a variety of conditions, the most common being primary osteoarthritis, which is discussed in detail in the following section. In addition to primary osteoarthritis, intra-articular fractures of the knee joint (traumatic), inflammatory arthropathy (rheumatoid arthritis and its variants, gout, synovial chondromatosis, haemophilia, etc.) and avascular necrosis cause secondary arthritis requiring treatment.

3.2.1 Pathophysiology of osteoarthritis Osteoarthritis is the most common form of arthritis. Many other terms are also used to describe osteoarthritis, including ‘osteoarthrosis’, ‘arthrosis’ and ‘degenerative joint disease’. The knee and the hip joint are the joints

50

Orthopaedic bone cements

most commonly affected by osteoarthritis. When a joint develops osteoarthritis, the cartilage gradually roughens and becomes thinner. The surrounding bone reacts to this by growing thicker. The bone around the edge of the joint grows outwards (osteophytes or bony spurs). This bone growth can affect both the femur and the tibia, as well as the patella. The synovium may produce extra fluid, making the joint swell. The capsule and the ligaments slowly thicken and shrink and the muscle groups that move the knee joint (quadriceps and hamstrings) gradually weaken and become thin or wasted. This can make the knee joint unstable so that it ‘gives way’ when you put weight on it. Osteoarthritis is not the simple ‘wear and tear’ of a joint, although the term is commonly used to try to explain the phenomenon to the patient. The osteoarthritic joint is metabolically, biochemically and structurally different from an ageing joint. Primary osteoarthritis is a slow process that develops over many years. Healthy articular cartilage is made up of type II collagen fibres, which contain hydrophilic proteoglycans, which imbibe fluid and swell against a restricting net of collagen. This provides a sturdy but flexible and durable material, which spreads loads evenly. In osteoarthritis, the disruption of the collagen releases the proteoglycan leading to swelling of the articular surface, fragmentation and debris formation, followed by an inflammatory reaction from the synovium. The cartilage becomes so thin that it no longer covers the thickened bone ends. The bone ends then rub against each other, and start to wear away. One can feel and hear the crepitus produced by the bone ends rubbing against each other. The loss of cartilage, the wearing of bone and the bony overgrowth at the edges, all combine to change the shape and alignment of the joint. This forces the bones out of their normal positions and causes deformity. The most common deformity seen is a varus deformity and is caused by predominant affection of the medial or the inside compartment of the knee (Fig. 3.2). Sometimes the arthritis affects the lateral compartment, although this is far less common (Fig. 3.3). Many factors increase the risk of osteoarthritis developing in the knee joint; such as obesity, age, repeated trauma and injury to the menisci or the ligaments of the knee. The risk increases with age; however, osteoarthritis of the knee joint is not a problem in all elderly people but can often run in families. Osteoarthritis of the knee is twice as common in women as in men. It mainly occurs in women who are over the age of 50, but there is no strong evidence that it is directly linked to the menopause. Osteoarthritis of the knee is also more common in some racial groups than others; for example, it is more common in Afro-Caribbean and Asian people than in Caucasian people. Normal use does not routinely lead to osteoarthritis, and neither does exercise (including running) unless it is excessive. However, injuries to the

Knee replacements

Left

3.2 Radiograph showing medial compartment osteoarthritis.

Right Valgus

3.3 Radiograph showing lateral compartment osteoarthritis.

51

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Orthopaedic bone cements

knee joint often lead to osteoarthritis in later life. A common cause is a tear of the meniscal cartilage or ligaments after a twisting injury. This is a common injury in footballers, who may face additional risk. The damaged cartilage can lead to osteoarthritis in later life, and we now know that the operation to remove the torn cartilage (meniscectomy) substantially increases the risk of osteoarthritis developing after a number of years.1

3.3

Clinical and radiological assessment of an arthritic knee

Assessment begins with a clinical history, careful examination and an X-ray. The initial clinical assessment aims to answer three questions. 1 2 3

Is the knee the true source of the patient’s symptoms? Is the disease of the knee truly osteoarthritis? How disabled is the patient due to osteoarthritis?

Osteoarthritis of the knee affects people in different ways. Some patients have the disease in only one knee, others in both knees. For some patients pain is the main complaint, while others find their main problem is the difficulty in walking. Some patients may notice little change in their condition over the years, while in others the osteoarthritis progressively gets worse. Pain in the knee joint can be referred from elsewhere, i.e. the hip or the back, and it is necessary to rule out these possibilities. It is important to note the patient’s age, gender and past medical history; in particular, trauma to the knee, hip and back. In addition to assessing the nature and severity of the pain experienced, it is important to assess the functional limitations caused by the osteoarthritis. When taking a clinical history, attention must be given to the patient’s ability to walk, negotiate stairs, perform daily activities and continue working, as well as how many painkillers he or she consumes on a daily basis. An examination of the knee joint should assess the presence or absence of effusion, range of movement, stability, muscle atrophy, joint line tenderness, gait and patello-femoral irritability. Further assessment of the patient’s general health is also important when making decisions on available treatment options. In advanced osteoarthritis, a weight-bearing antero-posterior and lateral radiograph suffice to show the extent of the disease (Fig. 3.4). In particular cases, special views such as a skyline view may be used to assess the patellofemoral joint (Fig. 3.5). Stress views to assess the integrity of cartilage in a relatively well-preserved tibio-femoral compartment or magnetic resonance imaging (MRI) to assess soft tissue pathologies and early osteoarthritis are sometimes used. Blood investigations may also be required to exclude inflammatory arthropathy, e.g. rheumatoid arthritis.

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100 mm measuring tool standing

3.4 Radiograph showing osteoarthritis affecting both compartments.

3.5 Skyline view showing patello-femoral arthritis.

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Orthopaedic bone cements

3.4

Treatment options for osteoarthritis

Currently, there is no definitive cure for osteoarthritis. Various treatment options are available, depending upon the severity of the symptoms and extent of the disease. The treatment options are aimed at relieving discomfort and pain, reducing stiffness and, if possible, reducing any further damage to the joint. A change in lifestyle (including job), weight loss and use of a support around the knee joint are usually the initial treatment options. Although no drugs have been shown to halt the process of osteoarthritis, several painkillers (such as paracetamol) and anti-inflammatory drugs (such as ibuprofen) and creams to rub into the knee can help ease pain and reduce stiffness. Anti-inflammatory drugs (non-steroidal anti-inflammatory drugs or NSAIDs) reduce the inflammation but can cause stomach ulcers. New NSAIDs called COX-2s are less likely to cause stomach problems but have been linked with increased risks of heart attack and stroke, so they are not suitable for people who either have had in the past, or currently have uncontrolled high blood pressure. All NSAIDs may cause other side effects such as rashes, headaches and wheeziness. Sometimes an intra-articular injection of steroids may help, although the effect is usually temporary. Injections of hyaluronic acid derivatives may also help by supplementing the joint’s natural synovial fluid. Finally, various surgical options are available, including arthroscopic washout and debridement, high tibial osteotomy (HTO), unicompartmental knee replacement (UKR) and TKR. Arthroscopy of the knee allows a detailed assessment of the extent of pathology and can help relieve symptoms by removing loose or unstable pieces of cartilage, bone or meniscus from the knee joint. Localised chondral defects can be treated by newer treatment options like microfracture, abrasion arthroplasty, autologous chondrocyte implantation (ACI) and osteochondral autograft transplantation (OAT). Although successful in selected cases in expert hands, in general these treatment options are only applicable in patients with osteoarthritis limited to a localised area(s). Osteotomy around the knee for arthritis has a long history with many proponents. However, it has fallen into some disrepute recently due to the significant success rates of knee replacement and greater expectations of the patients. HTO to correct a varus deformity is most valuable in young, heavy male patients employed in a manual job. The results of osteotomy are rather unpredictable and there are significant complications including non-union, infection and common peroneal nerve palsy.2 In cases where the diseased or damaged surfaces are limited to a single compartment (medial or lateral), a UKR can be used (Figs 3.6 and 3.7). Early UKR designs had poor results which are attributed to poor technique and instrumentation.3–5 This meant that the procedure was not frequently

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Left

3.6 Radiograph showing medial unicompartmental knee replacement.

Left

3.7 Lateral radiograph of the knee showing unicompartmental knee replacement.

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performed. Recently, numerous reports of UKR showing improved longterm results have been published.6–11 These improvements are attributed to improved surgical techniques and instruments. Consequently the use of UKR has become increasingly popular3 in the past decade as it has advantages over TKR. Namely, the preservation of both cruciate ligaments (anterior and posterior), thereby ensuring normal knee kinematics and better proprioception, improved range of movement, less peri-operative blood loss and early recovery. It has the additional advantage of preserving the unaffected knee compartments. Its role is particularly well established in medial compartment osteoarthritis, the most common site of involvement. The ligaments around the knee joint should be functionally intact if UKR is to be considered as a treatment option. The Oxford UKR is one of the most popular implants and has a 95% survivorship in expert hands.9,12 Between one in three and one in four patients presenting with end-stage osteoarthritis are suitable for UKR. Although in many cases UKR can be a definitive treatment for knee arthritis, in some cases the implant may eventually fail either due to progression of arthritis in the lateral compartment or aseptic loosening of the components. In such cases, the implant can be easily removed and exchanged for a primary TKR. In cases of end-stage arthritis not responding to other treatment options and when the disease involves more than one compartment, TKR is commonly performed. TKR is an elective procedure. The risks and outcomes vary, and it is essential that patients are informed of the likely consequences of the surgery in terms that are specific to them. Every patient’s goals and expectations (i.e. hopes and fears) should be established before surgery to determine whether their goals are attainable and their expectations are realistic. Any discrepancies between the patient’s expectations and the likely surgical outcome should be discussed in detail before surgery. The indications for TKR surgery are discussed in greater detail in the following section.

3.5

Indicators for total knee replacement

Based on existing research, TKR is considered a safe and cost-effective treatment. Overall, TKR has been shown to be a very successful, relatively low-risk therapy despite variations in patient health status and characteristics, type of prosthesis implanted, orthopaedic surgeons and surgical facilities. Each year, approximately 300 000 TKR surgeries are performed in the United States13 and around 45 000 in the United Kingdom for end-stage arthritis of the knee joint.14 Primary TKR is most commonly performed for knee joint failure caused by osteoarthritis; other indications include rheumatoid arthritis, juvenile rheumatoid arthritis, osteonecrosis and other types of inflammatory

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arthritis. TKR aims to relieve pain and improve knee function. Candidates for elective TKR should have radiographic evidence of joint damage, moderate-to-severe persistent pain that is not adequately relieved by an extended course of non-surgical management and clinically significant functional limitation resulting in diminished quality of life. In patients with rheumatoid arthritis and other inflammatory arthropathies, additional disease-specific therapies may be needed to achieve control of disease activity before proceeding with a surgical procedure. In the past, patients between 60 and 75 years of age were considered to be the best candidates for TKR. Over the past two decades, however, the age range has been broadened. On the one hand, more elderly patients (e.g. octogenarians and beyond), many of whom have a higher number of comorbidities, are offered a knee replacement and, on the other hand, younger patients, whose implants may be exposed to greater mechanical stresses (due to higher levels of physical activity) over an extended time period are offered a knee replacement. This is partly due to the increasing confidence of the surgeon in his abilities to perform an operation that is safe and will give lasting relief from debilitating symptoms. There are few absolute contraindications for TKR, e.g. the presence of active local or systemic infection and other medical conditions that substantially increase the risk of serious peri-operative complications. Obesity is not a contraindication to TKR; however, there may be an increased risk of delayed wound healing and peri-operative infection in obese patients. Additionally, general health risks are more common in clinically obese patients. Severe peripheral vascular disease and some neurological impairment are also considered contraindications to TKR.

3.6

Evolution of knee replacements

The concept of improving knee joint function by modifying the articular surfaces is not new. Various structures, such as pig bladder, nylon, fascia lata, pre-patella bursa and cellophane have been used to reconstruct the articular surface of a joint with interposition of soft tissues. The results were disappointing in general. In 1860, Ferguson15 resected an entire knee joint (excision arthroplasty), which resulted in mobility of the newly created subchondral surfaces. When more bone was removed, the patient enjoyed better knee motion but lacked the necessary stability of the knee, whereas limited bone resection resulted in spontaneous fusion. The first artificial knee implants were tried in the 1940s15 as moulds (both metallic and plastic materials were used) fitted to the femoral condyles following similar designs for the hip. In the next decade, tibial articular surface replacement was also attempted, but both designs had problems with loosening and persistent pain. Combined femoral and tibial articular surface

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replacements appeared in the 1950s as simple hinged prostheses.16 These implants failed to account for the complexities of knee motion and consequently had high failure rates from aseptic loosening. They were also associated with unacceptably high rates of postoperative infection. In 1971, Gunston17 importantly recognised that the knee does not rotate on a single axis like a hinge, but rather the femoral condyles roll and glide on the tibia with multiple instantaneous centres of rotation. His polycentric knee replacement had early success with its improved kinematics over hinged implants but in the end was unsuccessful because of inadequate fixation of the prosthesis to the bone. Of the many designs investigated in the 1970s, the total condylar design proved to be the most successful.18–20 The total condylar prosthesis was a true total replacement of the knee in that the patello-femoral joint was replaced as well as the femoral tibial compartment. Designed from a functional perspective, the inherent geometry of the prosthesis was intended to substitute the anatomic function of the cruciate ligaments, native articular geometry and menisci. The femoral component made of cobalt chromium alloy contained a symmetrically grooved anterior flange that separated posteriorly into two symmetric condyles, each of decreasing radius posteriorly with a symmetric convex curve feature in the coronal plane. The tibial component was made up of highdensity polyethylene in one piece with two separate bi-concave tibial plateaus that are articulated precisely with the femoral condyles in extension, thus permitting no rotation in this position. In flexion the fit ceased to be exact and rotation and guiding motions were possible. The symmetric tibial plateaus were separated by an inter-condylar eminence designed to prevent sideways sliding movement. The under surface of the component had a central fixation peg. The patella component was made up of high-density polyethylene and was dome shaped on its articular surface. Based on the initial total condylar design concept, many improvements were made in an attempt to find an optimal balance between joint stability and the stresses transmitted through the joint21 (Figs 3.8 to 3.10). These improvements have led to much debate over the design of TKRs. Current debates in TKR revolve around cruciate (PCL) excision, retention or substitution, fixed bearing or mobile bearing knee replacement, patella resurfacing and UKR versus TKR. These design rationales are addressed in greater detail in the following section.

3.7

Implant design rationales

3.7.1 Patella resurfacing The need for patella resurfacing is a controversial issue. In general, there is consensus that the patella should be resurfaced in patients undergoing TKR

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3.8 Image of a total knee replacement implant.

3.9 Implanted total knee replacement: antero-posterior view.

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3.10 Implanted total knee replacement: lateral view.

for inflammatory arthropathy and in patients in whom the patella is worn and/or does not track normally. Patella abnormalities are often assumed to result in anterior knee pain in TKR patients. However, the exact aetiology of unexplained anterior knee pain in patients undergoing TKR remains unclear and patella resurfacing depends upon the surgeon’s preference. The general consensus remains that the advantages of patella resurfacing are nullified by the associated complications of patella resurfacing namely patello-femoral problems, patella fracture and extensor mechanism issues. Burnett et al.22 in 2004 published results of a randomised controlled clinical trial at a minimum of 10 years’ follow-up in which patella resurfacing was compared with non-resurfacing in TKR. In this study, 100 knees (90 patients) with osteoarthritis were enrolled in a prospective randomised clinical trial using a posterior-cruciate-retaining TKR. Patients were randomised to receive resurfacing or retention of the native patella. Evaluations were done preoperatively and yearly, up to a minimum of 10 years. Diseasespecific (Knee Society clinical rating score) and functional (stair climbing, flexion/extension, patellar examination) outcomes were measured. Patient satisfaction, anterior knee pain and patello-femoral questionnaires were

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completed. Intra-operative grading of the articular cartilage was carried out. No patients were lost to follow-up; 45 patients remained alive. Nine revisions (9 of 90 knees; 10%) were carried out: in 7 patients in the nonresurfaced group (15% of knees) and in 2 patients in the resurfaced group (5% of knees). No significant difference was found between the groups regarding revision rates, Knee Society clinical rating scores, functional outcome, and patient satisfaction, anterior knee pain, patello-femoral and radiographic outcomes. Intra-operative cartilage quality was not a predictor of outcome. This study is currently the longest follow-up of a randomised; controlled, clinical trial that examines patella resurfacing in TKR. The results showed no significant difference between the groups for all outcome measures at a minimum of 10 years’ follow-up.

3.7.2 Fixed versus mobile bearing TKRs using well-designed, fixed-bearing prostheses have produced good long-term results. However, problems with polyethylene wear, osteolysis and fixation failure have occurred with some fixed-bearing designs. Mobilebearing TKR was designed to provide dual-surface articulation at both the upper and lower surfaces of the polyethylene insert. These designs offer the advantage of conforming geometry resulting in a reduction of contact stresses in the polyethylene, which may reduce wear. Goodfellow and O’Connor23 suggested that a mobile bearing design should reduce bone– prosthesis stress at the tibial surface. Despite the above-stated advantages, mobile bearing designs also introduce risks specific to the mobile bearing; i.e. bearing dislocation and increased backside wear. However, long-term studies of mobile- and fixed-bearing knees have shown no difference in the rate of wear and osteolysis. A recent Cochrane review also found no indication that either the mobile- or fixed-bearing TKR results in better functional performance of the patients.24

3.7.3 Cruciate excision, retention or substitution The cruciate ligaments provide static anterior and posterior stability to the knee, but more importantly in addition they also impose certain movements on the joint surfaces relating to one another. The ACL is often absent in arthritic knees and is not thought to be of much consequence in TKR although its role is pivotal in the success of UKR. On the other hand, the PCL, although often attenuated in arthritic knees, is usually present. It is also considered the collateral ligament for the medial compartment of the knee. The PCL causes the femoral condyles to glide and role back on the tibial plateau as the knee is flexed. In a normal knee the shape of the plateau does not restrain this motion, and laxity of the meniscal attachments allows

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the menisci to move posteriorly with the femur. This femoral roll back is crucial in prosthetic design. If the cruciates are excised, a more conforming tibial polyethylene component can be used to provide some degree of anterior and posterior stability. However, without the function of the PCL, femoral roll back will not occur and this theoretically will limit the ultimate knee flexion that can be achieved. If the PCL is retained, the tibial surface must be flat or even posteriorly sloped. If a more conforming component is used in these circumstances, posterior impingement will occur. Substitution of the PCL with a cam and post mechanism is one possible solution to recreate the femoral roll back and also prevent posterior impingement. It would be fair to say that there is still no consensus among surgeons for the regular use of either a cruciate-retaining or cruciate-substituting design for primary TKR. It is mainly a matter of personal preference as well as personal experience as to whether a cruciate-retaining or a posterior-stabilised design is used. Those in favour of cruciate retention argue that the PCL has a beneficial effect on femoral roll back, range of movement, quadriceps efficiency, joint stability and reduced tibial shear forces. Those who favour cruciate substitution question the function of the PCL in isolation (as almost invariably the ACL is excised with all the designs of TKR) but acknowledge the need for posterior stabilisation and use a cam and post mechanism to replicate the physiological functions of the PCL.

3.7.4 Constrained prosthesis In 1951, Walldius developed the hinged prosthesis that bears his name.25 The device was originally made up of acrylic and later made of metal. A hinged prosthesis has considerable appeal because it is technically easy to use as the intramedullary stems makes the prosthesis largely self-aligning and all the ligaments and other soft tissue constraints can be sacrificed because of this self-stabilising feature. The extent of damage to the knee is therefore of no consequence and even the most extreme deformities can be reasonably easily corrected by dividing the soft tissues and resecting sufficient bone. Because of inherent limitations with a simple hinge, including limited range of motion and transmission of stress to the prosthesis cement or prosthesis–bone interface, the early hinged prosthesis were replaced by rotating hinged devices. These constrained the prosthesis in the coronal and sagittal planes but allowed rotation in the axial plane. In cases in which the maximal constraint offered by a linked hinge is not required, unlinked but constrained devices have been used in revision as well as complex primary TKR. The primary characteristic of these devices is the presence of a cam and post mechanism, similar to that found in the posterior-stabilised prosthesis but it is thicker and taller, and provides resistance not only to posterior translation but also to varus and valgus stress. At the

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present time, the use of cemented or uncemented stems with both hinged and constrained but unlinked devices is based on surgeon preference.

3.7.5 Materials Most commonly, TKR bearing surfaces are made of ultra-high molecular weight polyethylene (UHMWPE). UHMWPE has been the standard material for TKR bearings for many years. UHMWPE wear is a common problem of implanted TKRs and has been linked to adverse biological conditions such as bone resorption, and even bone necrosis. The metallic components of TKR devices are most commonly made of cobalt chrome (CoCr) and less frequently titanium. There is no agreement as to which is the better metal. However, there is universal agreement that it is better if the metallic component in contact with the bearing (UHMWPE) is made of cobalt chrome. Two manufacturers have developed ceramic femoral components. One manufacturer (Kyocera) uses alumina ceramic, while the other manufacturer (Smith & Nephew) uses oxidised metal zirconium (Oxinium), which is also a ceramic surface. In laboratory tests, the ceramic-on-polyethylene bearing produced low quantities of polyethylene particles compared with the usual metal-on-polyethylene bearings. As yet, there are no reports about the long-term results of the ceramic TKRs.

3.7.6 Fixation Optimal and long-lasting sound fixation of the TKR implants to host bone is crucial to achieve a good functional outcome. One of the most common reasons for revision of TKR is failure of fixation leading to component loosening. If the fixation fails, the implant gets ‘debonded’ from the host bone and in a weight-bearing joint, like the knee, a loose implant cannot effectively bear and transmit loads. The knee joint in such cases becomes swollen, painful and has a tendency to give way. The knee replacement may be ‘cemented’ or ‘cementless’ depending on the type of fixation used to hold the implant in place. Cemented implants are fixed to the bone immediately when implanted as the cement is a quickhardening grout. Cementless designs are initially fixed with a press-fit or screws and long-term fixation is achieved by bone growth into the implant. The cementless designs tend to have a porous titanium and/or hydroxyapatite coating on their back surface to stimulate bony ingrowth.

3.8

The cemented total knee replacement

A total condylar design with good cementing technique is considered the ‘gold standard’ in TKR today. Various authors have reported excellent long-

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term results in patients undergoing cemented TKR.19,26–28 Despite the success of cemented condylar knee replacements, some authors have advocated a change to cementless fixation for knee replacement patients. This section will look critically at the available evidence to determine what type of fixation is best for TKR patients with respect to age and activity levels and an evidence-based analysis of the current literature regarding knee fixation will help define the role of cemented and cementless knee fixation. The key to achieving a lasting cemented TKR is proper cementing technique. Fixation of the bone cement to the cancellous bony surface is achieved by the irregular configuration of the bony surface and the penetration of the bone cement into the cancellous bone. Thus preparation of the cut bony surfaces is critical. The resected bony surfaces should be cleaned with pulsatile lavage to remove loose bony fragments, fat and blood. The surfaces should be dry to allow good cement penetration. The ideal cement penetration for TKR is 1–2 mm; however, deeper penetration may occur in softer rheumatoid bones, but with hard sclerotic bone this may be less. In such cases, the hard bony surface can be abraded to allow the cement to get a hold on the grooves in the bone. Concerns about the degradation products and debris of poly(methylmethacrylate) (PMMA), deterioration of the bone–cement interface and third-body wear led investigators to search for alternative means of fixation. The success of cementless hip fixation in young patients led to increased interest in cementless fixation in knee replacement. Proponents of cementless fixation in TKR believe that biological fixation has the potential to achieve a more durable bond of the implant to the bone and improve success over cemented fixation. The supposed advantages of cementless fixation include shorter operating time, ease of revision and improved longevity for younger patients. Reduced operative time is probably the most seductive reason for a surgeon to use this technology. Use of cementless implants will eliminate the 10–20-minute time period needed for polymerisation of the cement depending upon whether all the components are fixed at the same time or in sequential manner. Another potential advantage is the ease of revision. In the absence of cement inter-digitation, component removal is supposed to be simplified. The interface between the component and the bone can be divided and there is no need to remove embedded cement fragments. This is particularly relevant in cases of infection where a thorough cleaning of any foreign material is important to achieve eradication of infection. However, it is debatable as to whether it is easy to remove a well-fixed cementless prosthesis or not and also whether the amount of bone loss that may occur in removing a component is of a lesser degree with a cemented component or with a cementless component.

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The main suggested advantage of cementless fixation in young patients is the potential for improved longevity. This outcome is attractive for young and/or active patients. There is also concern that the bone–cement interface has the potential for late deterioration especially in young, active patients. Although this failure mechanism is possible, it occurs infrequently. A cemented TKR is loaded primarily in compression, a force well tolerated at the bone–cement interface. This situation is distinctly different from that of total hip replacement, in which the forces at the cement–bone interface are a combination of tension, compression and shear. Speculating that the level of activity would influence the longevity of the cemented TKR, Diduch et al.29 evaluated the long-term results in patients who were 55 years or younger at the time of index procedure. This was an active group of patients participating regularly in activities, which placed high stresses on the cement interface. The 18-year cumulative survivorship in this study was 94%. There was one case of polyethylene wear and no cases of component loosening in this series. The long-term results of cemented TKR are outstanding.30–32 A well-designed and properly positioned cemented TKR has a greater than 90% chance of surviving more than 15 years. Scuderi et al., looking at 1200 posterior stabilised knees, had a 98% good or excellent result31 at a 14-year survivorship of cemented TKRs, and had a 95% success rate. Font-Rodriguez et al.32, looking at more than 2000 posterior-stabilised metal-backed knees at 14-year follow-up, noted a success rate of 98%. A cemented TKR performed in an elderly patient should have a service life longer than the life of the patient, barring technical failure or infection. The true test of longevity comes, however, from examining the results of cemented TKR in young patients. Ranawat et al.33 reported a 94% 10-year survivorship in patients younger than 55 years old using cemented fixation. Gill et al.34 reported a 98% good or excellent result in their cemented TKR patients younger than 55 years old. Diduch et al.29, in evaluating 114 knees (88 patients) younger than 55 years old, had a 94% good or excellent result at 8 years using cemented fixation. Ritter et al.35 published mediumterm results in patients of 55 years or younger, treated with cruciateretaining cemented TKR. The overall survival rate for these 207 knees was 97.6% with an average follow-up of 9.1 years. These results of cemented fixation in the demanding patient subset of the young and active are encouraging and fail to substantiate the theory of cement interface deterioration with time. Many articles have directly compared cemented fixation with cementless fixation in TKR. Rand and Trouesdale36 looked at more than 11 000 TKRs and performed a survivorship analysis at 10 years. Of these, 92% were successful when cemented fixation was used, whereas only 61% were successful without cement (P < 0.001). Rorabeck37 compared more than

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300 hybrid or cementless TKRs with 224 cemented TKRs and found a 9.6% revision rate in the cementless group at 3 years and only 1.8% revision rate in the cemented group. Barrack et al.38 looked at 82 cementless rotating platform knees and compared them with 76 cemented rotating platform, mobile-bearing knees; 8% of the cementless knees were revised, whereas no cemented knees were revised. The cementless knees also had significantly lower Knee Society scores. Gioe et al.39 evaluated 5760 knees looking at various implants and methods of fixation and found that cementless TKRs had the lowest survival rate of all implants reviewed. Berger et al.,40 in evaluating 113 cementless TKRs at a mean follow-up of 11 years, found that 8% of the tibial components never achieved ingrowth. Rosenberg et al.,41 in a clinical and radiographic comparison of cemented and cementless fixation of the Miller Galante prosthesis (Zimmer, Warsaw, Indiana), found no fixation failures in the cemented group while three cementless cases failed due to lack of tibial bone ingrowth. In other series comparing these two fixation techniques, cementless fixation has shown a precipitous decline in successful results with longer follow-up.42,43 Fehring and Griffin44 also compared patients with cementless TKRs revised for lack of ingrowth with patients with cemented TKR revised for aseptic loosening. There were 27 patients with cementless implants and 36 patients with cemented implants. The authors found that the cementless TKR group was being revised much earlier than the cemented group. Of the cementless patients, 52% had a revision of their arthroplasty within 2 years of their index surgery. The average pain-free interval for the cementless group was only 11 months with 63% having no pain relief after their index procedure. In contrast, only 5 out of 36 (14%) of the patients in the cemented group had a revision in the first 2 years, whereas the average pain-free interval for the cemented group was 64 months. In cementless designs, while femoral ingrowth seems to be more consistent, the tibial component is the usual site of failure because of limited bone ingrowth. Retrieval analysis of cementless tibial components has revealed very little bone ingrowth, bringing into question the longevity of these implants. The use of screws can enhance initial tibial fixation; however, it also introduces new problems of osteolysis and polyethylene wear due to screw migration. The polyethylene debris can migrate through the screw holes in the tibial base plate generating a biologic response, which in turn can cause bone loss and osteolysis. Osteolysis has not been reported in long-term follow-up studies of cemented TKR28 and this is probably attributable to specific designs that produce insignificant numbers of wear particles. Some surgeons prefer the so-called hybrid fixation techniques. There are two such techniques currently used in clinical practice.

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1 Cementation of the tibial component leaving the femoral component cementless. 2 Partial cementation of the tibial component – cementing the base plate and not the tibial stem. Neither of these hybrid fixation methods has evidence-based literature support. Campbell et al.45 looked at 74 hybrid TKRs in which the tibia was cemented and the femur left cementless. They found that the femoral component survivorship was only 87%. They concluded that cementless femoral fixation is unreliable and that this type of hybrid fixation should be abandoned. Gioe et al.39 evaluated 5760 knees and found that cemented metalbacked components had 96% survival whereas hybrid knee replacement had only an 89% success rate. The clinical results of partial cementation of the tibial component have been less predictable than fully cementing the tibial component. Bert and McShane46 found that the tibial tray with partial cementation had significantly greater micromotion compared with the fully cemented construct. Diagnosis of aseptic loosening of cemented or cementless knee TKRs can be difficult. This is particularly true for the cementless implants (Figs 3.11 to 3.13). Although plain radiographs are usually helpful and can be

3.11 Radiograph of a cementless total knee replacement.

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3.12 Radiograph of same patient as in Fig. 3.11 showing component loosening and change in implant position.

3.13 Radiograph showing the same knee after revision with a cemented total knee replacement.

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diagnostic of aseptic loosening, even a slight obliqueness of the film can obscure the radiolucent line edges into the prosthesis. Plain radiographs with a divergence of the beam of only 3° from a plane parallel to the bone-implant interface do not detect a 2 mm lucent line beneath the tibial component. Use of fluoroscopic radiographs of the knee to evaluate aseptic loosening is a better method. The fluoroscopic guidance allows the radiographer to position the X-ray beam parallel to the bone–prosthetic interface so that the presence and extent of radiolucent lines beneath the cementless component can be measured.47 In contrast to failed cemented TKRs in which the cement fragmentation may be evident and the cement– bone interface is not as close to the obscuring metal of the implant, the interface beneath a cementless implant may be more difficult to assess. The porous surface may obscure a demarcation line, and sclerosis beneath the fixed porous implant can become confused with consolidation and spot welds. In cementless TKR, the presence of any line between the implant and the bone shows a lack of bony incorporation in that area. If such lines are extensive or progressive, this line is loose and may be the source of symptoms.47 Another area of controversy is the use of cemented or cementless fixation in revision knee replacement. Stable fixation is an integral part of any TKR, revision surgery being no exception. In revision surgery, there is usually bone loss and the peri-articular bone is compromised. To enhance stability, implants with extended stems have been used. The use of such stems transfers the stress from the deficient plateau to the shaft.48 Whether to cement the stem or not is a question. Excellent mid-term results have been reported using cemented stem fixation. Other authors have reported similar results with cementless stems.49 Fehring et al.50 evaluated a group of patients to determine whether cemented or cementless stem fixation was superior in a series of revision TKR implants with metaphyseal engaging stems. They looked at 107 cemented stems and compared them with 95 press-fit metaphyseal engaging stems. From this study the authors concluded that cementless metaphyseal engaging stems were unpredictable for fixation. Murray et al.51 have also found similar results. Other important considerations are bone quality and cost. The cemented implants are cheaper than the cementless implants even when one takes into consideration the cost of cement, the mixing bowls and the extra theatre time. Surgeons are also aware of the potential limitations of cementless fixation in patients with poor bone quality. This is not the case in cemented fixations. Cemented TKRs have equally good results in patients with rheumatoid arthritis as they do in those with osteoarthritis.52 In conclusion, cemented TKRs have consistently provided superior results as compared with cementless implants and they continue to be the gold standard against which any future products should be compared.

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3.9

Future trends

The use of biological solutions such as ligament reconstruction and cartilage repair or replacement, either in combination or on their own, along with other surgeries like bone realignment, may become popular. Similarly, advances will be made to produce wear-resistant, longer-lasting implants with better metallurgy. Assessment tools will also improve to identify potential problems early enough to prevent component failure. Medication to improve component fixation or to prevent bone loss may come into regular use. Customised implants to suit an individual patient’s anatomy and requirements may become popular. Overall, a combination of biological solutions and compartmental knee replacement may be the preferred treatment option.

3.10

References

1 gillquist, j. and k. messner, Anterior cruciate ligament reconstruction and the long-term incidence of gonarthrosis. Sports Med, 1999. 27(3): 143–56. 2 griffin, t., n. rowden, d. morgan, r. atkinson, p. woodruff and g. maddern, Unicompartmental knee arthroplasty for the treatment of unicompartmental osteoarthritis: a systematic study. ANZ J Surg, 2007. 77(4): 214–21. 3 gioe, t.j., k.k. killeen, d.p. hoeffel, j.m. bert, t.k. comfort, k. scheltema, s. mehle and k. grimm, Analysis of unicompartmental knee arthroplasty in a communitybased implant registry. Clin Orthop Relat Res, 2003. (416): 111–9. 4 marmor, l., Unicompartmental knee arthroplasty. Ten- to 13-year follow-up study. Clin Orthop Relat Res, 1988. (226): 14–20. 5 marmor, l., Unicompartmental arthroplasty of the knee with a minimum tenyear follow-up period. Clin Orthop Relat Res, 1988. (228): 171–7. 6 berger, r.a. r.m. meneghini, j.j. jacobs, m.b. sheinkop, c.j. della valle, a.g. rosenberg and j.o. galante, Results of unicompartmental knee arthroplasty at a minimum of ten years of follow-up. J Bone Joint Surg Am, 2005. 87(5): 999–1006. 7 berger, r.a., d.d. nedeff, r.m. barden, m.m. sheinkop, j.j. jacobs, a.g. rosenberg, and j.o. galante, Unicompartmental knee arthroplasty. Clinical experience at 6- to 10-year followup. Clin Orthop Relat Res, 1999. (367): 50–60. 8 heck, d.a., l. marmor, a. gibson and b.t. rougraff, Unicompartmental knee arthroplasty. A multicenter investigation with long-term follow-up evaluation. Clin Orthop Relat Res, 1993. (286): 154–9. 9 murray, d.w., j.w. goodfellow and j.j. o’connor, The Oxford medial unicompartmental arthroplasty: a ten-year survival study. J Bone Joint Surg Br, 1998. 80(6): 983–9. 10 squire, m.w., j.j. callaghan, d.d. goetz, p.m. sullivan and r.c. johnston, Unicompartmental knee replacement. A minimum 15 year followup study. Clin Orthop Relat Res, 1999. (367): 61–72. 11 svard, u.c. and a.j. price, Oxford medial unicompartmental knee arthroplasty. A survival analysis of an independent series. J Bone Joint Surg Br, 2001. 83(2): 191–4.

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12 price, a.j., j.c. waite and u. svard, Long-term clinical results of the medial Oxford unicompartmental knee arthroplasty. Clin Orthop Relat Res, 2005. (435): 171–80. 13 aaos, number of minimally invasive total knee replacements performed annually in USA. Web-based information service, 2007. Available from: http://orthoinfo. aaos.org/topic.cfm?topic=A00405. 14 department of health. National Joint Registry (England and Wales), 2006. 15 shetty, a.a., a. tindall, p. ting and w. heatley, The evolution of total knee arthroplasty. Part 1: introduction and first steps. Curr Orthop, 2003. 17: 322–5. 16 shetty, a.a., a. tindall, p. ting and w. heatley, The evolution of total knee arthroplasty. Part 2: the hinged replacement and the semi-constrained knee replacement. Curr Orthop, 2003. 17: 403–7. 17 gunston, f.h., Polycentric knee arthroplasty. Prosthetic simulation of normal knee movement. J Bone Joint Surg Br, 1971. 53(2): 272–7. 18 insall, j.n. and m. kelly, The total condylar prosthesis. Clin Orthop Relat Res, 1986. (205): 43–8. 19 scuderi, g.r., w.n. scott and g.h. tchejeyan, The Insall legacy in total knee arthroplasty. Clin Orthop Relat Res, 2001. (392): 3–14. 20 shetty, a.a., a. tindall, p. ting and w. heatley, The evolution of total knee arthroplasty. Part 3: surface replacement. Curr Orthop, 2003. 17: 478–81. 21 insall, j., w.n. scott and c.s. ranawat, The total condylar knee prosthesis. A report of two hundred and twenty cases. J Bone Joint Surg Am, 1979. 61(2): 173–80. 22 burnett, r.s., c.m. haydon, c.h. rorabeck and r.b. bourne, Patella resurfacing versus nonresurfacing in total knee arthroplasty: results of a randomized controlled clinical trial at a minimum of 10 years’ followup. Clin Orthop Relat Res, 2004. (428): 12–25. 23 goodfellow, j. and j. o’connor, The mechanics of the knee and prosthesis design. J Bone Joint Surg Br, 1978. 60-B(3): 358–69. 24 jacobs w.c.h., a.p. limbeek j and a.a.b. wymenga Mobile bearing vs fixed bearing prostheses for total knee arthroplasty for post-operative functional status in patients with osteoarthritis and rheumatoid arthritis. Cochrane Rev, 16 February 2001. (cited 2007; 23 April 2001). Available from: http://www.mrw.interscience. wiley.com/cochrane/clsysrev/articles/CD003130/frame.html. 25 van nieuwenhuyse, w., m. goossens and h. claessens, Total knee replacement with the Walldius hinge prosthesis. A 14 years’ review of 48 cases. Acta Orthop Belg, 1985. 51(4): 520–8. 26 scott, r.d. and t.b. volatile, Twelve years’ experience with posterior cruciateretaining total knee arthroplasty. Clin Orthop Relat Res, 1986. (205): 100–7. 27 scuderi, g.r. and h.d. clarke, Cemented posterior stabilized total knee arthroplasty. J Arthroplasty, 2004. 19(4 Suppl 1): 17–21. 28 colizza, w.a., j.n. insall and g.r. scuderi, The posterior stabilized total knee prosthesis. Assessment of polyethylene damage and osteolysis after a ten-yearminimum follow-up. J Bone Joint Surg Am, 1995. 77(11): 1713–20. 29 diduch, d.r., j.n. insall, w.n. scott, g.r. scuderi and d. font-rodriguez, Total knee replacement in young, active patients. Long-term follow-up and functional outcome. J Bone Joint Surg Am, 1997. 79(4): 575–82.

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30 ranawat, c.s., w.f. flynn, jr. and r.g. deshmukh, Impact of modern technique on long-term results of total condylar knee arthroplasty. Clin Orthop Relat Res, 1994. (309): 131–5. 31 scuderi, g.r., j.n. insall, r.e. windsor and m.c. moran, Survivorship of cemented knee replacements. J Bone Joint Surg Br, 1989. 71(5): 798–803. 32 font-rodriguez, d.e., g.r. scuderi and j.n. insall, Survivorship of cemented total knee arthroplasty. Clin Orthop Relat Res, 1997. (345): 79–86. 33 ranawat, c.s., d.e. padgett and y. ohashi, Total knee arthroplasty for patients younger than 55 years. Clin Orthop Relat Res, 1989. (248): 27–33. 34 gill, g.s., k.c. chan and d.m. mills, 5- to 18-year follow-up study of cemented total knee arthroplasty for patients 55 years old or younger. J Arthroplasty, 1997. 12(1): 49–54. 35 ritter, m.a., j.d. lutgring, k.e. davis, p.m. faris and m.e. berend, Total knee arthroplasty effectiveness in patients 55 years old and younger: osteoarthritis vs. rheumatoid arthritis. Knee, 2007. 14(1): 9–11. 36 rand, j.a. and r.t. Trousdale, Factors affecting the durability of primary total knee prostheses. J Bone Joint Surg Am, 2003. 85-A(2): 259–65. 37 rorabeck, c.h., Total knee replacement: should it be cemented or hybrid? Can J Surg, 1999. 42(1): 21–6. 38 barrack, r.l., s.j. nakamura, s.g. hopkins and s. rosenzweig, Winner of the 2003 James A. Rand Young Investigator’s Award. Early failure of cementless mobilebearing total knee arthroplasty. J Arthroplasty, 2004. 19(7 Suppl 2): 101–6. 39 gioe, t.j., k.k. killeen, k. grimm, s. mehle and k. scheltema, Why are total knee replacements revised?: analysis of early revision in a community knee implant registry. Clin Orthop Relat Res, 2004. (428): 100–6. 40 berger, r.a., j.h. lyon, j.j. jacobs, r.m. barden, e.m. berkson, m.b. sheinkop, a.g. rosenberg and j.o. galante, Problems with cementless total knee arthroplasty at 11 years followup. Clin Orthop Relat Res, 2001. (392): 196–207. 41 rosenberg, a.g., r. barden and j.o. galante, A comparison of cemented and cementless fixation with the Miller–Galante total knee arthroplasty. Orthop Clin North Am, 1989. 20(1): 97–111. 42 moran, c.g., i.m. pinder, t.a. lees and m.j. midwinter, Survivorship analysis of the uncemented porous-coated anatomic knee replacement. J Bone Joint Surg Am, 1991. 73(6): 848–57. 43 nafei, a., s. nielsen, o. kristensen and i. hvid, The press-fit Kinemax knee arthroplasty. High failure rate of non-cemented implants. J Bone Joint Surg Br, 1992. 74(2): 243–6. 44 fehring, t.k. and w.l. griffin, Revision of failed cementless total knee implants with cement. Clin Orthop Relat Res, 1998. (356): 34–8. 45 campbell, m.d., g.p. duffy and r.t. trousdale, Femoral component failure in hybrid total knee arthroplasty. Clin Orthop Relat Res, 1998. (356): 58–65. 46 bert, j.m. and m. mcshane, Is it necessary to cement the tibial stem in cemented total knee arthroplasty? Clin Orthop Relat Res, 1998. (356): 73–8. 47 fehring, t.k. and g. mcavoy, Fluoroscopic evaluation of the painful total knee arthroplasty. Clin Orthop Relat Res, 1996. (331): 226–33. 48 stern, s.h., r.d. wills and j.l. gilbert, The effect of tibial stem design on component micromotion in knee arthroplasty. Clin Orthop Relat Res, 1997. (345): 44–52.

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49 bertin, k.c., m.a. freeman, k.m. samuelson, s.s. ratcliffe and r.c. todd, Stemmed revision arthroplasty for aseptic loosening of total knee replacement. J Bone Joint Surg Br, 1985. 67(2): 242–8. 50 fehring, t.k., s. odum, c. olekson, w.l. griffin, j.b. mason and t.h. mccoy, Stem fixation in revision total knee arthroplasty: a comparative analysis. Clin Orthop Relat Res, 2003. (416): 217–24. 51 murray, p.b., j.a. rand and a.d. hanssen, Cemented long-stem revision total knee arthroplasty. Clin Orthop Relat Res, 1994. (309): 116–23. 52 meding, j.b., e.m. keating, m.a. ritter, p.m. faris and m.e. berend, Long-term followup of posterior-cruciate-retaining TKR in patients with rheumatoid arthritis. Clin Orthop Relat Res, 2004. (428): 146–52.

4 Vertebroplasty and kyphoplasty J. Y E H, St Bartholomew’s and The Royal London Hospitals, UK

Abstract: Vertebroplasty and kyphoplasty are two commonly performed, minimally invasive percutaneous vertebral body augmentation procedures using injectable bone cements. These procedures are increasingly being used for the treatment of spinal fractures. This chapter describes the first-hand clinical experience with these procedures using injectable bony cements. The future trends as seen from a clinical perspective are also considered. Key words: vertebroplasty, kyphoplasty, vertebral body augmentation, injectable bone cements.

4.1

Introduction

Vertebroplasty and kyphoplasty are two common percutaneously performed, minimally invasive, vertebral body augmentation procedures using injectable bone cements. These procedures have gained increasing acceptance by patients and clinicians for the treatment of spinal fractures (Manson and Phillips, 2007). This chapter describes the first-hand experience with these procedures. Section 4.2 gives a brief summary of the health issues involved. Section 4.3 describes the nature and treatment goals of vertebroplasty and kyphoplasty. This is followed in Section 4.4 by a description of the clinical outcome of balloon kyphoplasty in 124 consecutive patients; the procedures were performed at The Royal London Hospital, London, UK. Section 4.5 describes and discusses clinical experience with the use of injectable bone cements in these patients. In the final section, the future trends as seen from a clinical perspective are considered.

4.2

Vertebral compression fractures

The back, or the spine, is made up of a column of bone blocks (the vertebrae) with interconnecting specialized soft tissues, the intervertebral discs (see Fig. 4.1). This arrangement of the vertebral column allows the spine to bend and twist. Each vertebra has a large bony body in the front with a bony arch attached behind. In the spine, these bony arches form a tunnel (the spinal canal) in which the spinal cord and spinal nerves reside. The vertebral column provides a bony shield for the protection of the important 74

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Spinal canal

Pedicle

(b) Nerve root (a)

4.1 Photograph showing models of the vertebral column (a) and a vertebra (b).

but vulnerable neural tissues. A common type of vertebral body fracture is a compression fracture resulting from osteoporosis, tumour or trauma. Compression fractures often result in wedge-, biconcave- or crushed-shaped vertebral bodies (see Fig. 4.2). As well as the personal costs, vertebral body compression fractures impose significant health, financial and social burdens in the UK. The incidence of spinal trauma is about 3500 patients per year with one-third of these patients having paralysis of the limbs (Aspire, 2007). There are currently 40 000 people paralysed with spinal cord injury in the UK. The three most common causes of these spinal traumas are road traffic accidents, sports injuries and falls from heights (Spinal Injuries Association, 2007). It is particularly poignant that the age of these patients is often between 15 and 40 years old and their promising futures are often blighted by the injury. Although spinal cord injuries continue to be a medical challenge with effective treatment yet to become available, effective treatment for vertebral fractures to alleviate pain, stabilize the spine and correct deformity is available, and continues to improve. As well as spinal trauma, spinal tumours are also an important cause of vertebral compression fractures. The most common spinal tumours are tumours that have spread from the prostate gland, breasts, lungs, kidneys and the stomach (Davies and Wang, 2003). The incidence is 184–368 cases per 100 000 people per year. It is estimated that 40–80% of cancer patients

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(a)

(b)

(c)

4.2 Schematic drawing showing wedge (a), biconcave (b) and crushed (c) vertebral body compression fractures.

will have tumours spreading to the spine. On the other hand, tumours arising from the spine itself are relatively rare, and have an incidence of about 50–170 per 100 000 people per year. The most common primary spinal tumour is lymphoproliferative tumours, especially myeloma (Shaffer, 2003). Finally, osteoporosis is common in the UK, with 1 in 3 women and 1 in 12 men over the age of 50 having osteoporosis (National Osteoporosis Society, 2006). As the size of the ageing population increases, osteoporosis is becoming increasingly prevalent. It has been estimated that an osteoporotic fracture occurs every 3 minutes with 70 000 hip fractures, 50 000 wrist fractures and 120 000 spinal fractures per year. This health issue costs the National Health Service and the UK. Government over £1.7 billion per year, i.e. £5 million per day (National Osteoporosis Society, 2006)! Vertebral compression fractures cause pain, deformity of the spine and, sometimes, neurological deficits. The spinal deformity often results in changes of body posture and balance. This in turn leads to increased muscular fatigue with risk of falls and bone fractures. Furthermore, the

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alteration in spinal biomechanics can lead to increased local force at and around the fractured vertebral body leading to further fractures and deformity. It has been shown that people with spinal fracture tend to have decreased mobility, poor body posture, decreased lung function (if the fracture involves the thoracic spine), decreased appetite, sleep disorders, depression, diminished social roles, lower self-esteem and an increase in mortality rate of 23–34% (Gold, 1996; Lyles et al., 1993; Silverman, 1992).

4.3

Kyphoplasty and vertebroplasty

The considerable physiological demands and the long recovery period associated with traditional open spinal surgery have prompted clinicians and patients to seek less invasive alternatives. Vertebroplasty and kyphoplasty, introduced to clinical practice in 1984 and 1998, respectively, are the two commonly performed minimally invasive, X-ray-guided, percutaneous techniques of augmenting fractured vertebral bodies in clinical practice (Galibert et al., 1987; Watts et al., 2001). In vertebroplasty, fine cannulae are introduced into the collapsed vertebral body percutaneously under Xray control. Augmentation materials such as radio-opaque bone cement are then injected into the collapsed vertebral body through the cannulae under pressure (see Fig. 4.3(a)). The biomechanical properties of the vertebral body are therefore improved though the vertebral body remains collapsed. Kyphoplasty is similar to vertebroplasty although there is an additional attempt to restore the height of the collapsed vertebral body. In balloon kyphoplasty, deflated balloons are introduced into the vertebral body. Usually for each vertebral body, two balloons are introduced, one on each

Collapsed vertebral body Working cannula

Inflated balloon

Bone cements

Bone cements (a)

(b)

4.3 Schematic drawing showing vertebroplasty (a) and kyphoplasty (b) procedures.

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side of the vertebral body. The two balloons are then inflated with contrast medium so that the balloon inflation can be closely monitored with X-rays. As the balloons inflate, the collapsed vertebral body is slowly jacked open restoring its height (see Fig. 4.3(b)). The balloons are then deflated and withdrawn leaving voids in the vertebral body, which are then filled with viscous radio-opaque bone cements under minimal pressure injection. As well as balloon bone tamps, other bone tamps are being tried clinically. However, the vast majority of clinical experience is with balloon kyphoplasty. In both vertebroplasty and kyphoplasty, once the bone cements cure in the vertebral body, the spinal fractures are stabilized, and the patients are then able to be mobile. The procedures can be performed under local or general anaesthetic. The recovery period is short, and the patients often go home on the same or next day resuming normal daily activities. The primary clinical indication for vertebroplasty and balloon kyphoplasty is back pain arising from collapsed vertebral bodies. The procedures can also be used to arrest or correct progressive fracture deformity of the vertebral body. The use of these procedures prophylactically, to treat vertebral bodies at risk of collapsing, has not been established in clinical practice (Becker et al., 2007). As the procedures are performed under X-ray guidance, they are generally not suitable for pregnant women because of possible adverse radiation-related effects on the foetus.

4.4

Clinical outcomes

Over a period of 3 years and 8 months between August 2003 and March 2007, 358 balloon kyphoplasty bone tamps were used on 181 vertebral bodies in 124 consecutive patients at The Royal London Hospital, London, UK (J. Yeh, unpublished data, 2007). These patients were followed up prospectively with a mean follow-up period of 18 months (range: 2 weeks–3.5 years). There were 69 males and 55 females with mode age being in the sixties (range: 19–90 years old). There were 59 patients suffering from vertebral body collapses secondary to spinal tumours (49 patients with myeloma, 2 patients with leukaemia and 8 patients had solid tumours), 34 patients had osteoporotic vertebral compression fractures and 31 patients had traumatic spinal fractures. Most of the patients had had back pain related to the spinal fractures for 3–4 weeks (range: 1 day–5.7 years) (J. Yeh, unpublished data, 2007).

4.4.1 Pain relief In this series of patients, there was a decrease in pain from a pre-operative mean of 7.1 points on the visual analogue pain scale (0 being pain free, and 10 being the worst pain imaginable) to 1.9 points in the post-operative

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follow-up. The mean pain relief was 5.2 (±2.8) points (J. Yeh, unpublished data, 2007). None of the patients had worsening pain from the balloon kyphoplasty procedure per se. There was significant pain reduction within the first 24 hours of the procedure being done with further, although of smaller magnitude, pain reduction over a 1-month period. Many patients had significant pain relief immediately after balloon kyphoplasty. This was particularly observed in patients with osteoporosis or spinal tumours. In patients with spinal trauma, the onset of pain relief can take extra few days. This may be due to the presence of other concomitant painful tissue injuries such as spraining of ligaments or bruising of muscles. As well as pain alleviation, there were also a reduction in the analgesic usage following balloon kyphoplasty. Some 53% of the patients stopped analgesics use after balloon kyphoplasty while 30% had a reduction in analgesic usage; 14.5% of patients had no change in their analgesic usage (J. Yeh, unpublished data, 2007). Most of these patients took analgesics for pain arising from other spinal fractures or tissue injuries not amenable to balloon kyphoplasty. Three patients had increased analgesic usage due to progression of their disease. Reduction of analgesic usage is important as analgesics have potentially serious side effects such as drowsiness with opiates, and gastric bleeding with non-steroidal anti-inflammatory drugs. The pain relief with balloon kyphoplasty seems to be maintained with time in the follow-up period, though eight patients had a late relapse of spinal pain. In seven of these patients, this was due to underlying disease progression and one patient had a further fracture. This stresses the importance of treating the underlying pathology, and the fact that cement augmentation affects only the mechanical properties of the vertebral body but not the underlying disease. Optimal tumour and osteoporosis treatment remains vital. It is also important to note that the vertebral body augmentation procedures do not preclude other treatment modalities. Treatment of spinal tumours and osteoporosis can take place alongside the cement augmentation. Both vertebroplasty and balloon kyphoplasty are effective in relieving vertebral body pain in about 90% and 95% of cases, respectively (Brodano et al., 2007; Fournol et al., 2007; Garfin et al., 2006; Hulme et al., 2006; Serra et al., 2007; Taylor et al., 2006). This is particularly impressive if one considers that most of these patients have not responded to conservative management such as spinal bracing, bed rest and potent analgesics. The good results together with small risks have prompted clinical debate and trials as to whether or not these procedures should, in some cases, be considered as first-line treatments rather than after failed conservative management (De Negri et al., 2007; Grafe et al., 2005; Grohs et al., 2005; Ming et al., 2007).

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4.4.2 Functional improvement Functional outcome was assessed using the Oswestry Disability Index (ODI). The ODI assesses ten aspects of daily functions viz. pain intensity, personal care, lifting, walking, sitting, standing, sleeping, sex life, social life and travelling. An ODI of 0–20% indicates minimal disability; the patients can cope with most living activities and usually no treatment is indicated, apart from advice on lifting, sitting and exercise. An ODI of 21–40% indicates moderate disability; the patients experience more pain and difficulty with sitting, lifting and standing; travel and social life are more difficult and they may be disabled from work; personal care, sexual activity and sleeping are not grossly affected. An ODI of 41–60% indicates severe disability; pain remains the main problem in this group of patients; the activities of daily living are affected. Patients with an ODI of 61–80% are severely crippled in function with back pain impinging on all aspects of the patient’s life. Finally, an ODI of 81–100% indicates that the patients are bed-bound. There was a reduction from a mean ODI of 62% before the procedure to 15% ODI in the post-operative follow-up period, a mean reduction of 47% (±5.4%) ODI. This was a reduction from severely disabled/crippled status to minimally disabled status. The pattern of the reduction in disability is very similar to that of pain alleviation. It is likely that the improved function is secondary to the pain relief achieved. There is a significant reduction in the ODI in the first 24 hours to a mean of 35% with further improvement over a one-month period to a mean 15% ODI. It seems that this functional improvement was maintained throughout the follow-up period, except in three patients who had deterioration in their functions secondary to disease progression. There was no worsening of function due the balloon kyphoplasty itself, though in six patients the functional status remained unchanged.

4.4.3 Vertebral body height restoration, wedge and kyphotic angles In this calculation, the average of the adjacent vertebral body heights above and below was taken to be the estimated pre-collapsed height of the collapsed vertebral body. The mean anterior, mid and posterior vertebral height restoration was 6.5%, 6.0% and 1.9%, respectively, of the estimated pre-collapsed vertebral body height. The relatively smaller height restoration posteriorly was because the posterior part of the vertebral body tended to collapse the least compared with the anterior and middle portions of the vertebral body. The mean kyphotic and wedge angle corrections were 3.8° and 3.7°, respectively. The wedge angle is the angle formed by the superior and inferior surfaces of the vertebral body. The kyphotic angle is formed

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by the adjacent vertebral bodies above and below the collapsed vertebral body. Poor height restoration is seen in tumour patients, while near full restoration of vertebral body height is seen in acute traumatic spinal fractures of less than 2 weeks duration. The high proportion of tumour patients and old age of fractures (mean symptom duration of 4 weeks) accounted for the poor height restoration and angle correction seen in this series of patients.

4.4.4 Patient satisfaction When questioned, 98% of the patients were satisfied with their results, and would undergo the procedures again if necessary. Furthermore, they would recommend the procedures to other patients.

4.4.5 Length of hospital stay The mode length of hospital stay post-procedure is 1 day (range: 0–33 days). Often patients were admitted on the day of surgery and discharged the next day. It is also possible to perform the vertebral body augmentations as day cases in selected patients. Patients with acute spinal trauma often have other associated injuries necessitating longer hospital stay.

4.4.6 Complications There were three non-procedure-related mortalities during the follow-up period. The causes of death were myocardial infarction, chest infection and terminal cancer. There were three cases of further vertebral collapses (two due to tumour progression and one due to trauma), three cases of asymptomatic local cement extravasation and one case of transient spinal nerve root irritation that recovered spontaneously. The last two complications were avoidable, and occurred because of poor radiological visualization. The safety of percutaneously invasive procedures depends greatly on the ability to visualize structures clearly on X-ray images. Therefore, good quality imaging, and easily visible instruments and bone cements under Xray are important prerequisites for safe percutaneous vertebral body augmentation procedures.

4.5

Clinical experiences with injectable bone cements

4.5.1 Choice of bone cements The use of human or animal natural bone granules in vertebral body augmentation procedures is difficult. It has its obvious advantages. However,

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the use of such granules does not provide immediate stability of the fractured vertebral body especially in shear and torque loads. There is also the anxiety of the risk of transmissible diseases. Poly(methylmethacrylate) (PMMA) and calcium phosphate bone cements are currently the two commonly used injectable bone substitutes in vertebral body augmentation procedures. Other bone cements are also available clinically such as silicon phosphate and silica-based bone cements. The long-term effects of these artificial bone cements on the vertebral bodies are unknown. Intuitively, in most cases a resorbable bone cement that is progressively replaced by natural bone with time is preferable. The natural bone will have the correct biomechanical properties as well as the ability to remodel and adapt in response to stress load. However, in patients with spinal tumours, an inert material such as PMMA may be desirable as it cannot be invaded by tumours. As there are many drawbacks regarding the biomechanical properties and handling characteristics of current injectable calcium phosphate cements, they are infrequently used. However, the long-term effects of calcium phosphate cements on the vertebral body is unknown, these cements, which are eventually replaced by the natural bone, tend to be used in younger patients. In The Royal London Hospital, injectable calcium phosphate cements are often used in spinal trauma patients under 40 years of age. The calcium phosphate cements cure by crystallization. However, the crystallization process is easily hampered by movement; failure of crystallization results in biomechanically inferior powdery bone cements within the vertebral body. Great care is therefore required in the operating room to ensure that undesirable movement of the spine does not occur during the first 30 minutes of crystallization. Often bed rest is recommended for up to 24 hours. On the other hand, PMMA cements cure quickly in the vertebral body usually within a few minutes, and patients are free to mobilize straight away. The PMMA cements have good resistance to compressive, tensile, shear and torsional loads (Hong et al., 2006; Lieberman et al., 2005). On the other hand, although calcium phosphate has good resistance to compression loads, it has poor resistance to shear loads. It is important that calcium phosphate is only used where there is little likelihood of relative movement between the fractured bone fragments. Relative movement between fractured bone fragments will cause breakage of the calcium phosphate cement in the vertebral body.

4.5.2 Mechanism of pain relief The mechanism of vertebral pain relief of the cement augmentation is unclear (Lewis, 2007). Clinical observation suggests that the presence of bone cement rather than the amount of cement is the key factor for pain

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relief. There seems to be no correlation between pain relief and the amount of bone cement injected (Kaufmann et al., 2006). This is particularly the case with tumours of the spine. In most cases, pain relief is almost immediate following the procedures. Some patients have further pain improvement within days. One of the outstanding key features of vertebroplasty and kyphoplasty is the fast onset of pain relief. Several hypotheses have been put forward regarding the pain relief mechanism. As cement cures within the vertebral body, heat is generated. With PMMA, the temperature generated can reach as high as 70°C (Aebli et al., 2006; Aydin et al., 2006; Baroud et al., 2006; Bartels, 2006; Dunne and Orr, 2002; Li et al., 2004; Stanczyk and van Rietbergen, 2004). This will be sufficient to cause damage to the nerve endings resulting in anaesthesia. However, less exothermic injectable bone cements seem to be equally effective in pain relief when used (Baroud et al., 2006). The monomer, methylmethacrylate, in the PMMA is toxic to surrounding tissues including nerve endings, which may contribute to the anaesthetic effect (Frazer et al., 2005; Revell et al., 1998). However, it is doubtful that there is sufficient quantity of methylmethacrylate present for this effect to be significant. Finally, stabilization of the spinal fracture may account for the pain relief. Given that the quantity of cement required for pain relief is small (viz. 0.5 ml), the amount of bone cement required for fracture stabilization seems to be small, too (Kaufmann et al., 2006). However, fracture stabilization remains the most plausible explanation.

4.5.3 Handling characteristics Both PMMA and calcium phosphate cements need mixing in the operating theatre prior to injection. This poses several problems. The ratio of cement powder and liquid cannot always be guaranteed. The time taken for the mixture to reach the right viscosity partly depends on the room temperature, and may take some time. It is therefore not conveniently available to the clinician. The calcium phosphate cement has a relatively short working time for injection, and does not always allow for carefully controlled injections. The working time of an injectable bone cement is the time duration that the cement remains in a more or less constant and desirable viscosity for the delivery to, and infiltration of, the vertebral body. A bone cement that is ready-mixed with reasonable length of working time, short curing time in the vertebral body and predictable viscosity is highly desirable clinically. A viscous bone cement allows timely filling of the cancellous bone or bone voids in the vertebral body. However, a viscous cement may be difficult to deliver through the injecting cannulae. High-pressure injection to force the cement through the cannulae can result in extravasation of bone

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cement, which is not desirable. The PMMA cement has much better flow characteristics both within the delivery cannulae and in the vertebral body than calcium phosphate cement (Frazer et al., 2005; Noetzel and Kielbassa, 2005). Being able to visualize the bone cement in the vertebral body under Xray control is a very important safety feature as it allows early detection of cement extravasation. Current clinically available PMMA cement preparations are much better visualized than calcium phosphate cement under X-ray.

4.5.4 Bone cement extravasation Bone cement extravasation can occur locally around the vertebral body or it can embolize remotely to other organs (Duran et al., 2007; Groen et al., 2004; Koh et al., 2007; Laredo and Hamze, 2005). Not all cement extravasation causes clinical sequelae. Local extravasation to non-vital tissues is rarely of consequence. However, extravasation of cement into spinal canal or nerve root foramina may compress neural tissues causing paralysis. Local extravasation into the intervertebral discs with advanced degeneration may not be significant. However, the effect on the relatively young or healthy disc is not clear. Intuitively, the presence of bone cement in these discs may alter their biomechanical properties, and evoke unwanted reactions causing pain and destruction of disc materials. Local extravasation can occur though the fracture lines, along cannulae tracts or through the nutrient foramina of the vertebral body. Remote embolization of bone cement occurs because of the high vascularity of the vertebral bodies with their rich venous plexus. As the bone cement enters the venous system, it can cause transient hypotension, and often embolizes to the lungs, which can sometimes be fatal. In vertebroplasty, fine cannulae are used for the passage of the bone cement. Less viscous cements are therefore used, which are injected under pressure. The semi-liquid cements infiltrate the vertebral body more extensively than viscous ones. The combination of less viscous cement and relatively high-pressure injection leads to increased cement extravasation rates. The clinically significant cement extravasation rate for vertebroplasty is about 3% per vertebral level. In kyphoplasty, viscous cements are simply placed into the voids created by the balloons through larger working cannulae under minimal pressure. There is, however, less diffuse infiltration of the bone cement. The clinically significant cement extravasation rate is about 0.3% per level for balloon kyphoplasty (Coumans et al., 2003; De Negri et al., 2007; Greene et al., 2007; Ledlie and Renfro, 2006; Lieberman et al., 2001; Majd et al., 2005; Phillips et al., 2002). In our series of 124 patients, three patients had clinical, non-significant, local extravasation of cement. There was no remote cement embolization.

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4.5.5 Biomechanical changes of the spine and subsequent vertebral body fractures The vertebral body has a hard outer cortical bone shell with an inner cancellous bone structure (see Fig. 4.4). The cancellous bone structure consists of horizontal and vertical connecting trabeculae. The vertebral body is therefore relatively light-weight and strong. With ageing, there is a relative loss of the horizontal trabeculae that support the vertical trabeculae, thinning of cortex and loss of bone mass. This makes the vertebral body weaker and more prone to fractures. Traumatic compression fracture of a young, healthy vertebral body may result in a stiffer and stronger vertebral body, although it is geometrically deformed. In osteoporosis and tumour invasion, the collapsed vertebral body often remains weak and less stiff. In vertebroplasty, there is diffuse infiltration of the cancellous bone by the bone cement while in kyphoplasty, commonly, two blocks of bone cements are formed within the vertebral body. Bone cement augmentation of the vertebral body is often able to increase the strength of the fractured vertebral body above its pre-fracture state, while the stiffness rarely exceeds the pre-fracture value (Kayanja et al., 2006a; Kayanja et al., 2006b; Rotter et al., 2007; Tomita et al., 2003; Tomita et al., 2004; Wilke et al., 2006). Stiffness is a mechanical behaviour of a structure dependent on its material, shape and size. It is a measure of the resistance offered by the structure to the external loads as it deforms. The strength is the force necessary to cause breakage. Stiffness is partially determined by the nature and the quantity of the bone cement injected. The stiffness of the vertebral body plays an important part in energy absorption during loading. Correct stiffness allows optimal energy absorption without risk of fracture or excessive deformation. In addition to local changes in biomechanics, the geometry of the collapsed vertebral body may have a profound effect on the spinal biomechanics in the region. Restoration of vertebral body height can have an important effect on spinal alignment. Correct spinal alignment is important in the

Cortical bone

Cancellous bone

4.4 Photograph showing a model of a transacted vertebra.

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efficient load transfer from one region of the spine down to the adjacent region. In addition, it may also affect effficient function of nearby organs, e.g. forward angling (kyphosis) of the thoracic spine can severely compromise lung function. Some degree of vertebral height restoration may be possible with posture manipulation of the spine on the operating table, e.g. by hyperextending the spine. Vertebroplasty does not have an inherent capability for vertebral height restoration, while balloon kyphoplasty is able to restore the vertebral height partially or fully. The degree of height restoration depends on the age and configuration of the fracture. A deranged load transfer pattern along the spine often leads to excessive focal stress load, and leads to increased risk of further fracture in the already fractured vertebral body or adjacent vertebral bodies. Currently, it is still not clinically clear whether augmentation of the vertebral bodies with bone cement leads to increased risk of further spinal fracture (Fribourg et al., 2004; Lavelle and Cheney, 2006; Pflugmacher et al., 2006; Trout et al., 2006; Voormolen et al., 2006). However, from a biomechanical standpoint, it seems that the risk of further spinal fracture is more related to spinal malalignment than to the mechanical properties of the bone cements (Kim et al., 2004).

4.5.6 Tissue injury The PMMA cement polymerizes and solidifies with an exothermal reaction. In some cases, the core temperature can reach as high as 90°C with a peripheral temperature of 40–60°C (Aebli et al., 2006; Li et al., 2004). Such temperatures can damage the surrounding tissues, which often heal with fibrous scars. Such fibrous tissue may weaken the cement–bone interface (Stanczyk and van Rietbergen, 2004). The fibrous surrounds can be seen clinically on the plain X-rays as a lucent line around the bone cement in the vertebral body. The spinal cord is often protected from this thermal damage by the dura and the continuously flowing cerebrospinal fluid around it (Aydin et al., 2006; Bartels, 2006). In normal healthy intervertebral discs, the nutrient supply is primarily by diffusion across the end plates. Filling the vertebral body extensively with inert bone cements may compromise this diffusion process leading to starvation of the intervertebral discs causing premature degeneration. Theoretically, this may become an important issue in young patients with both vertebral bodies on either side of the intervertebral discs being filled extensively by bone cement.

4.6

Future trends

The minimally invasive percutaneous vertebral body augmentation procedures have created specific requirements and demands for injectable bone

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cements. An ideal injectable bone cement should be both patient and clinician friendly. Physically, the cement should: be ready-mixed with choices of different viscosities; have good flow characteristics; allow sufficient working time; be easily visualized under X-ray; result in a strong cement–bone interface; be non-injurious to surrounding tissues; and be strong with optimal stiffness. Biologically, the cement should be non-toxic, and readily replaced by natural bone without loss of biomechanical integrity. Some of these requirements have begun to be fulfilled in clinical practice though many are still lacking. Already, the possibilities of using the bone cement as a carrier for other therapeutic agents such as antibiotics, chemotherapy drugs and radioactive agents in combination with radiofrequency ablation for tumours are beginning to be explored. In the long term, not only the biomechanical but also the biological requirements of the vertebral body will need to be met, such as injection of stem cells for bone regeneration. Whatever the injected materials, there is a continuing need for safe containment of the injected materials. Technology to reduce cement extravasation by injecting cements into bone tamps that can be left in situ is already in progress. The potential for bone tamps to be left in situ to act also as carriers for other therapeutic agents is yet to be explored. Percutaneous vertebral body augmentation techniques have opened many potential clinical applications with new challenges for biomaterials. Meeting and overcoming these challenges promise a future that should be interesting, fruitful and fulfilling.

4.7

References

aebli n, goss b g, thorpe p, williams r and krebs j (2006), ‘In vivo temperature profile of intervertebral discs and vertebral endplates during vertebroplasty: an experimental study in sheep’, Spine, 31 (15), 1674–1678. aspire (2007), About Aspire, Middlesex, UK, http://www.aspire.org.uk/. aydin s, bozdag e, sunbuloglu e, unalan h, hanci m, aydingoz o and kuday c (2006), ‘In vitro investigation of heat transfer in calf spinal cord during polymethylmethacrylate application for vertebral body reconstruction’, Eur Spine J, 15 (3), 341–346. baroud g, swanson t and steffen t (2006), ‘Setting properties of four acrylic and two calcium-phosphate cements used in vertebroplasty’, J Long Term Eff Med Implants, 16 (1), 51–59. bartels r h (2006), ‘Letter to the editor concerning “in vitro investigation of heat transfer in calf spinal cord during polymethylmethacrylate application for vertebral body reconstruction” (by Aydin S et al.)’, Eur Spine J, 15 (12), 1859. becker s, garoscio m, meissner j, tuschel a and ogon m (2007), ‘Is there an indication for prophylactic balloon kyphoplasty? A pilot study’, Clin Orthop Relat Res, 458, 83–89.

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brodano g b, cappuccio m, gasbarrini a, bandiera s, de salvo f, cosco f and boriani s (2007), ‘Vertebroplasty in the treatment of vertebral metastases: clinical cases and review of the literature’, Eur Rev Med Pharmacol Sci, 11 (2), 91–100. coumans j v, reinhardt m k and lieberman i h (2003), ‘Kyphoplasty for vertebral compression fractures: 1-year clinical outcomes from a prospective study’, J Neurosurg, 99 (1 Suppl), 44–50. davies m r and wang j c (2003), ‘Metastatic spine tumors’. In V J Devlin (Ed.), Spine Secrets, Philadelphia, Hanley & Belfus Inc., pp. 428–433. de negri p, tirri t, paternoster g and modano p (2007), ‘Treatment of painful osteoporotic or traumatic vertebral compression fractures by percutaneous vertebral augmentation procedures: a nonrandomized comparison between vertebroplasty and kyphoplasty’, Clin J Pain, 23 (5), 425–430. dunne n j and orr j f (2002), ‘Curing characteristics of acrylic bone cement’, J Mater Sci Mater Med, 13 (1), 17–22. duran c, sirvanci m, aydogan m, ozturk e, ozturk c and akman c (2007), ‘Pulmonary cement embolism: a complication of percutaneous vertebroplasty’, Acta Radiol, 48 (8), 854–859. fournol m, amoretti n, novellas s, caramella t, chevallier p and bruneton j n (2007), ‘Percutaneous vertebroplasty in symptomatic osteoporotic vertebral compression fractures: review of 50 patients’, J Radiol, 88 (6), 877–880. frazer r q, byron r t, osborne p b and west k p (2005), ‘PMMA: an essential material in medicine and dentistry’, J Long Term Eff Med Implants, 15 (6), 629–639. fribourg d, tang c, sra p, delamarter r and bae h (2004), ‘Incidence of subsequent vertebral fracture after kyphoplasty’, Spine, 29 (20), 2270–2276. galibert p, deramond h, rosat p and le gars d (1987), ‘Preliminary note on the treatment of vertebral angioma by percutaneous acrylic vertebroplasty’, Neurochirurgie, 33 (2), 166–168. garfin s r, buckley r a and ledlie j (2006), ‘Balloon kyphoplasty for symptomatic vertebral body compression fractures results in rapid, significant, and sustained improvements in back pain, function, and quality of life for elderly patients’, Spine, 31 (19), 2213–2220. gold d t (1996), ‘The clinical impact of vertebral fractures: quality of life in women with osteoporosis’, Bone, 18 (3 Suppl), 185S–189S. grafe i a, da fonseca k, hillmeier j, meeder p j, libicher m, noldge g, bardenheuer h, pyerin w, basler l, weiss c, taylor r s, nawroth p and kasperk c (2005), ‘Reduction of pain and fracture incidence after kyphoplasty: 1-year outcomes of a prospective controlled trial of patients with primary osteoporosis’, Osteoporos Int, 16 (12), 2005–2012. greene d l, isaac r, neuwirth m and bitan f d (2007), ‘The eggshell technique for prevention of cement leakage during kyphoplasty’, J Spinal Disord Tech, 20 (3), 229–232. groen r j, du toit d f, phillips f m, hoogland p v, kuizenga k, coppes m h, muller c j, grobbelaar m and mattyssen j (2004), ‘Anatomical and pathological considerations in percutaneous vertebroplasty and kyphoplasty: a reappraisal of the vertebral venous system’, Spine, 29 (13), 1465–1471. grohs j g, matzner m, trieb k and krepler p (2005), ‘Minimal invasive stabilization of osteoporotic vertebral fractures: a prospective nonrandomized comparison

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of vertebroplasty and balloon kyphoplasty’, J Spinal Disord Tech, 18 (3), 238–242. hong s j, park y k, kim j h, lee s h, ryu k n, park c m and kim y s (2006), ‘The biomechanical evaluation of calcium phosphate cements for use in vertebroplasty’, J Neurosurg Spine, 4 (2), 154–159. hulme p a, krebs j, ferguson s j and berlemann u (2006), ‘Vertebroplasty and kyphoplasty: a systematic review of 69 clinical studies’, Spine, 31 (17), 1983– 2001. kaufmann t j, trout a t and kallmes d f (2006), ‘The effects of cement volume on clinical outcomes of percutaneous vertebroplasty’, Am J Neuroradiol, 27 (9), 1933–1937. kayanja m, evans k, milks r and lieberman i h (2006a), ‘The mechanics of polymethylmethacrylate augmentation’, Clin Orthop Relat Res, 443, 124– 130. kayanja m m, schlenk r, togawa d, ferrara l and lieberman i (2006b), ‘The biomechanics of 1, 2, and 3 levels of vertebral augmentation with polymethylmethacrylate in multilevel spinal segments’, Spine, 31 (7), 769–774. kim s h, kang h s, choi j a and ahn j m (2004), ‘Risk factors of new compression fractures in adjacent vertebrae after percutaneous vertebroplasty’, Acta Radiol, 45 (4), 440–445. koh y h, han d, cha j h, seong c k, kim j and choi y h (2007), ‘Vertebroplasty: magnetic resonance findings related to cement leakage risk’, Acta Radiol, 48 (3), 315–320. laredo j d and hamze b (2005), ‘Complications of percutaneous vertebroplasty and their prevention’, Semin Ultrasound CT MR, 26 (2), 65–80. lavelle w f and cheney r (2006), ‘Recurrent fracture after vertebral kyphoplasty’, Spine J, 6 (5), 488–493. ledlie j t and renfro m b (2006), ‘Kyphoplasty treatment of vertebral fractures: 2-year outcomes show sustained benefits’, Spine, 31 (1), 57–64. lewis g (2007), ‘Percutaneous vertebroplasty and kyphoplasty for the stand-alone augmentation of osteoporosis-induced vertebral compression fractures: present status and future directions’, J Biomed Mater Res B Appl Biomater, 81 (2), 371–386. li c, mason j and yakimicki d (2004), ‘Thermal characterization of PMMA-based bone cement curing’, J Mater Sci Mater Med, 15 (1), 85–89. lieberman i h, dudeney s, reinhardt m k and bell g (2001), ‘Initial outcome and efficacy of “kyphoplasty” in the treatment of painful osteoporotic vertebral compression fractures’, Spine, 26 (14), 1631–1638. lieberman i h, togawa d and kayanja m m (2005), ‘Vertebroplasty and kyphoplasty: filler materials’, Spine J, 5 (6 Suppl), 305S–316S. lyles k w, gold d t, shipp k m, pieper c f, martinez s and mulhausen p l (1993), ‘Association of osteoporotic vertebral compression fractures with impaired functional status’, Am J Med, 94 (6), 595–601. majd m e, farley s and holt r t (2005), ‘Preliminary outcomes and efficacy of the first 360 consecutive kyphoplasties for the treatment of painful osteoporotic vertebral compression fractures’, Spine J, 5 (3), 244–255. manson n a and phillips f m (2007), ‘Minimally invasive techniques for the treatment of osteoporotic vertebral fractures’, Instr Course Lect, 56, 273–285.

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ming j h, zhou j l, zhou p h and zhou j p (2007), ‘Comparison of therapeutic effect between percutaneous kyphoplasty and pedicle screw system on vertebral compression fracture’, Chin J Traumatol, 10 (1), 40–43. national osteoporosis society (2006), Osteoporosis Facts and Figures, Bath, UK. Available at: http://www.nos.org.uk/news/facts_and_figures.htm. noetzel j and kielbassa a m (2005), ‘Calcium phosphate cements in medicine and dentistry – a review of literature’, Schweiz Monatsschr Zahnmed, 115 (12), 1148– 1156. pflugmacher r, schroeder r j and klostermann c k (2006), ‘Incidence of adjacent vertebral fractures in patients treated with balloon kyphoplasty: two years’ prospective follow-up’, Acta Radiol, 47 (8), 830–840. phillips f m, todd w f, lieberman i and campbell-hupp m (2002), ‘An in vivo comparison of the potential for extravertebral cement leak after vertebroplasty and kyphoplasty’, Spine, 27 (19), 2173–2178. revell p a, braden m and freeman m a (1998), ‘Review of the biological response to a novel bone cement containing poly(ethyl methacrylate) and n-butyl methacrylate’, Biomaterials, 19 (17), 1579–1586. rotter r, pflugmacher r, kandziora f, ewert a, duda g and mittlmeier t (2007), ‘Biomechanical in vitro testing of human osteoporotic lumbar vertebrae following prophylactic kyphoplasty with different candidate materials’, Spine, 32 (13), 1400–1405. serra l, kermani f m, panagiotopoulos k, de rosa v and vizioli l (2007), ‘Vertebroplasty in the treatment of osteoporotic vertebral fractures: results and functional outcome in a series of 175 consecutive patients’, Minim Invasive Neurosurg, 50 (1), 12–17. shaffer w o (2003), ‘Primary spine tumors’. In V J Devlin (Ed.), Spine Secrets, Philadelphia, Hanley & Belfus Inc., pp. 420–427. silverman s l (1992), ‘The clinical consequences of vertebral compression fracture’, Bone, 13 (Suppl 2), S27–S31. spinal injuries association (2007), Causes and Frequency of SCI in the UK, Milton Keynes, UK. stanczyk m and van rietbergen b (2004), ‘Thermal analysis of bone cement polymerisation at the cement-bone interface’, J Biomech, 37 (12), 1803–1810. taylor r s, taylor r j and fritzell p (2006), ‘Balloon kyphoplasty and vertebroplasty for vertebral compression fractures: a comparative systematic review of efficacy and safety’, Spine, 31 (23), 2747–2755. tomita s, kin a, yazu m and abe m (2003), ‘Biomechanical evaluation of kyphoplasty and vertebroplasty with calcium phosphate cement in a simulated osteoporotic compression fracture’, J Orthop Sci, 8 (2), 192–197. tomita s, molloy s, jasper l e, abe m and belkoff s m (2004), ‘Biomechanical comparison of kyphoplasty with different bone cements’, Spine, 29 (11), 1203– 1207. trout a t, kallmes d f and kaufmann t j (2006), ‘New fractures after vertebroplasty: adjacent fractures occur significantly sooner’, Am J Neuroradiol, 27 (1), 217– 223. voormolen m h, lohle p n, juttmann j r, van der graaf y, fransen h and lampmann l e (2006), ‘The risk of new osteoporotic vertebral compression fractures in the year after percutaneous vertebroplasty’, J Vasc Interv Radiol, 17 (1), 71–76.

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watts n b, harris s t and genant h k (2001), ‘Treatment of painful osteoporotic vertebral fractures with percutaneous vertebroplasty or kyphoplasty’, Osteoporos Int, 12 (6), 429–437. wilke h j, mehnert u, claes l e, bierschneider m m, jaksche h and boszczyk b m (2006), ‘Biomechanical evaluation of vertebroplasty and kyphoplasty with polymethyl methacrylate or calcium phosphate cement under cyclic loading’, Spine, 31 (25), 2934–2941.

5 Antibiotic-impregnated polymethylmethacrylate (PMMA) spacers in hip surgery K. A N AG N O S TA K O S and J. K E L M, Universitätsklinikum des Saarlandes, Germany

Abstract: The implantation of antibiotic-loaded spacers is a standard procedure in the treatment of hip joint infections. This chapter reviews the current experience of antibiotic-impregnated hip spacers and focuses particularly on the pharmacokinetic and mechanical properties of these interim protheses as well as on the clinical outcome and possible related complications. Key words: hip joint infection, poly(methylmethacrylate) (PMMA), bone cement, hip spacers, antibiotic elution.

5.1

Introduction

Despite the progress in the prophylaxis and management of infections, late infections after total hip arthroplasty (THA) still remain a hazardous problem in orthopaedic surgery. Depending on the experience of the orthopaedic surgeon, one- and two-stage procedures can be performed for infection management. In the one-stage procedure (Wagner and Wagner, 1995), the infected prosthesis is removed, meticulous debridement of the situs and jet-lavage are performed and a new prosthesis is reinserted in a single procedure. Antibiotic-loaded bone cement is used for fixation of the prosthesis and a systemic antibiotic treatment is given for 4–6 weeks for infection management. In the two-stage procedure, the infected prosthesis is removed and the new prosthesis is reimplanted after 3–4 months. Between stages, local antibiotic therapy using antibiotic-loaded cement media and an intravenous antibiotic regimen should eradicate the infection. A two-stage procedure consists of either a resection exchange arthroplasty, the so-called Girdlestone procedure, or the insertion of an antibiotic-loaded interim prosthesis/spacer (Wagner and Wagner, 1995, Joseph et al., 2003, Anagnostakos et al., 2006). In the Girdlestone procedure, an excision arthroplasty of the femoral head/neck is performed and antibiotic-loaded beads are inserted into the 92

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5.1 Compared with the Girdlestone procedure, the spacer implantation prevents any soft-tissue contracture and the joint mobility is preserved.

femur or the resection cavity for local antibiotic therapy. The patient is left with an unstable joint and with a leg length discrepancy resulting from softtissue contracture, and the energy cost during gait is higher (Fig. 5.1). Furthermore, commercially available beads are loaded only with gentamicin, so that, in infections caused by gentamicin-resistant bacteria, no adequate local antibiotic therapy can be provided. Over the past two decades, antibiotic-impregnated spacers have become a very popular procedure in the treatment of late infections after THA. Despite the fact that the idea of a ‘spacer’ had already been discussed in the late 1970s (Hovelius and Josefsson, 1979), it was not until the early 1990s that spacers became a widespread technique in the management of such infections (Zilkens et al., 1990, Abendschein, 1992, Duncan and Beauchamp, 1993). The benefits of this method are: • immediate treatment of the infection source by locally reaching high antibiotic levels; • optional impregnation of the bone cement according to the antibiotic sensitivity profile of the pathogenic organism; • maintance of joint mobility; • protection of bone stock; • limitation of scar formation;

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• absence of soft tissue contraction (usually resulting in a leg length discrepancy); • ease for reimplantation. Before deciding between the one- and two-stage options, the acuteness or chronicity of the infection, the type of infecting organism, its antibiotic sensitivity profile, its ability to synthesize a glycocalyx, the occurrence of severe soft-tissue damage and the extent of bone loss have to be considered. Over the past decades, numerous studies have compared one- and twostage procedures and demonstrated that the latter is superior with regards to infection eradication. In the study of Garvin and Hanssen (1995), direct exchange arthroplasty had an 82% success rate in the eradication of infection compared with the 91% success rate observed in the two-stage technique. Ure et al. (1998) emphasized that a direct exchange arthroplasty can only be carried out in early infections, if the patient is not immunocompromised, the infecting organism is of low virulence (no methicillinresistant or Gram-negative bacteria), the surgeon is experienced and there are no major skin, soft-tissue or existing osseous defects. A two-stage approach permits identification of the infecting organism, determination of antibiotic sensitivity and appropiate adjustment of antibiotic therapy before reimplantation. Other benefits include adequate debridement of necrotic or infected tissues and removal of cement plus greater flexibility in reconstructive options. However, the prolonged hospitalization and its associated costs, the delayed mobilization and rehabilitation, and the risks of an additional surgery may be drawbacks, especially in elderly patients. Antibiotic-impregnated hip spacers are reported to have success rates of over 90% (Anagnostakos et al., 2006). However, confusion still reigns regarding various details in the use of these constructs. Are standardized spacers better than custom-made ones? Which antibiotic(s) is/are ideal for impregnation of the cement? Can the insertion of any metallic components enhance the mechanical properties of the spacers? Has the insertion of such metallic endoskeletons any influence on the elution properties of the spacers? What complications may occur between stages? In this chapter, a systematic review of the literature should clarify these questions and demonstrate the current experience of hip spacers.

5.2

Construction of hip spacers

Currently, three types of hip spacers are in use: • • •

manually shaped; standardized, prefabricated, commercially available (e.g. Spacer G); standardized, moulded, not commercially available (Fig. 5.2).

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5.2 Mould for standardized production of hip spacers. This particular mould produces a hip spacer with a total surface area of 13 300 mm2 (stem length 10 cm, head diameter 50 mm).

The disadvantage of the commercially available spacers lies in the fact that no further addition of an antibiotic into the cement, in response to the sensitivity profile of the pathogenic organism, is possible.

5.3

Diagnosis of infection

Many authors have emphasized the value of clinical history and physical examination, radiological evaluation and laboratory data – including Creactive protein (CRP), erythrocyte sedimentation rate (ESR) and leucocyte blood count – in confirming the diagnosis of a hip joint infection (Koo et al., 2001, Takahira et al., 2003, Durbhakula et al., 2004, Hsieh et al., 2004). An ESR of >30–40 mm/h, combined with a CRP level of >10–20 mg/l, is considered highly suggestive of infection (Durbhakula et al., 2004, Hsieh et al., 2004). However, these values could be false-negative if the patient has been already treated with antibiotics. Radiological findings depend on the stadium of the infection and may be normal. A preoperative hip aspiration was routinely performed by some authors (Isiklar et al., 1999, Takahira et al., 2003), whereas others found it of limited value, since the reported rates of negative preoperative aspiration range from 7 to 50% (Koo et al., 2001). In addition, intra-operative findings may suggest an infection despite a negative preoperative aspiration (Kraay et al., 1992). Multiple biopsy samples from the infected area should clarify and distinguish a contamination from an infection of any clinical significance. Besides the hip-joint-related infection parameters, it should be born in mind that general patient factors – including nutritional status, diabetes mellitus, rheumatoid arthritis, previous sepsis and concurrent illness – may swing the balance in favour of manifestation of the infection. Therefore, cooperation with other medical faculties might be required in order to provide an adequate premise for infection eradication.

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5.4

Consistency and function of antibiotic-loaded hip spacers

In the early years of the use of hip spacers, most of them were hand-formed. In recent years, the majority have been replaced with standardized spacers, produced using moulds. Most studies have used Palacos bone cement due to its excellent elution characteristics. Cephalosporins, penicillin G, aminoglycosides and glycopeptides have been added to the cement, the two latter having been used most often (Table 5.1). Almost three-quarters of all placed spacers reported in the literature have had mechanical support from metallic components (Fig. 5.3). Most spacers function as a hemiarthroplasty, whereas only a few offer the advantages of a THA (Duncan and Beauchamp, 1993, Duncan and Masri, 1994, Masri et al., 1994, 1998a, 1998b, Younger et al., 1997, 1998, Haddad et al., 1999, Hsieh et al., 2004, 2005, 2006a, 2006b). Almost half of the studies report on fixation of the spacer by cementation to the proximal part of the Table 5.1 Possibilities for enhancement of the mechanical stability of hip spacers by insertion of metallic components Authors

Metallic component used

Barrack 2002 Bertazzoni Minelli et al. (2004) Deshmukh et al. (1998) Duncan and Beauchamp (1993) Hsieh et al. (2004) Kraay et al. (1992) Morimoto et al. (2003) Pearle and Sculco (2002) Shin et al. (2002) Takahira et al. (2003) Zilkens et al. (1990)

Rush pin Hollow cylindrical rod Küntscher nail Stainless steel endoskeleton Kirschner wires Cerclage wires Gamma locking nail Steinmann pins Femoral endoprosthesis Ender nail Metallic telescopic shaft

5.3 Insertion of a titanium pin for improvement of the mechanical properties of a hip spacer and, in the case of a fracture, prevention of luxation of the fragments.

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femur, the other half used snap-fits as the method of fixation. Unfortunately, to date there is no study comparing the clinical performance and the complications among spacers with respect to their articulation and the fixation method. Although a proximal cementation of the spacer to the femur might preserve leg length and prevent rotation, no study has demonstrated which one of the two methods is best. Moreover, THA-like spacers may improve the congruence of the joint compared with hemiarthoplasty-like systems; however, no reports have studied whether clinical performance is better with THA-like spacers.

5.5

Mechanical stability and behaviour of hip spacers

The mechanical stability of spacers is determined and influenced by many parameters, including: geometry; the ageing, storage and type of cement; the type and content of antibiotic supplements; the presence of an endoskeleton (Table 5.1); and the standardization of spacer preparation (such as atmospheric composition during mixing, and the frequency and duration of the particular mixing process). Mechanical stress experiments with gentamicin-loaded spacers showed an average failure load of 1.6 kN (Schöllner et al., 2003). The inclusion of K-wires in the spacers prevented any dislocation of the fragments, but could not significantly improve the mechanical properties. In contrast, Kelm et al. (2001) reported an average failure load of 20 kN with antibiotic-loaded cement that did not include any supporting metal components. Recently, Affatato et al. (2003) investigated in an in vitro model the changes in the polymethylmethacrylate (PMMA) polymer conformation induced by wear. Although the wear behaviour varied with the area of the spacer and the amount of debris was higher than in tests with no temporary prostheses, the authors concluded that partial weight bearing can be allowed, since at the time of the prosthesis reimplantation any particle debris can be removed by jet lavage. The mechanical strength of cement is not only influenced by the type of antibiotic and atmospheric pressure but also by the ratio at which the antibiotics are mixed into the cement (Lautenschlager et al., 1976). Proportional weights of up to roughly 5% display a negligible influence on the mechanical strength of the resulting cement (Levin, 1975, Lautenschlager et al., 1976, Murray, 1984), whereas larger amounts of antibiotic powder will make the cement harder to mix and increase the possibility of inhomogenities. There is no quantitative information available for an ideal antibiotic/cement ratio, but most surgeons do not exceed a ratio of 10%. However, Hsieh et al. (2005) reported an antibiotic mixture ratio of up to 20% and had no difficulties in the fabrication of the prosthesis. In addition to the manufacturing process, other factors might compromise the function of the spacer, including the residual bone quality after

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the first surgery or deficient soft tissue. Although numerous initial strength tests for antibiotic-impregnated cement blocks or discs have been carried out (Lautenschlager et al., 1976, Lidgren et al., 1987, Askew et al., 1990, Armstrong et al., 2002, Lewis, 2003), only the studies by Kelm et al. (2001) and Schöllner et al. (2003) attempted to provide hard static data on these constructs. Unfortunately, the clinical relevance of these reports is hindered by the limited number of spacers investigated and the experimental conditions (an axial strength test does not represent the stress that a spacer is exposed to in the human body). The mechanical strength of the spacers could be increased by the addition of a metallic endoskeleton, although the benefits of such a construct have not been thoroughly determined. In addition, a spacer fracture exposing the metal may lead to bacterial recolonization and/or reinfection. It is important to note here that the improvement of the mechanical properties of hip spacers should not exceed the point where the spacer becomes harder than the bone, because, if it does, a femoral fracture might be expected in the case of a spacer luxation or dislocation. In addition, studies need to be conducted to investigate the antibiotic elution profile of spacers with metal inclusion in order to clarify a possible alteration of release characteristics. Beyond the enhancement of the spacer’s stability by inserting a metallic endoskeleton, the geometry of the spacer itself might also play a role in the emergence of clinical complications. Unfortunately, only a single report exists in the literature in which the various causes of clinical failures of spacers were investigated. Leunig et al. (1998) demonstrated that, with regard to the geometry, a relatively small spacer neck/head ratio should be aimed for (<0.73). In order to avoid dislocations, deep insertion of the spacer into the femur is recommended (>57 mm).

5.6

Pathogenic organisms

Most of the reported hip joint infections were caused by Staphylococus aureus, Staphylococus epidermidis, methicillin-resistant S. aureus (MRSA) and Escherichia coli. In a few cases, a polymicrobial infection has been diagnosed. Depending on the antibiotic sensitivity profile of the infecting organism, local therapy was complemented with a wide range of oral and/or parenteral antibiotics, with cephalosporins being the most frequently administered.

5.7

Antibiotic choice

An important topic in the use of antibiotic-loaded spacers is the impregnation of bone cement. Not every antibiotic qualifies equally for incorporation into bone cement: desirable characteristics include (Murray, 1984, Wahlig, 1987):

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Table 5.2 Antibiotic combinations for impregnation of PMMA hip spacers for local antibiotic therapy Authors

Antibiotic combination used (per 40 g PMMA)

Anagnostakos et al. (2005)

0.5 g gentamicin + 1 g vancomycin 0.5 g gentamicin + 0.2 g teicoplanin 0.5 g gentamicin + 0.75 g synercid

Barrack (2002)

3.6 g tobramycin + 1 g vancomycin

Haddad et al. (1999)

2.4–3.6 g tobramycin + 1–1.5 g vancomycin

Hsieh et al. (2004)

Vancomycin, teicoplanin, gentamicin, aztreonam, piperacillin

Isiklar et al. (1999)

2–3 g vancomycin

Kelm et al. (2006)

0.5 g gentamicin + 2 g vancomycin

Koo et al. (2001)

1 g gentamicin, 1 g vancomycin, 1 g cefotaxime

Masri et al. (1998b)

1.2–4.8 g tobramycin + 1–2 g vancomycin

Younger et al. (1997)

Tobramycin, penicillin G, vancomycin, ceftixozime, streptomycin

• availability in powder form; • wide antibacterial spectrum, bactericidal at low concentrations; • elution from PMMA in high concentrations for prolonged periods; • thermal stability; • low or no risk of allergy or delayed hypersensitivity; • low influence on the mechanical properties of the cement; • low serum protein binding. Aminoglycosides and glycopeptides are known to be the two groups of antibiotics that fulfil most of these criteria (Table 5.2). Unfortunately, the recent increase in the number of resistant bacteria strains (Neu, 1992, Krcmery et al., 1996, Cercenado et al., 1996, Wendt et al., 1998) requires an adequate treatment strategy in order to avoid failures and multiple surgical interventions. Thus, although an antibiogram might be useful, it does not always appear helpful since some of the recommended antibiotics are inactivated when mixed with the cement or during its polymerization process (chloramphenicol, tetracycline) (Buchholz and Engelbrecht, 1970, Ger et al., 1977). Equally, a low antibiotic release from the cement might limit its application. Not only is the right antibiotic choice important for an adequate local antibiotic therapy, but also the amounts of each antibiotic that are incorporated into the cement. Depending on the ratio, for example, between aminoglycosides and glycopeptides, different synergistic effects between these

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antibiotic groups should be expected with regard to their elution properties (Masri et al., 1998b, Anagnostakos et al., 2006). Furthermore, it is necessary to distinguish between the use of antibiotics in Europe and in the United States and Canada, since in Europe gentamicin is mostly used for impregnation of PMMA, whereas in the United States and Canada tobramycin is used (Table 5.2).

5.8

Antibiotic elution

Masri et al. (1994) measured the intra-articular antibiotic concentrations in the first days after the insertion of vancomycin–tobramycin-loaded spacers. They confirmed the in vitro result reported by Greene et al. (1998) and showed that tobramycin eluted better than vancomycin. Peak concentrations on day 1 were 107 μg/ml for tobramycin and 19 μg/ml for vancomycin, determined from the wound drainage fluids. These concentrations were 10–30 times higher than the minimal inhibiting concentrations (MICs) of the infecting organisms. An increase in the tobramycin dose enhanced the elution of tobramycin and vancomycin, whereas an increase in the vancomycin concentration lacked such an effect (Masri et al., 1998b). A sufficient elution of antibiotics from PROSTALAC (prosthesis of antibiotic-loaded acrylic cement) could be measured over a period of at least 4 months. The duration of the spacer implantation did not have a statistically significant influence on the elution characteristics of both antibiotics (Masri et al., 1998b). Isiklar et al. (1999) reported mean concentrations of 57 μg/ml for vancomycin eluted on day 1 from vancomycin-impregnated spacers in the treatment of orthopaedic implant-related S. epidermidis infections, also determined from the drainage fluid. Hsieh et al. (2006b) reported recently on the elution of vancomycin and aztreonam from hip spacers. Vancomycin peak concentrations were initially 1538 μg/ml and fell after 7 days to a mean value of 519 μg/ml. These high concentrations could be attributed either to the high amount of antibiotics incorporated into the cement (4 g vancomycin–4 g aztreonam–40 g PMMA) or the fact that aztreonam might have a different influence on the pharmacokinetics of vancomycin than aminoglycosides have, so that another synergistic effect might result. A recent in vitro study (Anagnostakos et al., 2005) investigated the efficacy of mono- and bi-antibiotic-loaded spacers with regard to bacteria growth inhibition and antibiotic release. Gentamicin–vancomycin-loaded spacers were the most effective in the growth inhibition of S. epidermidis and MRSA, whereas gentamicin–teicoplanin-impregnated spacers demonstrated the best results against E. faecalis and S. aureus. Reports on the elution characteristics of spacers are encouraging but require further investigations about the precise in vivo release mechanism. In vivo and in vitro studies showed a sufficient antibiotic elution; however, these studies have

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limitations. In in vivo studies (Masri et al., 1998b), no data were reported on the antimicrobial properties at such antibiotic concentrations; in in vitro studies (Anagnostakos et al., 2005), the experimental conditions are different from the normal in vivo conditions; these limitations may be overcome by a suitable in vivo animal model. Some concerns have been expressed with regard to the ideal reimplantation time and whether too early or too late spacer removal with prosthesis reimplantation might be associated with an infection persistence or recurrence. Bertazzoni Minelli et al. (2004) and Kelm et al. (2006) studied the residual antibiotic and antimicrobial properties of explanted spacers in vitro. In the first study (Bertazzoni Minelli et al., 2004), 0.05–0.4% and 0.8– 3.3% of the initial amounts present of gentamicin and vancomycin, respectively, were released in vitro over a time period of 10 days, indicating that a sufficient antibiotic release can persist over several months. Kelm et al. (2006) reported similar elution values of gentamicin and vancomycin and their spacers demonstrated sufficient antimicrobial properties for at least 14 days in vitro. Hence, an infection persistence should not necessarily be attributed to the insufficient elution kinetics of the spacers. Following surgery, the outcome of infections can be particularly dramatic because of the extensive size of soft tissue, and thus systemic antibiotic therapy should be given. Unfortunately, such supporting antibiotic treatments are not without risks, as reported by Koo et al. (2001). In that study the authors described the occurrence of a transient liver dysfunction and bone marrow depression following the simultaneous implantation of an antibiotic-loaded spacer and intravenous treatment with antibiotics. Holtom et al. (1998) found a positive correlation between the surface area of the vancomycin-loaded spacers and their elution characteristics in vitro. By increasing the surface/volume ratio from 0.24 (control spacer) to 0.30, a 40% enhancement in antibiotic release could be achieved. Greene et al. (1998) showed that tobramycin displayed a higher release rate than vancomycin and that Palacos bone cement allowed higher antibiotic elution than Simplex bone cement. Last but not least, an important issue still remains unclear with regard to the antibiotic elution from spacers. The local antibiotic concentrations vastly exceed those that result after intravenous application; however, it is unknown whether too high local concentrations could have a toxic or necrotic effect on the soft tissues. These concerns may be addressed by histopathologic findings providing us with some answers on this topic.

5.9

Clinical experience

It is not easy to evaluate objectively the clinical outcome of patients who have suffered from a hip joint infection. Most patients are limited by their

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general medical condition and are handicapped in the hip joint function by the many surgeries they have undergone. All studies about hip spacers are retrospective and, in combination with different hip scores, the joint function cannot be always compared directly among the published studies. Only a few reports demonstrate the various dimensions of each hip score (Haddad et al., 1999, Hsieh et al., 2004, 2005, 2006a, Younger et al., 1998). In one of the latest reports by the group developing PROSTALAC, Haddad et al. (1999) reported infection eradication in 77 of 81 hips. The mean Harris hip score increased from 34 points before spacer implantation to 56 during interim stages and to 76 at the time of reimplantation. In a study prior to this one, the great majority of patients were reportedly satisfied with the outcome of the treatment (43/48) (Younger et al., 1997). Such subjective reports were supported by the study of Koo et al. (2001), which used the Merle d’Aubigne hip score and showed an increase from 5.9 to 14.6 points. Takahira et al. (2003) confirmed a similar improvement after spacer implantation using the hip score of the Japanese Orthopaedic Association, which increased from 30 to 73 points. Furthermore, treatment of infection by using a spacer seems to offer great advantages, even in cases with massive bone loss of the proximal femur or the acetabulum (Duncan and Beauchamp, 1993, Younger et al. 1998, Isiklar et al., 1999, Hsieh et al., 2004). Prosthesis reimplantation can be a technically demanding procedure due to leg length discrepancy, softtissue shortening and disuse osteoporosis. The implantation of a spacer in such cases gives the patient a functional joint during the interim period. Duncan and Beauchamp (1993) and Hsieh et al. (2004) reported encouraging results in the treatment of hip infections by spacer implantation followed by reconstruction with allografts. None of them observed any recurrence of infection, and the joint function was enhanced between stages. In addition to their use in late infections after THA, spacers can also be used in the management of infections of the proximal femur after osteosynthesis. Recently, Hsieh et al. (2006a) reported very encouraging results in the treatment of deep hip infections following intratrochanteric fracture by insertion of antibiotic-loaded spacers. In a series of 12 patients no infection persistence or reinfection could be observed at an average follow-up of 4.8 years. Rodriguez and Ziran (2007) covered a gamma nail with antibiotic-loaded cement and implanted it as a spacer during staged reconstruction of an infected proximal femur non-union. In our own patient series (previously unpublished), we were able to make similar observations to those mentioned above in the treatment of infections of the proximal femur after osteosynthesis or destructive bacterial coxitis. A significant hip score increase could be seen between ‘infection situation’ and ‘between stages’, and ‘between stages’ and ‘follow-up’, at 1

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5.4 Left: destructive bacterial coxitis. Middle: infection treatment by implantation of a standardized, articulating spacer. Right: 3 months later, the spacer was explanted and a prosthesis implanted.

year after prosthesis implantation (Figs 5.4 and 5.5). In particular, the dimension ‘score’ demonstrated a highly significant increase between the time periods which correlated with the literature data (Younger et al., 1998, Isiklar et al., 1999). In addition to the material properties of the spacers an optimal interaction and articulation between the bone and the spacer favours a positive result of the joint replacement. The insertion of a standardized, articulating and antibiotic-impregnated prosthesis into the bone reduces the risk of crepitus during hip mobilization and helps in preserving the bone stock (Ries and Jergesen, 1999, Shin et al., 2002). This method has been reported to increase hip mobility and reduce pain during interim stages and after reimplantation. Most patients receiving spacers are mobile during interim stages, either on crutches with toe touch and/or partial weight bearing. The incidence of wound healing complications is estimated at 8.3–28% (Koo et al., 2001, Wentworth et al., 2002). Leunig et al. (1998) reported dislocations of the hip in 5 of 12 patients, Magnan et al. (2001) a dislocation rate of 1 in 10 and Duncan and Beauchamp (1993) reported dislocations in 3 of 13 patients. Ries and Jergesen (1999), Koo et al. (2001), Shin et al. (2002) and Takahira et al. (2003) could not observe any dislocation during the spacer’s implantation (Fig. 5.6 to 5.8). Here, it should be noted that in the majority of the

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Orthopaedic bone cements Mayo Hip Score p=0.012

90 80 70 60 50 40 30 20 10 0 Infection situation

p=0.018

After infection eradication

p=0.012

Follow-up

5.5 Evaluation of hip joint function using the Mayo Hip Score in eight patients suffering from destructive bacterial coxitis or infections of the proximal femur after osteosynthesis before spacer implantation, between stages and at 1-year follow-up after prosthesis implantation.

5.6 Spacer luxation in situ.

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5.7 Asymptomatic spacer fracture with no dislocation in the lower part of the stem.

studies the term ‘dislocation’ is not clearly defined as ‘luxation’, ‘dislocation’ or even ‘fracture’, so that the precise rate of each of these complications remains unknown. Our experience shows that luxations can be reduced in a closed fashion, whereas luxations with fractures of the spacer neck or stem often require a surgical revision with spacer exchange. An approximate calculation shows rates of 8–16% for dislocations and fractures (Durbhakula et al., 2004, Hsieh et al., 2004). Furthermore, it is also unclear whether these ‘dislocations’ correlate with the way in which the spacer is fixed to the femur (cemented versus snap-fit). The period between stages varies since there are no evidence-based recommendations in place (Morimoto et al., 2003). Based on laboratory data and the progress of each infection, reimplantation was carried out after a mean period of 3–4 months. Usually, no difficulties were reported at the time of spacer removal. Reinfections either with the primary causative organism or with a new species were rare (Jahoda et al., 2003, Takahira et al., 2003, Hsieh et al., 2004). A prolonged implantation period might actually endanger the outcome of the treatment if subtherapeutic levels of antibiotic(s) are eluted from the spacer, and the antibiotic-impregnated cement itself provides an excellent environment for the development of

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5.8 Symptomatic spacer fracture with dislocation in the upper part of the spacer near the neck.

resistant bacterial strains, especially for those producing glycocalyx (Duncan and Masri, 1994, Thornes et al., 2002). Furthermore, some controversy exists regarding drainage insertion. Some authors attribute an infection persistence or new bacterial contamination to the drainage or avoid an insertion in order to maintain high local antibiotic tissue concentrations (Hofmann et al., 2005). However, what is not being taken into consideration is that the antibiotics can be eluted as long as a diffusion gradient in situ is available. If no drains are inserted, at some point the gradient is suspended since the tissue concentrations and those on the spacer surface are level. Furthermore, a drain prevents haematoma formation, which could create a medium for growth of micro-organisms. Moreover, the local antibiotic levels can be determined from the drain fluid, the efficacy of the spacer confirmed, or not and – if necessary – the complementary systemic antibiosis adjusted. With regard to infection sanitation, it is unclear how high the emergence of new resistant bacterial strains is after the use of spacers in cases of infection persistence or recurrence. In general, resistance is a potential problem when using antibiotic-loaded cement. Antibiotic-resistant organisms are found in 88% of infected total hip implants when gentamicin-loaded cement is used in the primary procedure and in only 16% of patients where gentamicin is not added (Hope et al., 1989). Perhaps, in the case of a spacer

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reimplantation, a different antibiotic combination should be incorporated into the cement than that used for the primary replacement. The overall success rates of hip replacements are reported to be 80–98% (these rates vary according to each definition of ‘reinfection’ and ‘infection persistence’). In order to achieve such high infection eradication rates and to avoid any possible haematogenous spreading of the infection, a systemic antibiosis complementary to the local antibiotic therapy is crucial. Most authors recommend an antibiosis over the first 6 postoperative weeks, consisting of 4 weeks of intravenous administration followed by 2 weeks of oral antibiosis (Durbhakula et al., 2004, Hsieh et al., 2004, Hsieh et al., 2005). Theoretically, a reduction of the antibiotic dose over the first 2 weeks might be possible, due to the locally high released concentrations, with no negative influence on the infection sanitation; however, current data are scarce regarding this treatment algorithm and further studies are clearly needed in order to clarify this.

5.10

Girdlestone or spacer?

Controversy also currently exists with regard to which of the two-stage techniques has the best clinical outcome. There is only one study, by Hsieh et al. (2004), that compared the two procedures of the two-stage protocol (Girdlestone versus spacer implantation). It was shown that the spacer implantation had a significantly lower complication rate. The patients achieved a higher hip score rate between stages, and at reimplantation there was a shorter operative time, less loss of blood and a lower transfusion requirement – indicating that the spacer is probably the superior technique in the treatment of late hip joint infections. Wentworth et al. (2002) compared their results using the PROSTALAC system with those of the literature concerning the Girdlestone procedure with regard to specific complications (persistent and recurrent infections at first and second stage). It was found that 16.1% of the patients treated by resection arthroplasty suffered from a persistent or recurrent infection, whereas this rate was 10.7% in the PROSTALAC group; however, this difference was found to be not significant. However, each orthopaedic surgeon should be aware of some limitations in the use of hip spacers. Firstly, the compliance of the patient is essential for a good functional outcome. Should the patient be non-compliant, several complications (spacer luxation, spacer fracture, femur fracture) might occur and hence endanger the outcome. Secondly, the patient should be willing to undergo a prosthesis reimplantation after spacer explantation. Patients who are multimorbid and are medically unfit to tolerate another surgery should be treated using a Girdlestone procedure rather than spacer implantation.

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5.11

New techniques

The increase in the popularity of the spacer for the treatment of hip joint infections in recent years has fostered research for new techniques for quick and simple intraoperative manufacture. Ries and Jergesen (1999) and Shin et al. (2002) recommended the use of a rubber bulb portion of an irrigation syringe as a mould to shape the proximal end of the spacer’, while Barrack (2002) manufactured different spacers by utilizing rod pins with various lengths and diameters. Such an approach would enable the surgeon to build up a stock of pre-manufactured spacers from which the appropriate shape could be selected during surgery. Ries and Jergesen (1999), Barrack (2002) and Shin et al. (2002) reported using such spacers and achieved successful eradication of infection and preservation of articulation between the spacer and the acetabulum in conjunction with satisfying mechanical strength of the prosthesis.

5.12

Conclusions

A choice between one-stage and two-stage procedures has to be made in the treatment of a hip joint infection. Should the patient undergo a twostage protocol, the implantation of antibiotic-loaded spacers is currently the method carried out by a large range of surgeons. In this chapter, we reviewed the current data available and the following conclusions are drawn. • Standardized spacers appear to be superior to hand-made spacers. • The combination of two antibiotics is advisable due to the enhanced elution of both agents and the wider antimicrobial spectrum. • In the case of an aminoglycoside–glycopeptide combination, the glycopeptide should be included at a higher dose due to its inferior release characteristics. • The total dose of the additive antibiotics should not exceed 10% of the weight of the cement in order to maintain the mechanical strength of the spacer. • The insertion of metallic components into the cement during the spacer’s manufacturing process may increase its stability, but the release of antibiotics after metal insertion has not been evaluated. • A complementary, systemic antibiotic treatment during the first weeks between stages is necessary to prevent an infection due to haematogenous spread, but it should be noted that the selected antibiotics ought to differ from the spacer’s antibiotics. • The insertion of a drain is advisable. Better data on all these issues are clearly desirable. The ideal antibiotic/ cement ratio for hip spacers is still unknown and requires quantification.

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In addition, future studies should pay more attention to possible improvements of the mixing process, which can result in an increasingly homogeneous antibiotic/cement mixture. We would also welcome further investigations into the enhanced strength of spacers by the inclusion of metal endoskeletons and the effects of different antibiotic concentrations on this enhanced strength. Nevertheless, it should be noted that, despite the importance of in vitro studies for all these parameters, clinical trials are inevitable to confirm and substantiate such experience.

5.13

References

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ries md, jergesen h (1999), ‘An inexpensive molding method for antibioticimpregnated cement spacers in infected total hip arthroplasty’, J Arthroplasty, 14 (6), 764–5. rodriguez h, ziran bh (2007), ‘Temporary antibiotic cement-covered gamma nail spacer for an infected nonunion of the proximal femur’, Clin Orthop Relat Res, 454, 270–4. schöllner c, fürderer s, rompe j-d, eckardt a (2003), ‘Individual bone cement spacers (IBCS) for septic hip revision – preliminary report’, Arch Orthop Traum Surg, 123 (5), 254–9. shin ss, della valle cj, ong bc, meere pa (2002), ‘A simple method for construction of an articulating antibiotic-loaded cement spacer’, J Arthroplasty, 17 (6), 785–7. takahira n, itoman m, higashi k, utsiyama k, miyabe m, naruse k (2003), ‘Treatment outcome of two-stage revision total hip arthroplasty for infected hip arthroplasty using antibiotic-impregnated cement spacer’, J Orthopedic Sci, 8 (1), 26–31. thornes b, murray p, bouchier-hayes d (2002), ‘Development of resistant strains of Staphylococcus epidermidis on gentamicin-loaded bone cement in vivo’, J Bone Joint Surg Br, 84 (5), 758–60. ure kj, amstutz hc, nasser s, schmalzried tp (1998), ‘Direct-exchange for the arthroplasty treatment of infection after total hip arthroplasty’, J Bone Joint Surg Am, 80 (7), 961–8. wagner h, wagner m (1995), ‘Infizierte Hüftgelenksprothesen. Gesichtspunkte für den einzeitigen und zweizeitigen Prothesenwechsel’, Orthopäde, 24 (4), 314–8. wahlig h (1987), ‘Über die Freisetzungskinetik von Antibiotika aus Knochenzementen – Ergebnisse vergleichender Untersuchungen in vitro und in vivo’, Aktuelle Probl Chir Orthop, 31, 221–6. wendt c, rüden h, edmond m (1998), ‘Vancomycin-resistente Enterokokken’, Deutsches Ärzteblatt, 95, 1284–91. wentworth sj, masri ba, duncan cp, southworth cb (2002), ‘Hip prosthesis of antibiotic-loaded acrylic cement for the treatment of infections following total hip arthroplasty’, J Bone Joint Surg Am, 84 (Suppl 2), 123–8. younger ase, duncan cp, masri ba, mcgraw rw (1997), ‘The outcome of two-stage arthroplasty using a custom-made interval spacer to treat the infected hip’, J Arthroplasty, 12 (6), 615–23. younger as, duncan cp, masri ba (1998), ‘Treatment of infection associated with segmental bone loss in the proximal part of the femur in two stages with use of an antibiotic-loaded interval prosthesis’, J Bone Joint Surg Am, 80 (1), 60–9. zilkens k-w, casser h-r, ohnsorge j (1990), ‘Treatment of an old infection in a total hip replacement with an interim spacer prosthesis’, Arch Orthop Trauma Surg, 109 (2), 94–6.

6 Commercial aspects and delivery systems of bone cements R. K O WA L S K I and R. S C H M A E H L I N G, DePuy CMW, UK

Abstract: The chapter begins by discussing the commercial aspects of mixing and delivery systems of bone cements, in hip and knee arthroplasty, together with the impact of cemented and cementless implants. This chapter includes a summary of the evolution of bone cement mixing and delivery systems and discusses the features and benefits of such systems. Key words: bone cement, mixing, vacuum, porosity, commercial aspects.

6.1

Introduction

As a result of longer life expectancy and an ageing, more demanding population musculoskeletal disorders are affecting millions of people around the world. Joint diseases, for example, account for more than half of all chronic conditions in persons aged 60 years and over; and back pain is the second leading cause of sick leave.1 Conditions such as rheumatoid arthritis, osteoarthritis, osteoporosis, and spinal disorders, and an increase in the number of serious injuries, have been identified as leading causes of morbidity and disability, costing billions in terms of healthcare expenditure and loss of earnings. According to the American Agency for Healthcare Research and Quality,2 in 2005 musculoskeletal procedures were performed in over 3.4 million hospital stays, causing aggregate costs of $31.5 billion. Knee arthroplasty, hip replacement, and spinal fusion have been the most common musculoskeletal procedures, accounting for about 1.2 million hospital stays. From 1997 to 2005, the volume of knee and hip replacements in the US rose by about 69% (from 328 800 procedures to 555 800 procedures) and about 32% (from 290 700 procedures to 383 500 procedures), respectively.2 Similar growth rates can be currently seen in Europe and Asian countries. The demand for these procedures is projected to double in the next two decades. As a result, societies and healthcare systems face huge financial burdens which need to be addressed. The level of reimbursement has changed in most societies, and healthcare providers are challenged to deliver 113

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high-quality patient care with limited financial resources. Hence, healthcare providers review internal processes and costs more closely than ever before. In this challenging economic environment, bone cement plays a vital role. Still, to date, about 90% of all knee arthroplasties and about 60% of all hip arthroplasties worldwide are fixated by bone cement. Failure of fixation causes the need for revision, which will increase total costs for the healthcare systems significantly. It is the objective of this chapter to discuss the commercial aspects, and mixing and delivery systems, of bone cement in hip and knee arthroplasty. The role of bone cement in spinal applications is not part of this chapter. While discussing the need for bone cement by comparison with cementless implants, all steps for a successful modern cementing technique will be described under commercial aspects. As the fixation of the implant also relies heavily on the mixing and delivery of the cement, the features and benefits of systems currently available will be emphasized.

6.2

Commercial aspects: cemented versus cementless fixation

It was Sir John Charnley in the 1960s who revolutionized orthopaedic surgery with his ‘low friction arthoplasty’. In seeking a method of stable fixation of the hip implant, he found a cold curing dental acrylic cement, which showed very promising results.3 A few years later using a bone cement had become the method of choice for fixation for orthopaedic surgeons around the world. Despite convincing long-term results of cemented implants, bone cement and its associated risks have always been questioned and scientists try to find other methods of fixation without relying on bone cement and related accessories. Changes in design, surface finish, and material lead to the development of cementless implants. Their fixation relies exclusively on the bony ingrowth into the implant. While cementless fixation is increasingly successful in hip arthroplasty, due to good long-term clinical results, numerous clinical studies in knee arthroplasty show either inferior or equal performance. On the other hand, cementless implants may be more expensive than their cemented counterparts and, especially in price-sensitive markets, the choice of implant may be a very commercial, non-clinical decision. In order to understand the commercial role that bone cement plays in total joint arthroplasty, one needs to understand all cost-associated factors in both the cementless and the cemented implant environment. Critical consideration must be given to four major aspects:

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clinical performance; implant costs; professional education; operating room time.

6.2.1 Hip arthroplasty Clinical performance Several clinical studies have been performed on hip arthroplasties. With a survival rate of 80% at 25 years4 the cemented option remains unsurpassed; however, the survival of circumferentially coated uncemented stems has significantly improved. In recent years, a tendency towards cementless hip implants has been seen. In a literature review5 comparing the clinical performance of cementless and cemented hip implants, no advantage was found for either procedure when revision of either one or both components was defined as the endpoint. In large subsets of study populations cemented fixation still outperforms the cementless counterpart; however, this literature suggests an improved performance of cementless implants.5 These findings are also supported by the Swedish National Arthroplasty Hip Register.6 Cementless implants (see Fig. 6.1) are especially selected for younger, still very active patients, while there is a tendency to use cemented implants (see Fig. 6.2) for older patients with osteoporotic bone quality.

6.1 Example of a cementless hip implant.

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6.2 Example of a cemented hip implant.

Implant costs While the costs for the uncemented femoral implant relate to the actual stem only, based on modern cementing techniques the cemented stem requires further products for successful fixation. 1

A plug for sealing the femoral canal. While some surgeons may use bone graft taken from the femoral head, many surgeons prefer polyethylene or biodegradable cement restrictors. Plugging improves the ability to pressurize the cement and limits the size and extent of the cement column. This increases the uniformity of the cement mantle. 2 A pulse lavage system for cleaning the medullary canal. Cleaning the cancellous bone has been proven to increase penetration of cement into the bone and has been considered critical in achieving an adequate cement interdigitation. Furthermore, the risk of pulmonary embolism can be reduced. 3 Vacuum mixing of the bone cement can decrease cement porosity and fume exposure. Porosity reduction has been well documented to increase tensile and fatigue strength in the cement, theoretically increasing the cement’s longevity. This will be discussed in more detail later in this chapter. Depending on the surgeon’s preference, bone cement can be mixed in various devices: in an open bowl, a vacuum bowl, or in a syringe. Costs vary depending on the chosen mixing device. 4 The bone cement itself. Depending on the size of the patient’s femoral canal and surgeon’s individual preferred cementing technique, the

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quantity of bone cement used may vary. While some surgeons consider a quantity of 40 g sufficient, others may choose up to 100 g for a primary hip. The quantity used can be a considerable cost factor, given the fact that antibiotic-loaded bone cement is strongly favoured by a growing number of surgeons. Antibiotic-loaded bone cement can be priced three times higher than its non-antibiotic-loaded counterpart. However, there is strong clinical evidence that antibiotic-loaded bone cement may reduce the risk of infection.7 5 Pressurization with special devices helps to increase further the interdigitation into the cancellous bone. For the cemented implants, all these steps may result in additional costs. However, due to higher manufacturing costs, cementless implants may be more expensive than cemented stems. This may still be the case even when bone cement and accessories are added to the total implant cost. It must be mentioned that the above discussion relates to the stem insertion only. If a cemented cup is chosen for the acetabulum, further mixing systems, bone cement, and pressurizers may be required.

Professional education While the implantation technique of cementless implants is focused on the stem only, the cemented procedure requires a far more complex implantation technique. After preparation, plugging, and cleaning of the medullary canal, it is the responsibility of the nurse to initiate the mixing process of the bone cement. As bone cement is sensitive to external factors such as humidity and temperature, the nurse needs to be trained well to fully understand the function of the mixing device and the behaviour of the bone cement. Any major mistake or time delay during the mixing process might result in inferior characteristics of the bone cement viscosity. If the viscosity is too low, blood may penetrate the bone cement; if the viscosity is too high, the bone cement may not be interdigitated ideally into the cancellous bone, or even worse, the bone cement may set too early in the medullary canal before the implant insertion. This may result in a time-consuming revision. However, assuming an optimal mixing process, the nurse inserts the mixing device into an application gun and hands it over to the surgeon. It is now up to the surgeon to check the viscosity of the bone cement. According to modern cementing techniques, the surgeon fills the femoral canal with the bone cement. Then, to achieve the required interdigitation, he pressurizes the cement with a special pressurization device. After sufficient pressurization, the implant must be inserted centrally to achieve a circumferent cement mantle. After insertion, the surgeon must keep the implant

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in its central position until the cement is fully set. This may take a couple of minutes in which micro-motion should be prevented. Compared with the cementless implantation technique, the cemented insertion technique is considerably more demanding. Hence, nurses and surgeons need to be well educated in order to achieve optimal fixation. Operating room time Differences in duration of the operation should be considered in explaining the differences in costs between cemented and cementless implants. In a review of the Norwegian Arthroplasty Register, the operating time for the uncemented prostheses was, on average, 15 min shorter than that for the cemented total hip arthroplasties (THAs).7 Additionally, other studies reported the cost per minute in the operating room to be between $20 and $25.8 This cost saving certainly needs to be considered when investigating the commercial aspects of implant choice in THA.

6.2.2 Knee arthroplasty Clinical performance Total knee arthroplasty (TKA) is well established for relieving pain, correction of deformity, and improving function. Long-term results, however, are often related to cemented fixation and show an excellent survivorship of over 97.2% at 14–17 years.9 Cementless implants (see Fig. 6.3) were designed as an alternative to cemented implants. The aim was to provide long-term fixation without the fear of cement debris particle generation and cement degradation resulting in late prosthetic loosening and failure. For

6.3 Example of a coating used on cementless knee implants.

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6.4 Example of a cemented tibial plateau knee implant.

younger patients in particular, they were meant to serve as a superior solution. However, critical studies revealed a unique set of complications, i.e. poor fixation as evidenced by frequent occurrence of radiolucent lines, aseptic loosening, osteolysis, and patellar polyethylene dissociation from metal-backed cementless patellar components. At the same time, the cemented counterpart (see Fig. 6.4) continues to produce excellent long-term results. In recent years, more porous as well as hydroxyapatite (HA)-coated cementless implants have been designed. Although early clinical results are promising, these devices must be measured against cemented fixation. To date, more than 90% of all TKAs worldwide are fixed by bone cement. Implant costs Clinical outcomes are very decisive for the choice of the implant. As cementless implants have yet to prove their long-term survival, surgeons tend to prefer the cemented fixation. In TKA, costs of implants may vary depending on the surgeon’s preferences. In most cases, as this is supported by clinical outcomes, the tibial component will be fixed with bone cement. The femoral component, however, may be implanted either with or without cement, depending on the surgeon’s preference. The replacement of the patella is a controversial subject and opinions vary from country to country. In the case of patella replacement, there is a strong tendency for choosing the cemented option. Compared with THA, the accessories required to achieve an optimal cementing technique in TKA are not as complex. 1

A pulse lavage system for cleaning the bone, especially the tibial plateau. Cleaning the cancellous bone has been proven to increase penetration

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of cement into the bone and has been considered critical in achieving an adequate cement interdigitation. 2 Similar to the hip, vacuum mixing of the bone cement is used to decrease cement porosity and fume exposure, and improve fatigue strength of the cement. Again, depending on the surgeon’s preference, bone cement can be mixed in various devices: in an open bowl, in a vacuum bowl, or in a syringe. Costs vary depending on the chosen mixing device. 3 The bone cement itself. Depending on the size of the implant chosen and the surgeon’s preferred cementing technique, the quantity of bone cement used may vary. While some surgeons consider a quantity of 40 g sufficient, others may choose up to 80 g for a primary knee. For the cemented implants, all these steps may cause additional costs. However, due to higher manufacturing costs cementless knee implants may be more expensive than their cemented counterparts. This may still be the case even when bone cement and accessories are added to the total implant cost. Professional education In comparison to THA, the implantation technique is not as complex for TKA. Still, nurses and surgeons need to understand the function of the mixing and delivery system used as well as understanding the behaviour of their bone cement. While the cement insertion process can occur quite quickly in THA, the challenge in TKA is quite different. Some surgeons aim to implant all three components with one mix of bone cement. This implies that the bone cement offers sufficient working time. However, some surgeons prefer to cement the tibial and the femoral component with two separate mixes of bone cement. Their argument relates to the fear that a hyperextension during the cement setting time may cause cement leakage anteriorly. Compared with the cementless implantation technique, the cemented insertion technique is not very much more demanding. However, as the working time is most crucial for this indication, the nurses in particular need to be trained well in the correct usage of the chosen mixing device. Operating room time Depending on surgeon’s preference and skills, the operating room time required for the implantation may vary. Implanting the components with two separate mixes of bone cement doubles the total amount of time required for cement setting. Furthermore, one more mixing device will need to be used.

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6.3

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6.3.1 Occupational health and risks associated with methylmethacrylate Bone cements are offered as a two-component system: a powder component and a liquid component. The powder component consists mainly of poly(methylmethacrylate) and/or a copolymer of methylmethacrylate. In addition, the powder contains a radiopaque agent (e.g. barium sulphate or zirconium dioxide), an initiator, benzoyl peroxide (BPO), and optionally an antibiotic (e.g. gentamicin sulphate or tobramycin). The liquid component mainly consists of methylmethacrylate monomer, but other methacrylates such as butylmethacrylate are sometimes used. Other ingredients include an activator (N,N-dimethyl-p-toluidine (DMPT)) and an inhibitor (e.g. hydroquinone). Once the powder and liquid are mixed together, the reaction between the initiator (BPO) and activator (DMPT) will commence through the formation of radicals, the radicals in combination with the methylmethacrylate monomer initiate a polymerization reaction resulting in the evolution of heat. At room temperature, methylmethacrylate is a liquid, it is highly flammable and has a boiling point of 100°C, a freezing point of −48°C, and vapour pressure of 35 mmHg at 20°C. The odour threshold of methylmethacrylate has been quoted as 0.08–0.21 ppm.10 Owing to its low odour threshold, methylmethacrylate odour can be noticed at levels significantly below its maximum exposure limit, therefore a strong odour of methylmethacrylate does not necessarily indicate that the worker is being overexposed to it; however, the odour is able to warn workers of the presence of methylmethacrylate before its maximum exposure limit is reached. Documented evidence has shown that some investigators had noted irritation to methylmethacrylate vapour at 170–200 ppm, although workers had tolerated 200 ppm without complaint.11 From the evidence, it was determined that 100 ppm could be tolerated continuously for 8 hours. Exposure levels of 2300 ppm were found to be intolerable for workers.12 The American Conference of Governmental Industrial Hygienists (ACGIH) recommends an 8-hour time weighted average (TWA) of 50 ppm (205 mg/m3) and a shortterm exposure limit (STEL) of 100 ppm.13 A study was carried out to measure methylmethacrylate vapour during total joint arthroplasty. In this study measurements were taken around the surgeon’s face during two consecutive operations; during each surgery two mixes of cement were used, one in the hip and one in the acetabulum. The theatre operated at 20 air changes per hour. Peak levels of methylmethacrylate measured were less than 1 ppm.14

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No carcinogenic effects were observed in several inhalation and oral animal studies15,16 and methylmethacrylate is considered not likely to be carcinogenic to humans.16,17 Some animal inhalation studies have suggested that exposure to high levels of methylmethacrylate vapour can cause foetal abnormalities.15,16 There are no adequate studies to determine whether methylmethacrylate can affect pregnancy in humans. However, theatre staff who may be pregnant should avoid overexposure to methylmethacrylate. Commonly, female hospital staff, especially in the first trimester of pregnancy, may be encouraged to leave the operating room during cement mixing and application. A further study was carried out to investigate the concentrations of methylmethacrylate in serum and breast milk from two lactating surgeons. Bone cement was mixed in a vacuum mixing system and the surgeons were exposed to typical levels of methylmethacrylate vapour found in an operating theatre, by inhalation, without the use of personal exhaust systems during total joint arthroplasty. Samples were taken from the surgeons following surgery but methylmethacrylate was not detectable, at the 0.5 ppm level, in serum or breast milk following inhalation exposure during total joint arthroplasty.18 Overexposure to methylmethacrylate is known to cause irritation of the respiratory system, and skin sensitization. Workers exposed to methylmethacrylate have complained of soreness in the eyes, nose, and throat, and of headaches. Although methylmethacrylate is an irritant, it is still considered to be a relatively safe material for implantation into the body. However, measures should be taken, such as mixing the cement under vacuum, to maintain low levels of methylmethacrylate vapour in the operating theatre to protect the patient, surgeon, and nursing staff.

6.3.2 Mixing under vacuum Sir John Charnley is generally recognized as having developed the firstgeneration cementing technique.19 Charnley realized that the cement, when implanted into bone, is a grout and not glue. In order to achieve good stability of the implant, it was essential that the cement penetrates into the porous structure of the cancellous bone and then sets achieving good fixation to the bone. A summary of the technique he used is as follows.19 The nurse mixed a high viscosity bone cement (CMW 1) in an open bowl. Methylmethacrylate monomer vapour was not contained within the open bowl, but accumulated in the theatre during preparation of the doughy bone cement. Although the surgical staff wore full theatre dress, which bore some similarity to that of an astronaut, this would not completely protect them from the monomer vapour. In operating theatres where staff did not use the

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astronaut dress, there would be even higher exposure to methylmethacrylate vapour. Once the bone cement had doughed, the surgeon would knead the cement with his thumbs and pack the doughy bone cement immediately into the femoral medullary canal. The surgeon would apply his thumbs for packing the medullary canal and for pressurizing the bone cement into the cancellous bone. Once packed with cement, the surgeon would then push the shaft of the femoral implant into the cement dough.

Owing to the uncomfortable levels of methylmethacrylate monomer in the operating theatre, during the 1970s early mixing system designs20 were primarily focused at removing and minimizing monomer fumes in the theatre, but bowls were still essentially open to atmosphere with fume extraction ports around the body of the bowl. However, the benefits of mixing the bone cement under a closed vacuum, especially with respect to reducing the porosity in the cement21 and increasing the strength of the cement,22 were finally realized. It was not sufficient for a vacuum mixing system to simply remove the fumes through the vacuum pump and to deliver them to another part of the theatre. It is generally accepted in modern vacuum mixing systems to pass the monomer fumes through an activated carbon filter, where the monomer fumes are trapped, ensuring that monomer fumes are minimized in theatre. A number of mixing systems have been developed to mix the bone cement, under vacuum, in a mixing bowl or in a syringe, some of which will be discussed in the next sections.

6.3.3 Vacuum mixing bowls For the bowl to be commercially successful, it must address user requirements. The main user requirements are fume removal, reduction in porosity, ease of mixing, and the quality of mix. However, during the development of a vacuum mixing bowl a number of other factors need to considered, such as the volume of the bowl, the method of sterilization of the plastic components, sterilization of the carbon filter, the effect of the monomer on the plastic components, contaminants in the plastic components, sealing of the bowl to vacuum, the delivery of the cement to the bone, and the influence of vacuum and temperature on the setting characteristics of the cement (examples of vacuum mixing bowls are given in Table 6.1). A study was carried out, which clearly demonstrates that mixing of bone cement, in a bowl under a closed vacuum, has the advantage of significantly reducing the level of monomer fumes in the operating theatre. In the study, which compared the Stryker bowl (Stryker), and the UltraMix bowl (DePuy) to open-bowl mixing, it was demonstrated that lowered methylmethacrylate emissions during mixing under vacuum and curing resulted in a 73–90%

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Table 6.1 Examples of vacuum mixing bowls

Bowl system

Manufacturer

Method of mixing

Comments

Stryker bowl Stryker Ultramix bowl DePuy SmartMix Bowl DePuy

Rotational Central axis Rotational Central axis Rotational Gearing mechanism, is easy to (back-and- use, with every complete forward stroke giving 1.75 mix action) rotations resulting in quicker and more thorough cement mix. The paddle is designed to fit the contours of the bowl to minimize the risk of unmixed powder Mixevac III Stryker Rotational Dual-blade design to improve mixing, bowl and blade designed for lower porosity, 2 : 1 gear ratio, ensures quick, thorough mix Fusion Biomet Vertical and Cement collection under twisting vacuum HiVac Bowl Summit Rotational Geared rotational axis mixing mechanism HiVac Multimix Summit Rotational Geared rotational axis mixing (Distributed mechanism and contraas SmartMix rotating paddle to scrape Tower by bowl and feed cement back to DePuy in the the main paddle. Transfers US) the mixed cement into a cartridge Advanced Stryker Rotational Blade design to give lower Cement Mixing porosity (with Simplex P). (ACM) System Transfers the mixed cement into a cartridge

reduction in methylmethacrylate concentrations in the breathing zone of the preparer.23 In total hip surgery, the goal of porosity reduction in the preparation of acrylic bone cement is to provide a stronger, more fatigue-resistant material between the implant and bone. Conventional mixing of poly(methylmethacrylate) bone cement produces porosity of 5–16%, whereas vacuum mixing or centrifugation reduces the porosity to a range of 0.1–3.4%.24 A vacuum bowl removes a significant amount of porosity from the cement, particularly when compared with hand mixing in an open bowl.18 Reduced porosity in the bone cement leads to a denser cement,

(%)

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9 8 7 6 5 4 3 2 1 0 Hand mixed

SmartMix Bowl

6.5 Porosity of hand-mixed cement and cement mixed in a SmartMix Bowl.25

Tensile strength (MPa)

58 56 54 52 50 48

Hand mixed

SmartMix Bowl

6.6 Tensile strength of hand-mixed cement and cement mixed in a SmartMix Bowl.25

increasing its mechanical strength (see Figs 6.5 and 6.6). This increases the amount of stress that the cement mantle can withstand, which, along with modern pressurization techniques, may contribute to the long-term fixation of the implant.25–27 The volume of the bowl is dictated by the surgical procedure. This has been discussed in Sections 6.2.1 and 6.2.2; bowls are generally designed to hold up to three mixes of bone cement (i.e. 120 g of cement powder). Common methods of sterilization include gamma sterilization and ethylene oxide (EO) sterilization. Gamma sterilization can reduce the molecular weight of plastic components, which can result in embrittlement of those components. This can potentially lead to functional failure during surgery or breakage of the mixing system components during transit. During the development phase of a new mixing system the assembled components would undergo full functional testing and transit testing to ensure that the

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components do not fail in use or during transport to the end user. Although EO sterilization does not normally adversely affect the mechanical properties of the plastic components, it can adversely affect moisture-sensitive components and will be absorbed by, and denature, the carbon filter. The carbon filter may therefore have to be sterilized separately from the plastic components of the mixing system. Methylmethacrylate monomer is a strong solvent; it causes dissolution and swelling of methylmethacrylate-based polymers, dissolution of poly vinyl chloride, and crazing in polycarbonate. The solvent effects of methylmethacrylate must be assessed against any new plastic components or seals used in the mixing system, which may come into contact with the cement or methylmethacrylate monomer, during mixing, to ensure that the function of the mixing system is not compromised or that the cement is not contaminated by solvated plastics that may in turn adversely affect the mechanical properties of the cement or risk the safety of the patient. Polymers used in injection moulding may contain certain processing additives, such as phthalate plasticizers, and these may leach out of the polymer to contaminate the bone cement. Such materials may pose a risk to the patient. Polymers for use in vacuum mixing systems should be carefully selected to ensure that they do not contain such materials and contact trails may need to be carried out to measure whether materials are transferred from the plastic components of the mixing system to the powder, liquid, or mixed cement, with which they may come into contact. The bowl base must be sealed to the bowl lid and around the agitator mechanism to help achieve an adequate vacuum within the bowl. In turn the bowl must be designed to withstand vacuum: to prevent the bowl from imploding due to atmospheric pressure, or to prevent the bowl base and bowl lid from compressing together, applying pressure to the paddle and thus preventing the paddle from functioning correctly. The US market remains dominated by medium-viscosity cement, e.g. Simplex P (Stryker) and SmartSet MV Endurance (DePuy), whereas in Europe and the rest of the world high-viscosity cements have tended to dominate (e.g. Palacos (Hereaeus Medical), CMW1RO (DePuy), and SmartSet HV (DePuy)).This difference is reflected in mixing system designs. In general, systems that are able to mix high-viscosity cements are also able to mix medium-viscosity cement. However, paddles designed for mediumviscosity cement are not necessarily designed to cut through and thoroughly mix high-viscosity cement. A number of bowls are currently marketed that utilize different styles of mixing. These include a central axis agitator (see Fig. 6.7), geared agitators to increase the speed of the agitator (see Fig. 6.8) or to deliver back-and forth mixing, rotational axis mixing, and vertical and twisting mixing (see Table 6.1). Differences in cement mixing are known to influence the fatigue properties of the bone cement.28 In order to allow the

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6.7 Example of a bowl with a central axis agitator.

6.8 Example of a bowl with a geared agitator.

theatre staff to view the contents and quality of cement mixing easily, and to identify any areas of unmixed powder, improvements have been made in bowl clarity. Once the cement has been mixed in the bowl, the surgeon must then deliver that cement to the bone. This may be achieved in one of two ways: digitally or by syringe. If the surgical technique is to apply the cement digitally, then first the bone cement must be allowed to build up viscosity before using a spatula to remove the cement out of the bowl into the palm of the surgeon’s hand. Some spatulas are designed to match the contours of the inside face of the bowl, to ensure that the majority of cement is extracted

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6.9 Example of a combined mixing bowl and syringe design.

from the bowl. The cement may then be applied digitally, e.g. to the socket of the acetabulum. If the surgical technique is to apply the cement with a syringe, then following mixing the cement may be poured (especially in the case of a medium-viscosity cement) directly from the bowl into a nonvacuum syringe (e.g. DePuy non-vacuum syringe) or the cement can be allowed to build up viscosity before being transferred into the syringe with the spatula. Two systems, ACM (Stryker) and HiVac MultiMix (Summit), use a combined mixing bowl and syringe design, primarily for use with medium-viscosity cement, to facilitate transfer of the cement from the bowl to the syringe by the use of a simple valve mechanism (see Fig. 6.9). Temperature and the level of vacuum affect the polymerization reaction. An increase in temperature results in a quicker polymerization reaction and in a faster setting time for the cement. Generally, bowls are held in the hand of the user close to the body and heat can easily be transferred from the user, through the wall of the bowl to the cement. To prevent heat transfer to the mixing bowl, some bowls incorporate a handle with which to hold the bowl while turning the mixing handle, to minimize the heat transfer from the user. The bone cement comprises a powder and a liquid component. The liquid contains an inhibitor to avoid premature polymerization during storage. Increasing the level of inhibitor in the liquid is known to

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slow down the setting reaction of the cement. However, the inhibitor requires the presence of oxygen to inhibit the polymerization reaction. Removing the oxygen by using a vacuum interferes with the function of the inhibitor and results in a speeding up of the reaction. Increasing vacuum results in a gradual speeding up of setting time.

6.3.4 Vacuum mixing syringe Vacuum mixing syringes can be described as a family of systems in which the mixing of bone cement takes place within the syringe barrel of the system. As such, there is no necessity to transfer the mixed cement from one mixing system to another, mixing and delivery of the bone cement are carried out with the same system. For the syringe system to be commercially successful, it must meet user requirements. The main user requirements are fume removal, reduction in porosity, ease of mixing, and the quality of mix. However, during the development of a vacuum mixing syringe other factors need to be considered, similar to those considered for the vacuum mixing bowl, such as: the volume of the syringe, the method of sterilization of the plastic components, sterilization of the carbon filter, the effect of monomer on the plastic components, contaminants in the plastic components, sealing of the syringe to vacuum, the ease of mixing and the quality of mix, the delivery of the cement to the bone, the level of vacuum and porosity in the cement, and the influence of vacuum and temperature on the setting characteristics of the cement (examples of vacuum mixing syringes are given in Table 6.2). Removal of methylmethacrylate fumes from the operating theatre plays a large part in the design of mixing syringe systems. Published studies suggest that most vacuum mixing syringe systems significantly reduce methylmethacrylate exposure as compared with open-bowl hand mixing.29,30 Mixing the cement in a closed vacuum is one of the means of effectively reducing cement porosity;22 vacuum mixing has become widely accepted as the best method for reducing porosity in the cement. Cement porosity has been described as microporosity and macroporosity. Vacuum mixing systems are good at removing microporosity, but some mixing systems are not good at removing macroporosity.27 Collection of cement under vacuum31 is a step to reduce the number of macropores and micropores while increasing bone cement density. In addition, it was observed that the greater the pressure reduction applied, then the lower the porosity in the cement and the higher the cement density for a given mixing system.21 The optimum level of vacuum and the optimum porosity in the cement are still under debate; if the vacuum applied to the cement during mixing is too low then significant levels of porosity may remain. However, it has been suggested that, if the vacuum level is too high, then the final cement

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Table 6.2 Examples of vacuum mixing syringes Syringe system

Manufacturer

Method of mixing

Comments

Vacu-mix

DePuy

Vertical and twisting Rotational

Unique mixing action for complete mixing Mixes high- and mediumviscosity cement

Cemvac

DePuy

Vertical and twisting

Mixes high- and mediumviscosity cement

Optivac

Biomet

Vertical and twisting

Mixes high- and mediumviscosity cement (collection under vacuum)

HighVac

Summit

Vertical mixing translated to rotational mixing

Unique mixing action for reproducible mixing

HighVac 7

Summit

Vertical and twisting

Mixes high- and mediumviscosity cement

EASYMIX

Heraeus Medical

Vertical and twisting

Mixes high- and mediumviscosity cement

Revolution

Stryker

Powered mixing

Reduced mix times, uses a power mixer Mixes high- and mediumviscosity cement

MixOr

Smith & Nephew

Vertical and twisting

Mixes high- and mediumviscosity cement

Vortex

Smith & Nephew

Rotational

In-syringe mixing; three revolutions every time the handle is turned

may suffer from excessive shrinkage. This could lead to the formation of micro-cracks, which, under loading, may lead to fatigue failure.32 Other studies33 suggest that excessive cement shrinkage can lead to the formation of voids at the cement–stem interface. Fracture of the cement mantle has been found to be part of the failure pattern in many total hip prostheses requiring revision for loosening.24 It has been concluded that voids act as stress raisers and render the cement susceptible to early failure; the use of a reduced pressure mixing system for mixing cement was found to improve the compressive properties of the bone cement by 15–30% compared with using a manual mixing system22 and to improve the bending strength by 9–33%.28 The use of a void-free

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stronger material is expected to improve the longevity and stability of the implant. Studies have demonstrated that the mean fatigue strength of acrylic bone cement mixed under reduced pressure can be enhanced over that achieved when the cement was mixed in an open bowl, but the mixing mechanism was also found to influence fatigue strength.32 In order to deliver high-viscosity cement, the syringe needs to withstand the high pressures generated by the cement as well as the high vacuum; deformity of the barrel walls, while under pressure may result in mechanical failure of the syringe wall or may prevent the piston from sealing against the walls, resulting in cement leakage. Changing the wall material, increasing the thickness of the walls, or adding strengthening rings into the design of the wall, are ways that can be used to increase the strength and rigidity of the syringe. The diameter of the syringe and the diameter of the nozzle can significantly affect the pressure generated by the syringe and can change the ease of cement extrusion. The design of the paddle is important, especially if the mixing paddle must mix high-viscosity cement; paddle strength and the ability to cut through high-viscosity cement are essential. The strength of the gun is an important factor in delivering cement. Gun designs have shown a mixed performance with regards to cement delivery.34 While Charnley continued to use the original CMW high-viscosity cement, more fluid, medium- and low-viscosity cements began to rise in popularity in other centres. These lower viscosity cements could be used to fill the femur with a syringe,35 whereas early syringe systems struggled to either mix or deliver high-viscosity cement. Some high-viscosity cements began to be pre-chilled, effectively lowering the viscosity of the cement, to facilitate mixing and delivery.36 With these changes in the use of cement, cementing techniques changed accordingly. A retrograde filling technique was adopted, whereby bone cement was gradually syringed in a retrograde fashion from the distal plug to the proximal end of the femur using a long syringe nozzle. This technique not only ensured complete filling of the femur, but also helped to prevent the entrapment of air or fat in the distal end of the femur, which could then potentially lead to embolization of the canal contents into the venous circulation during the pressurization of cement. As part of the cementing technique, the nozzle is commonly then cut to a shorter length to allow pressurization of the cement. Some vacuum-mixing syringes have the ability to break the nozzle at predetermined points in line with the cementing technique (see Fig. 6.10). In order to achieve good penetration of cement into the bone, good cement pressurization was essential; however, lower viscosity cements could not be easily pressurized digitally without loss of cement and pressure, therefore a bone plug to seal the femoral canal was used together with a proximal pressurizer to help facilitate the pressurization of the cement.

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6.10 Nozzle with predetermined break-off point.

6.11 Proximal pressurizer on syringe nozzle.

Some proximal pressurizers are designed to fit onto the syringe nozzle (see Fig. 6.11). As syringe systems have become stronger and some are now quite capable of syringing high-viscosity cements, the lessons learned from the use of lowand medium-viscosity bone cements, e.g. retrograde filling of the femur and cement pressurization, are still used today and form the basis of the modern generation of cementing techniques (see Fig. 6.12 for an example of an empty syringe system).

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6.12 An example of an empty syringe system.

6.3.5 Pre-filled systems Pre-filled vacuum mixing systems can be described as a family of systems in which the powder is pre-filled in its mixing chamber. For the pre-filled syringe system to be commercially successful, it must meet user requirements. The main user requirements are fume removal, reduction in porosity, ease of mixing, and the quality of mix. However, the additional advantages of pre-filled systems over conventional syringe systems include: reduced monomer fumes due to complete enclosure of the ampoules, either in an ampoule cartridge or in the mixing system); fewer steps (e.g. no need to fill the system with powder); less chance of spillages (powder and liquid are already in the system); less chance of being cut while breaking the glass ampoules (glass ampoules are contained); less chance of contamination of the cement by broken glass particles from the ampoule (by the use of a particle filter); reduced hardware; reduced storage; and reduced weight (examples of pre-filled systems are given in Table 6.3). During the development of a pre-filled vacuum mixing system, the factors that need to be considered are similar to those that were considered for the vacuum mixing bowl and the vacuum mixing syringe. However, the additional factor for consideration is the method of sterilization of the powder and liquid. The liquid is sterilized by ultra-filtration; once it is aseptically

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Table 6.3 Examples of pre-filled vacuum mixing systems Syringe system

Manufacturer

Method of mixing

Comments

VacuMix Plus mixing syringe system

DePuy

Vertical and twisting/rotational

Powder pre-filled into a syringe barrel Pre-filled with CMW 1, CMW 1G bone cement, CMW 3, CMW 3G or Endurance (Japan only) bone cements

Pre-filled SmartMix Cemvac

DePuy

Vertical and twisting

Powder pre-filled into a syringe barrel and liquids pre-filled into a monomer cartridge Pre-filled with SmartSet HV, SmartSet GHV, CMW 1 or CMW 1G bone cement

Cemex system

Tekres SPA

Firmly striking the device against the palm of the hand and rotating at each strike

Powder and liquid in one enclosed mixing system Pre-filled with Cemex bone cement

Cemex automated mixing

Tekres SPA

Automated mixing

Used together with the Cemex system to deliver an optimal mix

SmartMix Vacuum mixing Bowl

DePuy

Rotational (backand-forward action)

Liquid ampoules pre-filled into a monomer cartridge; powder prefilled into a bowl Pre-filled with SmartSet MV or SmartSet GMV

VacPac mixing and delivery system

Biomet

Manipulation of powder and liquid digitally

Powder and liquid vacuum packed in plastic packaging Pre-filled with Generation 4 bone cement

packed, in glass ampoules or plastic packaging, sources of energy, such as heat and gamma radiation, can cause the liquid to polymerize inside that packaging. Glass ampoules containing methylmethacrylate monomer, which are packed within a blister, are commonly sterilized by EO gas sterilization. The plastic chamber, within which the glass ampoules are housed, must allow EO gas to penetrate fully within the chamber and to escape from the

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6.13 An example of a pre-filled system.

chamber to sterilize the plastic component and the ampoule. The plastic component must not be sensitive to EO gas. Bone cement powder is commonly sterilized by either EO gas or gamma sterilization. Gamma sterilization can be used to terminally sterilize the mixing system and the bone cement powder. However, the material for the mixing system must be carefully chosen so as not to discolour and to retain its mechanical strength following gamma sterilization. In addition, the bone cement powder must not be adversely affected. With regards to EO sterilization, the EO gas must be allowed to fully penetrate within, and escape from, the syringe barrel to sterilize the powder and plastic components (see Fig. 6.13 for an example of a pre-filled system).

6.4

Regulatory aspects

With regards to the regulatory classification, these mixing systems are regulated as medical devices. In Europe the classification is taken from the Medical Device Directive (MDD 93/42/EEC annex IX); in the US, the Food and Drug Administration (FDA) has established classifications. If the device is classified as Class I or II, and if it is not exempt, a 510 k will be required for marketing. For Class III devices, a pre-market approval application (PMA) is required, unless an approved pre-existing predicate device

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(or similar product) already exists on the market. In that case, a 510 k will be the route to market. The classification is broadly based on whether the device comes into contact with the patient, or whether the device contains either plain or antibiotic-loaded cement. In Europe, empty bowls are classified as Class Ia sterile devices. Because the nozzle of the empty syringe can come into transient contact with the patient, empty syringes are classified as Class IIa devices. Pre-filled systems that contain plain bone cement, are classified as Class IIb devices, since the bone cement is implanted into the patient; prefilled systems that contain antibiotic-loaded bone cement are classified as Class III devices, since the antibiotic is released into the patient from the bone cement. In the US all empty systems are classified as Class I devices; pre-filled systems are classified as Class II devices.

6.5

Future trends

Patient body mass index and weight have been shown to be significant predictors of survivorship of primary knee arthroplasties;37 with the trend to perform total joint replacement surgery on more active or obese patients there is now a greater requirement than ever to improve the long-term performance of the bone cement. Studies have shown that improved cement performance is not only dependent on the cement’s properties, but also on the level of porosity within the cement and the consistency of cement mixing.32 Through the reduction of porosity in the cement and improved mixing, the mechanical properties and the fatigue life of the cement can be much improved. Continual improvements in vacuum mixing systems can meet the future challenges that are developing as a result of changes in patient demographics. Empty and pre-filled systems are capable of ensuring a safe environment for the theatre staff and the patient by significantly reducing methylmethacrylate monomer fumes. Improvements in the design of pre-filled systems for totally closed mixing and delivery may increasingly minimize the exposure of the theatre staff and the patient to methylmethacrylate. The use of new materials in the design of the mixing system could offer more flexibility and simplicity, making the systems easier to use. Ease of use for the theatre staff is crucial, given the fact that any mistake during the mixing process can result in suboptimal handling characteristics of the bone cement. In addition, the busy operating room environment and staff fluctuation make proper training difficult. This lack of educational time needs to be addressed by the manufacturers. Improvements in the design and action of the mixing paddle could improve cement mixing, especially for high-viscosity cements, to give a better consistency of mix. Systems with improved wall clarity would enable

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theatre staff to have full visibility of the mix, to ensure that the cement is fully mixed and homogeneous. Automated mixing systems may potentially improve the ease of use of the mixing system for the user and give improved consistency of mix. Improvements in gun design and syringe design may improve ease of delivery of the cement, especially high-viscosity cement. The future trends in mixing and delivery systems will also be dependent on new indications, such as new methods of treating osteoporosis in an ageing population (many systems have recently been developed for the relatively new procedures of vertebroplasty and kyphoplasty, but these have not been reviewed in this chapter. See Chapter 4 for a more detailed account). Some of the lessons learned from orthopaedics may find applications in these new areas.

6.6

Conclusions

Bone cement still plays a vital role, especially in knee arthroplasty where more than 90% of all knees, worldwide, are fixed by cement. With changing patient demographics, there is now a greater requirement than ever to improve the long-term performance of the bone cement and this can be greatly enhanced by the performance of the vacuum mixing system. For vacuum mixing systems to be commercially successful, they must address user requirements and show benefits both to the user and to the patient. In general, the main requirements are fume removal, reduction in porosity, ease of use and the quality of mix. Although there are many vacuum mixing systems available commercially, there is still room for improvement in the design and in materials.

6.7

References

1 lidgren l. The bone and joint decade 2000–2010. Bulletin of the World Health Organization, 2003, 81 (9), 629. 2 chaya merrill m.p.h. and elixhauser a. Hospital stay involving musculoskeletal procedures, 1997–2005. Agency for Healthcare Research and Quality, Statistical brief 34, July 2007. 3 charnley j. Anchorage of the femoral head prosthesis to the shaft of the femur. Journal of Bone and Joint Surgery, 1960, 42B (1), 28–30. 4 berry d.j., harmsen w.s., cabanela m.e., and morrey b.f. Twenty-five-year survivorship of two thousand consecutive primary Charnley total hip replacements: factors affecting survivorship of acetabular and femoral components. Journal of Bone and Joint Surgery (American), 2002, 84-A (2), 171–7. 5 morshed s., bozic k., ries m.d., malchau h., and colford jr j.m. Comparison of cemented and uncemented fixation in total hip replacement. Acta Orthopaedica, 2007, 78 (3), 315–26. 6 the swedish national hip arthroplasty register. Annual Report 2005.

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7 engesæter l.b., espehaug b., lie s.a., furnes o., and havelin l.i. Does cement increase the risk of infection in primary total hip arthroplasty? Revision rates in 56,275 cemented and uncemented primary THAs followed for 0.16 years in the Norwegian Arthroplasty Register. Acta Orthopaedica, 2006, 77 (3), 351–8. 8 salvati e.a., wright t.m. Commentary & Perspective on ‘Effects of preheating of hip prostheses on the stem-cement interface’ by Kazuho Iesaka, MD et al. Journal of Bone and Joint Surgery, 2003, 85-A (3), 421–7. 9 rodricks d.j., patil s., pulido p., and colwell jr c.w. Press-fit condylar design total knee arthroplasty, fourteen to seventeen-year follow-up. Journal of Bone and Joint Surgery (American), 2007, 89, 89–95. 10 national institute for occupational safety and health (niosh). Occupational Health Guideline for Methyl Methacrylate – September 1978. 11 spealman c.r., main r.j., haag h.b., and larson p.s. Monomeric methyl methacrylate. American Journal of Industrial Medicine 14 (4), 292–8. 12 coleman a.l. Letter to the TLV Committee from State of Connecticut, Labor Department, Occupational Health Section (March 15, 1963). 13 american conference of governmental industrial hygienists (acgih). Guide to Occupational Exposure Values. Cincinnati, Ohio: ACGIH, 2006. 14 innova airtech instruments. Occupational hygiene in hospitals – measurement of methyl methacrylate during joint replacement operations. Available at http:// innova.dk/uploads/media/Aplic_MMA_in_Hospitals_PW.pdf, 1984. 15 us environmental protection agency. Health and Environmental Effects Profile for Methyl Methacrylate. EPA/600/x-85/364. Cincinnati, Ohio: Environmental Criteria and Assessment Office, Office of Health and Environmental Assessment, Office of Research and Development, 1985. 16 us environmental protection agency. Toxicological Review of Methyl Methacrylate (CAS No. 80-62-6) in Support of Summary Information on the Integrated Risk Information System (IRIS). Research Triangle Park, North Carolina: National Center for Environmental Assessment, Office of Research and Development, 1998. 17 us environmental protection agency. Integrated Risk Information System (IRIS) on Methyl Methacrylate. Washington DC: National Center for Environmental Assessment, Office of Research and Development, 1999. 18 linehan c.m. and gioe t.j. Serum and breast milk levels of methylmethacrylate following surgeon exposure during arthroplasty. Journal of Bone and Joint Surgery (American), 2006, 88, 1957–61. 19 charnley j. Low Friction Arthroplasty of the Hip. Berlin, Springer, 1979. 20 zimmer usa inc. Device for mixing bone cement. US patent 4,015,945, 1977. 21 alkire m., dabezies e., and hastings p. High vacuum as a method of reducing porosity of polymethylmethacrylate. Othopaedics, 1987, 10, 1533–9. 22 lidgren l., drar h., and moller j. Strength of polymethylmethacrylate increased by vacuum mixing. Acta Orthopaedica Scandinavica 1984, 55 (5), 536–41. 23 ungers l.j., vendrely t., and barnes, c. Control of methyl methacrylate during the preparation of orthopedic bone cements. Journal of Occupational and Environmental Hygiene, 2007, 4, 272–80. 24 wixson r.l., Do we need to vacuum mix or centrifuge cement? Clinical Orthopaedics and Related Research, 1992, 285, 84–90.

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25 van emden e. and spierings p.t.j. Comparative Study of Various Cement Mixing Systems. Spierings Medische Techniek, BV, The Netherlands, 2002. 26 smeds s., goertzen d., and ivarsson i. Influence of temperature and vacuum mixing on bone cement properties. Clinical Orthopaedics and Related Research, 1997, 334, 326–34. 27 mau h., schelling k., heisel c., wang j.-s., and breusch s. Comparison of various vacuum mixing systems and bone cements as regards reliability, porosity and bending strength. Acta Orthopaedica Scandinavica, 2004, 75 (2), 160–172. 28 dunne n.j. and orr j.f. Influence of mixing techniques on the physical properties of acrylic bone cement. Biomaterials, 2001, 22, 1819–26. 29 schiegel u.j., sturm m., ewerbeck v., and breusch s.j. Efficacy of vacuum bone cement mixing systems in reducing methylmethacrylate fume exposure. Comparison of 7 different mixing devices and handmixing. Acta Orthopaedica Scandinavica, 2004, 75 (5), 559–66. 30 breusch s and malchau h, eds. The Well-Cemented Total Hip Arthroplasty. Heidelberg: Springer Medizin Verlag, 2005, pp. 113–18. 31 wang j.-s., franzén h., jonsson e., and lidgren l. Porosity of bone cement reduced by mixing and collecting under vacuum. Acta Orthopaedica Scandinavica, 1993, 64 (2), 143–6. 32 dunne n.j., orr j.f., mushipe m.t., and eveleigh r.j. The relationship between porosity and fatigue characteristics of bone cement. Biomaterials, 2003, 24, 238–45. 33 bishop n.e., ferguson s., and tepic s. Porosity reduction in bone cement at the cement-stem interface. Journal of Bone and Joint Surgery, 1996, 78B (3), 349–56. 34 heisel c., schelling k., thomsen m., schneider u., and breusch s.j., Cement delivery depends on cement gun performance and cement viscosity. Zeitschrift für Orthopädie und Ihre Grenzgebiete, 2003, Jan–Feb, 141 (1), 99–104. 35 waugh w. John Charnley the Man and the Hip. London: Springer-Verlag, 1990, p. 149. 36 lidgren l., bodelind b., and möller j. Bone cement improved by vacuum mixing and chilling. Acta Orthopaedica Scandinavica, 1987, 57, 27–32. 37 mulhall k.j., ghomrawi h.m., mihalko w., cui q., and saleh k.j. Adverse effects of increased body mass index and weight on survivorship of total knee arthroplasty and subsequent outcomes of revision TKA. Journal of Knee Surgery, 2007, 20, 199.

7 Wear particles and osteolysis N. PAT I L and S. B. G O O D M A N, Stanford Medical Center, USA Abstract: Osteolysis with subsequent implant loosening is considered to be the most frequent cause of revision joint replacement surgery. Osteolysis is characterized by progressive periprosthetic bone resorption secondary to the biological response to wear particulate debris. Initially, osteolysis was referred to as ‘cement disease’ since it was considered to be the result of fragmented cement. However, with better understanding of the pathophysiology of the osteolytic process, it has been observed to develop from a combination of foreign body cellular responses to wear particulate debris from various origins – metallic, polyethylene and polymethylmethacrylate – with varying contributions of these different wear debris. This chapter provides a brief overview of the basic science related to osteolysis; the mechanism of generation, consequences and clinical implications of different types of wear particulates; and the future trends in clinical and basic science research to mitigate or circumvent osteolysis. Key words: osteolysis, wear debris, joint replacement, cement disease, polyethylene debris.

7.1

Introduction

Total joint replacement for end-stage arthritis has consistently demonstrated dramatic improvement in the quality of life of the patients.1,2 However, the longevity of total joint replacement is adversely influenced by osteolysis and aseptic loosening regardless of the method of fixation or location.3–5 Osteolysis is characterized by progressive periprosthetic bone resorption secondary to the biological response to wear particulate debris. Initially, ‘osteolysis’ was referred to as ‘cement disease’ since it was considered to be the result of fragmented cement.6 However, with better understanding of the pathophysiology of the osteolytic process over time as a result of rigorous scientific research, osteolysis has been observed to stem from a combination of foreign body cellular responses to wear particulate debris from various origins: metallic, polyethylene and polymethylmethacrylate (PMMA), although the contribution of these different types of wear debris to the osteolytic process can vary. The debris may be generated from several interfaces constituting the joint reconstruction including articulating surfaces, fixation surfaces, modular component 140

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surfaces and devices used for adjuvant fixation.7 Wear debris, in prosthetic joints is predominantly generated by adhesive, abrasive or fatigue mechanisms. This debris may originate as a result of relative motion between two primary (bearing) surfaces, a primary and secondary (non-bearing) surface or two secondary surfaces. At first, osteolysis is painless; with progressive bone loss, osteolysis may manifest clinically in the form of local pain, abnormal function, subluxation or dislocation and pathological fracture at the implant site. Radiologically, osteolysis demonstrates varying extent and pattern of ballooning radiolucency. Osteolysis with subsequent implant loosening is considered to be the most frequent cause for revision joint replacement surgery. Hence, osteolysis has been the main focus of clinical research in joint replacement recently due to the complexity and higher risk of complications associated with revision surgery for periprosthetic bone loss secondary to osteolysis and implant failures. This chapter will provide a brief overview of the basic science related to osteolysis; the mechanism of generation, consequences and clinical implications of different types of wear particulates with special emphasis on cement debris; and the future trends in clinical and basic science research to mitigate or circumvent osteolysis.

7.2

Cellular cascade and mediators of osteolysis

The cellular events leading to osteolysis and subsequent loosening have been studied extensively, but ongoing research continues to identify new pathways involved in the osteolytic process. It is expected that, with increasing use of alternative bearing surfaces, the incidence and severity of wear particulate-mediated osteolysis may reduce, but the role of the cellular mechanisms involved in the osteolytic process will remain central to the evolution of osteolysis. Recent cellular and molecular biological tools for studying the tissue samples retrieved from patients with and without osteolysis have provided vital information to elucidate the basic biology of periprosthetic bone loss.8–18 Osteolysis is thought to be principally mediated by the action of proinflammatory factors released by macrophages, fibroblasts and other cells in response to particulate debris (less than about 10 microns) from prosthetic materials or bone cement (Fig 7.1).18–22 Some recent studies have identified the intermediatory role of osteoblasts and fibroblasts in the osteolytic process. It has been proposed that small wear particles, <5 microns, can be phagocytosed by the cells of osteogenic lineage to induce dose-dependent down-regulation of factors involved in bone formation.22–27 Santavirta et al.28 observed that, in cases of cemented arthroplasty with progressive osteolysis, the periprosthetic tissue demonstrated a more aggressive

142

Orthopaedic bone cements Wear particulates (<12–15 microns) Negative Positive

GMCSF

Activated macrophages, fibroblasts, etc.

Proinflammatory cytokines (e.g. IL-1, IL-6, TNF-alpha, PGE2), chemokines

Degradative enzymes (e.g. MMP-2, -9)

Osteoblasts and mesenchymal precursors

Decreased bone formation

Increased RANKL and decreased OPG Local osteoclast mediated bone resorption

Periprosthetic bone loss (osteolysis)

7.1 Flow chart demonstrating the cascade of events and the role of various cellular mediators leading to osteolysis. MMP, matrix metalloproteinase; IL, Interleukin; PGE2, prostagland in E2.

granulomatous reaction composed of multinucleated giant cells and C3bireceptor and non-specific esterase positive monocyte-macrophages. Among the numerous cellular mediators, the cytokines interleukin 1, interleukin 6, tumor necrosis factor alpha (TNF-alpha), prostanoids such as prostaglandin E2, nitric oxide and other super oxide factors are considered to be the most important components capable of inducing cell proliferation and stimulating osteoclasts to resorb the adjacent bone.12,18 TNF-alpha also stimulates osteoblasts to produce granulocyte-macrophage colony stimulating factor (GM-CSF) which up-regulates transformation of immature macrophages into osteoclasts and multinucleated giant cells.29 All these factors initiate a cascade of events that ultimately influence the final effectors of bone metabolism, receptor activator of nuclear kappa B ligand (RANKL) and osteoprotegerin (OPG). Hence osteolysis occurs as a consequence of increased RANKL and decreased OPG in the periprosthetic bone microenvironment.30–33 Recently, the role of phosphoinositol-3-kinase (PI3K)– Akt pathway in the cellular cascade of osteolysis was elucidated in an in vitro study. The authors concluded that inhibition of the PI3K–Akt pathway in response to titanium particles with adherent endotoxin and

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endotoxin-free particles reduced macrophage TNF-alpha production, which can have significant implications in addressing osteolysis by pharmacological interventions.34 Farber et al.35 found that antigen-presenting cells (macrophages) from the tissues surrounding loose cemented acetabular components expressed B7-1 and B7-2, the key molecules of the B7–CD28 (CD28: cell differentiation antigens expressed on T cells that provide co-stimulatory signals required for T cell activation) co-stimulatory pathway, indicating activation of immune response. Furthermore, metal particles complexed with specific proteins may lead to activation of specific T lymphocytes in the immune system.35 Although immunochemical studies have documented the importance of T lymphocytes and related immune responses in osteolysis, the exact role of T cells in modulating periprosthetic events remains controversial.36 In addition, several degradative enzymes, metalloproteinases, collagenases and stromelysin released by activated macrophages are implicated in the sequence of events leading to osteolysis.37–46

7.3

Morphology and bioreactivity of wear debris

Retrieval analysis of tissue specimens from failed joint replacements has provided valuable insights into the pathogenesis of wear particle-induced osteolysis. Histopathological studies of periprosthetic tissues around loose joint arthroplasties have revealed the presence of a pliant interpositional membrane composed of fibrovascular stroma, abundant macrophages and foreign body giant cells and wear debris. Depending on the materials in the implant used, particles of PMMA, polyethylene and metallic debris are found in the retrieved specimens. Fibroblasts are the most predominant cellular elements observed (70–90%) in the periprosthetic tissues. Polyethylene debris less than 1–2 microns are phagocytosed by macrophages; while larger shards (>10 microns) are surrounded by multinucleated giants cells in fibrous stroma. Smaller polyethylene wear particles are thought to be more bioreactive and produce a greater release of proinflammatory cytokines.8,47–50 Haynes et al.51 and associates noted that stainless steel and cobalt chrome molybdenum particles induced more toxic response by human monocytes, while titanium particles did not affect cell viability. However, in a recent study, Goodman et al. concluded that titanium and PMMA particles inhibit bone cell viability and proliferation, and down-regulate markers of bone formation in a dose- and time-dependent manner. PMMA particles, although initially thought to be relatively less toxic due to their larger size, can contribute to periprosthetic bone loss by inhibiting the early phases of osteoprogenitor differentiation.27,52 Wear particles from ceramic prostheses are much smaller in size, fewer in number and appear to be more benign compared with polyethylene debris.

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Orthopaedic bone cements

Shanbag and associates22,53 demonstrated that the macrophage response to particulate debris was dependent on particle size, composition and dose as given by surface area ratio. Most studies have suggested that particles that can be phagocytosed are the most stimulatory and that higher doses elicit greater response up to the saturation level. Yamamoto et al.54 reported on the mechanical cytotoxicity of murine macrophages and fibroblasts exposed to ceramic particles of varying size and shape. It was observed that cytotoxicity was influenced by the shape of ceramic particles and not by their size, with dendritic particles having higher cellular toxicity than spindle-shaped and spherical particles. Hence, it is clear that the bioreactivity of particles is dictated by the size, composition and concentration of the wear debris but it is difficult to determine which particle type (polyethylene, metal, ceramic) will have the least clinical effect.

7.4

Osteolysis

The extent and pattern of osteolysis are influenced by the effective joint space.55 Effective joint space is the periprosthetic region that is accessible to joint fluid. Periprosthetic bone resorption seems to be a self-perpetuating process, which increases with higher concentrations of biologically active wear particles circulating in the effective joint space. This leads to formation of large spaces secondary to bone resorption, encouraging further ingress of joint fluid and additional bone resorption.56 Moreover, loosening of implants may be a reflection of hydraulic effects influencing the rate of progression of bone resorption. It is clear from long-term clinical studies that longer follow-up, preferably more than 10 years, is necessary to document clinically relevant periprosthetic bone resorption representing the true nature of the problem.57 Osteolysis can result in secondary effects such as periprosthetic fractures, and the extent of osteolysis can also determine the complexity and long-term success of subsequent revision surgery. Even with substantial bone loss, patients may be asymptomatic and may not present to the clinician until bone loss leads to implant failure, loosening or periprosthetic fracture.58 Pain associated with osteolysis with or without loosening manifests as groin, buttock or thigh pain according to the location (acetabular/femoral) of the involved component of the hip replacement. The pain is often associated with loading activity and abates with rest. On physical examination, the patient with implant loosening with or without associated periprosthetic fracture may have pain with range of motion and/or shortening and deformity of the involved extremity.59 Sometimes it may be possible to piston the implant in the femoral canal.60 Earlier detection and monitoring of osteolysis can prevent deleterious complications secondary to osteolysis such as periprosthetic fracture and

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loosening. Osteolysis may be an incidental finding on late postoperative radiographs or may manifest in the form of radiographic loosening. Advanced osteolysis may present as loss of cancellous bone in regions of lesser and greater trochanters.57 Radiographic presentation of osteolysis varies according to the type and location of implant fixation (Figs 7.2 and 7.3).

7.2 Anteroposterior radiograph of pelvis showing osteolysis on the femoral as well as on the acetabular side.

7.3 Lateral radiograph of the same patient revealing femoral and pelvic osteolysis.

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Plain radiographs are often a poor measure of the extent of osteolysis. Special radiographic views are often helpful, such as Judet oblique views of the acetabulum. Three-dimensional computed tomography (CT) is a more sensitive tool to detect osteolysis at an earlier or silent phase in order to quantify its size, and evaluate its progression with time.61 Although threedimensional CT is considered to be superior to conventional radiographs, its use is limited, due to cost and exposure of the patient, to high levels of radiation making it impractical for monitoring the osteolytic process.62,63

7.5

Cement debris

7.5.1 In vitro and in vivo studies on cement debris Animal studies have demonstrated that cement particles can produce a foreign body and chronic inflammatory reaction similar to that documented around loose cemented implants in humans.64,65 In an animal study on rabbit tibia, it was observed that cement particles decreased net bone formation and increased bone resorption.64 However, polyethylene wear particles were the predominant elements inhibiting bone formation in this study. This may be attributed to the higher bioreactivity of the smaller polyethylene particles compared with cement particles.66,67 Similarly, in vitro and in vivo studies have confirmed that cement debris can induce cellular responses associated with osteolysis.68–71 Recent laboratory studies have implicated small radio-opaque particles of barium sulfate, zirconium oxide and other components of bone cement in exacerbating the osteolytic process, although this is controversial.72–79 Furthermore, a subgroup of patients exhibited an allergic reaction to N,N-dimethylparatoluidine (DMT), an accelerator present in the cement polymerization process, as determined by skin patch testing. This manifested as rapid onset aseptic loosening, often associated with osteolytic lesions. The authors recommended skin patch testing in high-risk patients such as those with previous dental or occupational exposure to DMT to avoid catastrophic early loosening of the implant.80 The role of mature cells such as osteoclasts, macrophages and osteoblasts and the effects of particulate debris leading to osteolysis and implant loosening have been well documented. However, some recent studies have also elucidated the potential influence of PMMA particle debris on the osteoprogenitor cells, thus contributing to the process of periprosthetic bone loss.81,82 The activity of osteoblasts is considered an important factor influencing implant stability as the periprosthetic tissue is undergoing constant remodeling. PMMA particles have been demonstrated to inhibit osteoprogenitor cells involved in differentiation into osteoblasts.81,82 Osteoblastic differen-

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tiation is a process that comprises an initial phase of osteoprogenitor proliferation, an intermediate phase of matrix maturation and an ultimate phase of mineralization. The differentiation of the progenitor cells is considered a critical phase in the process of osteogenesis. Furthermore, the ability of these cells to attain an optimum colony size or population with recruitment of proper cellular elements is necessary for subsequent collagen synthesis, matrix mineralization and inducing appropriate matrix signals for proper progression of differentiation. In a recent study, murine bone marrow cells were treated with PMMA particles (0.30% v/v) for a duration of 1, 3, 5 and more than 5 days, respectively. PMMA particles induced complete irreversible inhibition of osteoprogenitor cells that were exposed for a duration of 5 days or more than 5 days. However, diminished inhibition of osteoprogenitor differentiation was seen when the bone marrow cells were exposed to PMMA particles for 1 or 3 days, respectively. Additionally, conditioned medium from particle-treated cells did not suppress osteogenic development, thus showing that suppression of osteogenesis was not due to inhibitory factors secreted by the other particle-activated cells such as macrophages or monocytes in the heterogeneous bone marrow cell population. It was concluded that the early phase of osteoprogenitor differentiation is a crucial time period during which exposure to PMMA particles causes irreversible inhibition of osteogenesis, which may be due to alteration in the internal morphology of the osteoprogenitor cells produced by the PMMA particles phagocytosed by the osteoprogenitor cells.81 The effect of PMMA particles on osteclasts has also been documented. Yamanaka et al.83 established that PMMA particles activate transcription factors NF-κB (nuclear factor-κappa B) and MAP (mitogen-activated protein) kinases in osteoclast precursors. The c-Jun N-terminal kinase (JNK) pathway was observed as the principal mediator in PMMA-induced osteoclastogenesis in this study. Their data indicated an essential priming action of RANKL coupled with increased osteoclastogenic signaling with the JNK pathway in response to particulate debris resulting in enhanced osteoclast activity and osteolysis.83

7.5.2 Cement disease There is clear evidence in the literature to suggest that PMMA bone cement is satisfactory for elderly patients with low activity levels.84–86 However, concerns have been raised about higher rates of aseptic loosening of cemented implants, especially with total hip arthroplasty in young, highdemand patients. It was thought that the biomaterial properties of the cement used for implant fixation could contribute to the pathological process of loosening, which leads to use of the terminology ‘cement disease’.6 Modern cementing techniques and improvements in implant design

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characteristics have been successful, to a certain extent, in alleviating the problem of cement disease. It has been documented that failure of cement mantle is commonly associated with fractures through micro-defects in the cement developed secondary to mixing of the monomer with the polymer.87,88 The shrinkage of acrylic during polymerization also may result in gaps between interfaces, which may lead to micro-motion of the implant.89 Additionally, cement fragmentation tends to occur in focal areas surrounding the femoral component, secondary to thinness of cement mantle, concentrated stresses, cyclic loading and fatigue failure of the cement.90,91 Furthermore, cemented knee arthoplasties are loaded primarily by compressive forces, which are well tolerated at the cement–prosthesis interface; whereas in cemented hip replacements, the forces on the interface are a combination of tension, compression and shear.92 Some studies have demonstrated that the cement– bone bond at the proximal lateral and distal medial interface in total hip arthroplasty is subjected to significant tensile stress that may exceed the strength of the cement.93,94 These biomechanical and biomaterial factors related to cemented implants, especially with cemented hip replacements, may lead to fragmentation and disruption of the bone–cement interface over a period of time, generating cement debris; and ultimately contributing to osteolysis and aseptic loosening of joint implants. Furthermore, the cement debris gaining access to the space between the two articulating surfaces may contribute to three-body wear, thereby increasing the overall particulate burden.95 Radiographically, loosening of cemented implants presents as component migration, cement fractures or radiolucency encompassing a large region (50–100%) of the implant–bone interface (Fig. 7.4). The possibility of infection may be suspected, but cultures of interface tissue are negative. Hence, research continues to focus on improvements in (a) cement handling, application, and (b) components of bone cement to address the issue of loosening of cemented implants. All modern cementing techniques are directed at improving the mechanical interlock between bone and cement in order to establish a durable interface (Table 7.1).96–98 This includes reduction of porosity by vacuum mixing/centrifugation, obtaining a uniform cement mantle around the prosthesis with the aid of a centralizer and incorporation of a new catalyst or reinforcement using additives to improve the properties of the bone cement. Retrograde introduction of bone cement via a cement gun has demonstrated significant reduction in air entrapment and thereby porosity of the cement mantle. Thorough cleansing of the bone bed by the use of pulsatile lavage has been shown to improve cement penetration and mechanical shear strength. Furthermore, the use of pressurizing devices to seal and contain the cement within the femur and acetabulum are proven steps to improve cement pressurization

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Cement fracture

7.4 Anteroposterior radiograph of pelvis with failed cemented total hip replacement (THR) showing cement fracture and radiolucency at the cement–bone interface. Table 7.1 Cementing techniques: characteristic features Cement technique

Features

First generation

Manual digital packing of cement in an unplugged femur. Irrigation with bulb syringe

Second generation

Retrograde cement delivery with cement gun in a plugged femur with a bone fragment or cement distally. Use of pulsatile lavage for irrigation

Third generation

Vacuum mixing or centrifugation for mixing bone cement. Use of centralizers and cement pressurization devices

and hence interdigitation. The use of modern cementing techniques has resulted in significant improvements in implant survivorship and a reduction in the number of revision surgeries.98–102 The surface finish of femoral cemented stems has also been observed to affect the survivorship in cemented hip replacements. Recently, there has been increasing evidence to suggest that cemented femoral implants with rough surface finish (e.g Iowa stems (Zimmer, Warsaw, IN), Exeter stems (Howmedica)) are associated with aggressive bone destruction and femoral lysis leading to rapid clinical deterioration compared with smooth, polished implants. This has discouraged the use of rough-surfaced implants, making smooth, polished stems a more favorable option for cemented fixation.103,104

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7.6

Polyethylene wear debris

The most common bearing configuration used in total hip replacement to date consists of a metal femoral head articulating with ultra-high molecular weight polyethylene. Polyethylene particles are the most extensively studied and most commonly implicated wear debris in the osteolytic process adjacent to joint implants. Polyethylene wear has been observed to be influenced by a plethora of factors, which can be broadly divided into endogenous and exogenous.105–108 Endogenous factors, which are determined by the biomaterial characteristics, include the method of fabrication of the polyethylene component, the thickness and toughness of the polyethylene, the sterilization conditions, the storage environment and the shelf age of the component. The exogenous factors are commonly related to design features of the implant and articulation, and include the level of conformity or congruence of the articulation, the modularity of the prosthesis, the condition of component surface articulating with polyethylene, and surgical factors such as alignment of the articulating components, retained ‘third-body’ particulates, etc. The mechanism of generation of wear debris appears to differ in total knee replacement (TKR) and total hip replacement (THR). In THR, polyethylene wear occurs primarily as a result of micro-adhesion and micro-abrasion (Fig. 7.5).109 However, in TKR, wear occurs from a combination of rolling, sliding and rotational motions ultimately leading to delamination with formation of large flakes, pitting and fatigue failure of the polyethylene surface (Figs 7.6 and 7.7).110,111 Additionally, backside or undersurface wear has been noted with modular polyethylene tibial inserts.112

Polyethylene wear

7.5 Anteroposterior radiograph of pelvis showing polyethylene wear in an uncemented THR.

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Osteolysis

Polyethylene wear

7.6 Anteroposterior and lateral radiographs of failed total knee arthroplasty following catastrophic polyethylene wear and extensive osteolysis in the proximal tibia.

7.7 Anteroposterior and lateral radiographs of the same patient following revision total knee arthroplasty.

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Various wear patterns of polyethylene have been described in clinical and experimental studies based on location (e.g. symmetric/asymmetric), surface alteration (e.g. abrasion, burnishing, scratching), presence or absence of fusion defects or spots, and presence or absence of subsurface white bands.105,113 The magnitude of linear polyethylene wear can be assessed by using sophisticated methods like radiostereometric analysis or computerassisted, edge-detection techniques,114–116 and volumetric wear which is derived by mathematical calculation using values of linear wear and head size. Numerous approaches have been adopted to minimize polyethylene wear, such as the use of cross-linked polyethylene, alterations or improvements in the manufacturing and sterilization process, and optimizing design characteristics – for example, reducing contact stresses by increasing the congruence of the articulating surfaces.117–120 The treatment of cross-linked polyethylene with antioxidants like alpha tocopherol has been observed to reduce oxidative degradation of cross-linked polyethylene in experimental studies which can be vital in improving the service life of polyethylene in the future.121,122 Further advances include changes in the fabrication process of highly cross-linked polyethylene such as X3 HXPE, also called the second-generation, highly cross-linked polyethylene, that use a sequential irradiation and annealing process.123 This type of processing preserves the mechanical strength properties of polyethylene but also yields very low free radical content and has been observed to possess the highest survivorship in functional fatigue testing.

7.7

Metallic wear debris

Historically, metallic wear debris is considered to play a minor role in periprosthetic osteolysis. There has been renewed interest in the use of metalon-metal bearing surfaces due to their better wear characteristics and the additional potential benefits of using large heads. Wear of metal-on-metal hip prostheses has been estimated to be 10 to 40 times less than that of metal-on-polyethylene prostheses.124–126 However, the level of metallic debris is often underestimated by conventional methods (e.g. scanning electron microscopy) used to characterize metal debris in retrieval studies. Fretting corrosion at the junction of metal-on-metal modular components has been observed to be a major source of metal debris. Corrosion is an electrochemical process in which metal ions are released from implant surfaces. These metal ions form metal salt precipitates and get deposited in the aqueous environment surrounding the implant. Other sources of generation of metal particulates include sites of impingement of the neck or collar with the acetabular rim, the inner surface of modular metal-backed cups and abrasion at the bone–implant interface.127–130 Titanium debris has been implicated in the loosening of cemented implants. This is supported

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by the observation that more metal particles were produced by titanium alloy implants than with cobalt chrome implants.124 The local tissue reaction around metal-on-metal prostheses, as determined by the number of macrophages, is about one grade lower than that around metal-on-polyethylene prostheses.131,132 Numerous explanations have been proposed to justify this difference. Metal particles are significantly smaller than polyethylene particles; hence macrophages are able to store larger numbers of metal particles, thus decreasing the total number of macrophages required for storage of the particles. As a result of the small size, metal particles may enter macrophages by pinocytosis instead of by phagocytosis, leading to altered cellular responses to the particles. Furthermore, CoCr particles have been recognized to have greater potential for cytotoxicity than polyethylene particles, rendering the cells incapable of an inflammatory response. There may also be a difference in the relative proportion of metal particles that are retained locally versus distributed systemically in comparison with polyethylene particles.133 Future clinical studies and experimental studies using metal particles with clinically relevant characteristics and concentrations will help to define the contribution of metallic debris in osteolysis and aseptic loosening.

7.8

Ceramic wear debris

Ceramic bearings have demonstrated superior wear resistance compared with any other bearings owing to their greater hardness. These bearings can be polished to a much lower surface roughness and have better wetability owing to their hydrophilic nature. Ceramic debris is considered more bioinert than metal alloy debris;134 however, the relative size, shape, number, reactivity and distribution of the wear particles remains to be elucidated. Hatton et al.135 observed the presence of intracellular particle agglomerates, macrophages and interspersed areas of necrosis in periprosthetic tissue samples from ceramic-on-ceramic primary total hip replacements undergoing revision surgeries. In vitro and retrieval studies on ceramic wear debris have indicated bimodal size range of this wear debris with particles as small as 5–90 nm to as large as 0.05–3.2 microns. These two types of ceramic debris are thought to be generated from two different wear mechanisms. The smaller debris is generated during normal articulating conditions and larger particles are generated under micro-separation conditions.136 Ceramic debris may not be so bioinert as initially assumed since osteolysis has been documented in some patients with ceramic-on-ceramic bearings.137–139 However, in well-functioning prostheses, there undoubtedly appears to be less inflammatory reaction compared with that associated with polyethylene or metallic debris. Furthermore, toxicity of ions is not an issue of concern with ceramics because of their corrosion resistance.

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7.9

Future trends

Strategies to combat wear and subsequent osteolysis can be broadly divided into two major categories. The first strategy includes preventing the generation of wear particulates by improvements in implant design features and fabrication processes. The second strategy is based on the philosophy of accepting the wear particles that are produced and treating osteolysis by surgical means, or by pharmacologically modulating the inflammatory reaction and subsequent bone destruction. The new bearing configurations such as ceramic-on-metal, ceramic-onceramic, or metal-on-metal offer great potential for reduction in wear rates in future joint replacements.124,140 Substantial improvements in the longevity and performance of arthroplasties can be anticipated if the notion that wear debris alone precipitates implant loosening holds true. This might result in younger patients being operated on for replacement surgeries and possibly minimizing, if not eliminating, the need for revision surgery to a large extent in patients of all age groups. This demands a close collaborative effort among surgeons, bioengineers and tribologists. Additionally, the use of polyethylene as a bearing surface option appears promising with improvement in manufacturing accuracy in relation to geometrical form, surface finish and development of polyethylene with very low wear rates. Lastly, thirdgeneration cementing techniques coupled with possible future improvements in the physical and chemical properties of PMMA may help to increase the survivorship in cemented implants. All these changes are expected to increase the cost of implants, making sophisticated economic analysis one of the deciding factors in the utility of these future forms of prostheses. Pharmacological measures to combat osteolysis are a potentially attractive option for elderly patients where revision surgery involves substantial risk, and for younger patients who may be exposed to multiple revision surgeries during their lifetime. The pharmacological modulation of osteolysis can be achieved by affecting osteoclast function and osteoclast signaling, inhibition of the inflammatory cascade/cytokines or by local infusion of growth factors – thus providing numerous options.139,141 Biphosphonates induce apoptosis of macrophages, inhibit osteoclastic function and influence the differentiation of osteoclast precursors. Preclinical studies have demonstrated encouraging results in inhibiting particulate-induced osteolysis.142,143 Modulation of osteoclast signaling by the use of osteoprotegerin (OPG) has shown also promising results in experimental studies.144 Non-steriodal anti-inflammatory drugs may potentially influence osteolysis by inhibiting cylco-oxygenase, which is a predominant isoenzyme involved in the inflammatory cascade at the sites of local tissue injury.145,146 Etanercept, a soluble inhibitor of TNF-alpha has been shown to inhibit osteoclastic bone

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resorption in laboratory studies and hence may potentially help to reduce osteolysis.147,148 A recently introduced AMG 162 human IG2 monoclonal antibody has been observed to possess very high affinity for RANKL and may help to modulate osteolysis by its action on the OPG–RANKL pathway. Local infusion or direct application of growth factors like transforming growth factor beta (TGFβ) and basic fibroblast growth factor-2 (FGF-2) also appear to be interesting treatment alternatives due to their stimulatory action on osteoblast-like cells, given the benefit that the systemic effects of other potent agents may be mitigated.149,150 Although pharmacological therapy appears an attractive treatment option, there is no clinical data in humans that show that these agents inhibit or reverse particulate-induced osteolysis. The use of alternative bearing surfaces, such as highly cross-linked polyethylene, and third-generation cementing techniques represent some of the advances that aim to address the problem of wear debris and the resulting osteolysis associated with joint replacements. Further clinical and basic science research is warranted to help the surgeon to offer patients the best possible joint prosthesis with optimal wear characteristics, and with superior longevity and performance, and which best suits their level of activity, needs and expectations.

7.10

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92 scott wn (Ed.) (2005), Insall & Scott, Surgery of the Knee, 4th Edition, Churchill Livingstone, Philadelphia, PA. 93 svesnsson nl, valliappan s, wood rd (1977), Stress analysis of human femur with implanted Charnley prosthesis. J Biomech, 10(9), 581–8. 94 crowninshield rd, tolbert jr (1983), Cement strain measurement surrounding loose and well-fixed femoral component stems. J Biomed Mater Res, 17(5), 819–28. 95 willert hg, bertram h, buchhorn gh (1990), Osteolysis in alloarthroplasty of the hip. The role of bone cement fragmentation. Clin Orthop, 258, 108–21. 96 nedungayil sk, mehendele s, gheduzzi s, learmonth id (2006), Femoral cementing techniques: current trends in the UK. Ann R Coll Surg Engl, 88(2), 127–30. 97 markel dc, critchfield j, chaffin a, wooley pm, grimm mj (2003), Cement pressurization after provisional repair of femoral cortical defects. Am J Orthop, 32(5), 229–33. 98 beckenbaugh rd, ilstrup dm (1978), Total hip arthroplasty. J Bone Joint Surg Am, 60(3), 306–13. 99 britton ar, murray dw, bulstrode cj, et al. (1996), Long-term comparison of Charnley and Stanmore design total hip replacements. J Bone Joint Surg Br, 78(5), 802–8. 100 ballard wt, callaghan jj, sullivan pm, johnston rc (1994), The results of improved cementing techniques for total hip arthroplasty in patients less than fifty years old. A ten-year follow-up study. J Bone Joint Surg Am, 76(7), 959–64. 101 roberts dw, poss r, kelley k (1986), Radiographic comparison of cementing techniques in total hip arthroplasty. J Arthroplasty, 1(4), 241–7. 102 russotti gm, coventry mb, stauffer rn (1988), Cemented total hip arthroplasty with contemporary techniques. A five-year minimum follow-up study. Clin Orthop, 235, 141–7. 103 grose a, gonzález della valle a, bullough p (2006), High failure rate of a modern, proximally roughened, cemented stem for total hip arthroplasty. Int Orthop, 30(4), 243–7. 104 della valle ag, zoppi a, peterson mg, salvati ea (2005), A rough surface finish adversely affects the survivorship of a cemented femoral stem. Clin Orthop, 436, 158–63. 105 lewis g (1997), Polyethylene wear in total hip and knee arthroplasties. J Biomed Mater Res, 38(1), 55–75. 106 bartel dl, rawlinson jj, burstein ah, ranawat cs, flynn wf (1995), Stresses in polyethylene components of contemporary total knee replacements. Clin Orthop, 317, 76–82. 107 sutula lc, collier jp, saum ka, currier bh, currier jh, sanford wm, et al. (1995), Impact of gamma sterilization on clinical performance of polyethylene in the hip. Clin Orthop, 319, 28–40. 108 rose rm, crugnola a, ries m, cimino wr, paul i, radin el (1979), On the origins of high in vivo wear rates in polyethylene components of total joint prostheses. Clin Orthop, 145, 277–86. 109 mckellop ha, campbell p, park sh, schmalzried tp, grigoris p, amstutz hc, et al. (1995), The origin of submicron polyethylene wear debris in total hip arthroplasty. Clin Orthop, 311, 3–20.

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129 jacobs jj, skipor ak, patterson lm, mallab nj, paprosky wg, black j (1998), Metal release in patients who have had a primary total hip arthroplasty. A prospective, controlled, longitudinal study. J Bone Joint Surg Am, 80(10), 1447–58. 130 silva m, heisel c, schmalzried tp (2005), Metal-on-metal total hip replacement. Clin Orthop, 430, 53–61. 131 doorn pf, campbell pa, worrall j, benya pd, mckellop ma, amstutz mc (1998), Metal wear particle characterization from metal on metal total hip replacements: transmission electron microscopy study of periprosthetic tissues and isolated particles. J Biomed Mater Res, 42(1), 103–11. 132 doorn pf, mirra jm, campbell pa, amstutz hc (1996), Tissue reaction to metal on metal total hip prostheses. Clin Orthop, 329(Suppl), S187–205. 133 macdonald sj, mccalden rw, chess dg, bourne rb, rorabeck ch, cleland d, et al. (2003), Metal-on-metal versus polyethylene in hip arthroplasty: a randomized clinical trial. Clin Orthop, 406, 282–96. 134 christel ps (1992), Biocompatibility of surgical-grade dense polycrystalline alumina. Clin Orthop, 282, 10–18. 135 hatton a, nevelos je, nevelos aa, banks re, fisher j, ingham e (2002), Aluminaalumina artificial hip joints. Part I: a histological analysis and characterisation of wear debris by laser capture microdissection of tissues retrieved at revision. Biomaterials, 23(16), 3429–40. 136 tipper jl, hatton a, nevelos je, ingham e, doyle c, streicher r, et al. (2002), Alumina–alumina artificial hip joints. Part II: characterisation of the wear debris from in vitro hip joint simulations. Biomaterials, 23(16), 3441–8. 137 wirganowicz pz, thomas bj (1997), Massive osteolysis after ceramic on ceramic total hip arthroplasty. A case report. Clin Orthop, 338, 100–4. 138 yoon tr, rowe sm, jung st, seon kj, maloney wj (1998), Osteolysis in association with a total hip arthroplasty with ceramic bearing surfaces. J Bone Joint Surg Am, 80(10), 1459–68. 139 horowitz sm, algan sa, purdon ma (1996), Pharmacologic inhibition of particulate-induced bone resorption. J Biomed Mater Res, 31, 91–6. 140 capello wn, dantonio ja, feinberg jr (2005), Alternative bearing surfaces: alumina ceramic bearings for total hip arthroplasty. Instr Course Lect, 54, 171–6. 141 goodman sb, trindade m, ma t, genovese m, smith rl (2005), Pharmacologic modulation of periprosthetic osteolysis. Clin Orthop, 430, 39–45. 142 huk ol, zukor dj, antoniou j, petit a (2003), Effect of pamidronate on the stimulation of macrophage TNF-alpha release by ultra-high molecular-weight polyethylene particles: a role for apoptosis. J Orthop Res, 21, 81–7. 143 thadani pj, waxman b, sladek e, barmada r, gonzalez mh (2002), Inhibition of particulate debris-induced osteolysis by alendronate in a rat model. Orthopedics, 25, 59–63. 144 goater jj, o’keefe rj, rosier rn, puzas je, schwarz em (2002), Efficacy of ex vivo OPG gene therapy in preventing wear debris induced osteolysis. J Orthop Res, 20, 169–73. 145 goodman sb, chin rc, chiou ss, lee js (1991), Suppression of prostaglandin E2 synthesis in the membrane surrounding particulate polymethylmethacrylate in the rabbit tibia. Clin Orthop, 271, 300–4.

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146 zhang x, morham sg, langenbach r, young da, xing l, boyce bf, et al. (2001), Evidence for a direct role of cyclo-oxygenase 2 in implant wear debris-induced osteolysis. J Bone Miner Res, 16, 660–70. 147 childs lm, goater jj, o’keefe rj, schwarz em (2001), Efficacy of etanercept for wear debris-induced osteolysis. J Bone Min Res, 16, 338–47. 148 childs lm, goater jj, o’keefe rj, schwarz em (2001), Effect of antitumor necrosis factor-alpha gene therapy on wear debris-induced osteolysis. J Bone Joint Surg, 83A, 1789–97. 149 goodman sb, song y, chun l, regula d, aspenberg p (1999), Effects of TGF beta on bone ingrowth in the presence of polyethylene particles. J Bone Joint Surg, 81B, 1069–75. 150 goodman sb, song y, yoo jy, fox n, trindade mc, kajiyama g, et al. (2003), Local infusion of FGF-2 enhances bone ingrowth in rabbit chambers in the presence of polyethylene particles. J Biomed Mat Res, 65A, 454–61.

Part II Materials

8 Acrylic bone cement: genesis and evolution S. D E B and G. K O L L E R, King’s College London, UK

Abstract: This chapter describes the story of modern bone cements and their evolution over the last five decades. The 1960s saw the application of poly(methylmethacrylate) (PMMA)-based cement for the fixation of both the femur and acetabulum in hip replacement surgery (Charnley, 1960). Since then PMMA acrylic bone cement has gained a distinctive place in the domain of synthetic biomaterials although the composition of the cements remains essentially unaltered, newer mixing and dispensing techniques are increasingly being used to improve the performance of the cement. Furthermore, the addition of additives such as antibiotics, fluoride salts and bioactive glass fillers has been researched to enhance the clinical function of the PMMA cement. Key words: acrylic bone cements, orthopaedics, joint replacement, poly(methylmethacrylate) (PMMA).

8.1

Introduction

Arthritis is a debilitating disease that poses a major problem for the ageing population and is also known to affect the younger generation. In a normal hip joint the cartilage is responsible for articulation between bones, transmission of load across the joint and enabling smooth movements across the joint. Disease conditions of the joint – such as osteoarthritis, osteoporosis or rheumatoid arthritis – cause degeneration of the joint, the final stage of which is the destruction of articular cartilage. Since cartilage has a limited regenerative capability, the damage is usually permanent and necessitates end-stage surgical replacement of the hip. Total joint replacement has been widely adopted to treat these conditions, as it is often the only course of treatment to alleviate pain and improve the quality of life. Surgeons can replace a dysfunctional joint with a functional, long-lasting prosthesis, using metal alloys and polymeric materials. Over the last five decades, there have been numerous advances in the design, construction, technique and implantation of artificial hip joints, resulting in a high percentage of successful long-term outcomes. The number of joint replacements worldwide is rising and is expected to continue to do so, given the increase in population of the elderly and associated disease and trauma. It is estimated that more than one and a half million patients undergo hip and knee replacements worldwide. In 2000, The Swedish National Hip Arthroplasty Register 167

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reported a total of 169 419 hip replacements performed between 1979 and 1998 of which 7% were revision procedures. The UK Department of Health in 1995 estimated that about 18% of hip replacements were revision cases. The survival rates of hip and knee replacement surgery deem them as successful procedures; however, more detailed data presented in the Norwegian Charter suggest that the survival rates are lower for secondary hip replacements than for the primary procedure. Revision surgery is more complex, more expensive, involves higher risk and needs a longer rehabilitation time; thus there is a need to enhance the survival rates of joint replacement surgery, which will be greatly beneficial not only to improve the quality of life of patients but also to make it viable for the younger patient. The reasons for failure of hip replacements are numerous and can be associated with inappropriate technique, due to the materials and technology applied, and/or poor post-operative care. The materials and technology applied are crucial to the success of joint replacement and some known problems are those associated with particulate debris from the prostheses, stress shielding due to improper transfer of mechanical loads causing changes in the distribution of bone mass, defects in fixation and/or infection.

8.1.1 History Arthritis is a debilitating disease that causes a tremendous amount of pain, hence surgeons have been trying for over a century to treat this condition and alleviate the associated pain. It became clear early on that surgery was an effective way to address these joint-related problems and the earliest recorded attempt at replacing a diseased joint was carried out by T. Gluck (1891) in Germany, who used ivory to replace the femoral head. In 1925, M. N. Smith-Petersen, a surgeon from Boston, USA used a piece of moulded glass in the shape of a hollow hemisphere to provide a smooth new suface to fit over the ball of a hip joint to allow smooth movement. The glass was found to be cytocompatible; however, it failed rapidly as it could not withstand the normal physiological stresses. Other materials were also explored, including plastics and metals, and these were used in mould arthroplasty in the 1940s. In the UK, hip replacements using a metal prosthesis were first attempted at the Middlesex hospital by Mr Philip Wiles; however, the major breakthrough came after the Second World War, with the efforts of Sir John Charnley. Sir John Charnley pioneered one of the greatest surgical advances of the twentieth century, the modern total hip arthroplasty. Sir Charnley was born in Bury, Lancashire in 1911, and gained his medical qualifications from the Victoria University of Manchester. He qualified in 1935 and also acquired a BSc in anatomy and physiology. He was the youngest surgeon

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8.1 Charnley hip.

to become a Fellow of the Royal College of Surgeons with residencies at Salford Royal Hospital, King’s College London and Manchester Infirmary. In 1958, Charnley directed his efforts to the development of hip surgery and invented the low-friction hip replacement in the early 1960s at the Centre for Hip Surgery at Wrightington, UK. He spent the next two decades refining the replacement technique and training surgeons. The idea of using a poly(methylmethacrylate) (PMMA)-based bone cement to anchor hip and knee replacements to bone was developed by Sir John Charnley and Professor Dennis Smith in the late 1950s. They coined the term ‘low-friction arthroplasty’ to emphasize the small diameter of the prosthetic head, 22 mm, which is essential to the underlying theory of the Charnley prosthesis (Fig. 8.1). Charnley continued his efforts to find methods of replacing the femoral head and acetabulum of the hip effectively and designed a Teflon cup with a metal ball component that was expected to simulate the smooth surface of the articulating joint; however, polyethylene was soon used to replace Teflon due to its poor function. By 1961 hip replacements were being carried out by Charnley with good results. He used a PMMA-based cement that was used by dentists for denture bases. Since then joint replacement has undergone numerous changes with regards to surgical techniques, materials used for prostheses, design of prosthesis, cementing techniques; however, the basic composition of the PMMA bone cement remains unaltered.

8.2

Hip and knee joint

In a total hip replacement (THR) an artificial hip joint is surgically implanted. There are two components of the hip joint: (a) the acetabulum or the cup that replaces the hip socket and is generally made out of ultrahigh molecular weight polyethylene, or highly cross-linked polyethylene, metals and ceramics (Fig. 8.2) or a combination thereof and (b) the metal or ceramic head attached to a metal stem that is inserted into the femur to provide stability to the prosthesis. The components can be placed with or without a grouting agent, commonly referred to as the ‘bone cement’, and are grouped in two categories: ‘the cemented hip’ or the ‘uncemented hip’. Knee joints also come in two forms: an uncemented and cemented prosthesis. The prosthesis consists of two parts and is implanted onto the ends

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8.2 Components of a hip replacement. Patellar component

Femoral component

Tibial bearing component

Tibial tray

CMMG 2003

8.3 Components of a knee replacement.

8.4 Acrylic bone cement.

of the femur, the tibia and the patella with or without bone cement (Fig. 8.3). Metal implants are generally used on the femur end with polymers such as polyethylene used on the tibia and patella surface; however, combinations of metal-on-metal, ceramic-on-ceramic and ceramic-on-plastic are in common use.

8.3

Acrylic bone cement

Bone cements have been traditionally used in placing hip and knee joints (Fig. 8.4). The cement fills the space between the implant and the joint and

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is introduced as a flowable mix, which hardens with time. The fixation of THRs with PMMA bone cements is currently regarded as the gold standard (Havelin et al., 2000). The main functions of the bone cement are to transfer body weight and service loads from the prosthesis to the bone, and the immediate immobilization of the prosthesis. The PMMA cement performs admirably due to its range of properties. PMMA bone cement provides adequate fixation of the femoral stem for up to 10–15 years, but beyond that time, failure is practically unavoidable, even using the latest methods (including femoral canal preparation, and improved cement mixing and delivery techniques). The loosening of the prosthesis can be related to many factors including sepsis, interfacial failure, bone remodelling and/or mechanical failure of the cement mantle. Failure of a joint replacement surgery is a multifactorial phenomenon and thus it is often difficult to identify one particular reason for failure. Since the application of PMMA in hip replacements, it has also been used for joint reconstruction, cranioplasties, investment of aneurysms, fixation of pathological fractures, artificial eyeballs and bone replacements. Bone cements are self-curing systems and are mixed prior to insertion and introduced at the surgical site in a doughy state that undergoes in situ polymerization. Acrylic bone cement is currently the most widely used biomaterial for anchoring cemented arthroplasties to contiguous bone; however, it does have some well-known disadvantages. The reaction is exothermic and generates temperatures in the range of 66–120°C. The high rise in temperature is often a cause of necrosis and impairment of blood circulation, and is one of the reasons for fibrous tissue formation around the bone–cement interface. MMA can also cause hypotensive effects that may induce adverse systemic effects if it enters the bloodstream. If incomplete polymerization occurs, unreacted monomer may leach into the surrounding tissues leading to chemical necrosis. Furthermore, PMMA is a brittle material with low fracture toughness and poor fatigue life; there is also a mismatch of the modulus between the cement and the bone, which can contribute to failure.

8.3.1 Evolution Acrylic bone cement has been used extensively in orthopaedic surgery for several decades. Essentially, the bone cement is dispensed as a powder and a liquid, which on mixing yields the cement due to the polymerization of the monomer, MMA, in the mixture. Cements are generally characterized in terms of their chemical composition, curing kinetics and physicomechanical properties.Throughout the last five decades, efforts have focused on improving bone cements with a view of enhancing mechanical properties, mixing methods and handling characteristics, and developing new formulations and delivery methods.

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The early cements were packaged as a powder, PMMA, in a sterile pouch and the liquid, MMA, was in brown glass ampoules (Fig. 8.4). Hand mixing of the components was usually carried out in a polypropylene bowl; the requisite amounts of powder and liquid were added, the mix was stirred for 40–120 s and the cement paste introduced into the bone cavity manually. One major disadvantage of this technique was the exposure of surgical staff to MMA vapour, which has a typical odour. MMA is a respiratory and skin irritant and may act as a sensitizer through skin contact; typical exposure limits lie in the range between 20 and 50 ppm. Although the monomer is not highly toxic, it can cause problems through continued exposure. MMA is not biocompatible, although PMMA is biocompatible, hence for a brief period during the curing phase the cement can cause adverse reactions. Thus, bone cement manufacturers concentrated on developing closed mixing systems that would minimize MMA exposure. Manual mixing also causes entrapment of air, which can inhibit the polymerization process and allows air bubbles to be incorporated in the matrix resulting in porosity. In addition, manual mixing lacked reproducibility with cement mantles of uncontrolled micro- and macroporosity being formed. Therefore, centrifugation methods were attempted where a hand-mixed cement dough is placed in a syringe, then promptly placed in a centrifuge and spun at a maximum speed of between 2300 and 4000 rpm for a time period between 30 and 180 s (Jasty et al., 1990; Trieu et al., 1994). Some manufacturers also introduced mechanical mixing under vacuum and some advocated a combination of centrifugation and vacuum mixing. A number of proprietary systems were also developed for closed mechanical mixing under vacuum using paddles and closed bowls. One such system is the Stryker ACM System (Stryker, Kalamazoo, MI), which provides a convenient and effective method of mixing cement under vacuum. The ACM system mixes the cement in a cartridge, minimizing the exposure to harmful fumes. Mixevac III (Stryker) represents an advanced, simple and universal mixer for all types of bone cements with a unique blade design that ensures a quality mix and reduces porosity for all bone cements. The EASYMIX® vacuum cementing system (Heraeus, Germany) is an effective system that is easy to use; it consists of an EASYMIX® vacuum pump, an EASYMIX® cement gun and various cartridge sets. Other proprietary systems include: the Simplex Enhancement Mixer (Howmedica, Rutherford, NJ), the Stryker High Vacuum System (Stryker), MITAB (Mitab Corp., Sjobo, Sweden), Optivac (Mitab) and Sterivac (SD, Germany). There are no generic steps in vacuum mixing, and the powder is added to the liquid in a mixing pot, which is then placed in the vacuum chamber. The MITAB vacuum system has a mixing chamber that can be then closed and a vacuum of 28 kPa (absolute) can be applied, while the cement constituents are being manually

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mixed with a spatula at 0.25 Hz for between 30 and 75 s. The pot is then removed from the chamber and lid, the nozzle applied and the cement is dispensed through a cement gun at atmospheric pressure (Lewis, 1997). When the Simplex Enhancement Mixer is used, the powder is added to the liquid in the mixing chamber. The mixture is then stirred manually with a special spatula until the powder is saturated. A vacuum of 15 kPa (absolute) is drawn into the chamber, while stirring at 1 Hz is carried out for 90 s. Research findings show that the cement mixing method has a marked effect on the physical and mechanical properties of bone cements; however, the variation in literature data related to fatigue and fracture toughness of bone cements (Lewis, 1997, 1999; Deb, 1999) is largely attributed to the difference in porosity of the cements.

8.3.2 Current status Aseptic loosening of a cemented arthroplasty can result from a large number of factors, including the failure of the cement mantle; although this has not been firmly established, however, failure of the cement mantle has been implicated in the failure of joint replacement surgery, thus the bone cement has always been treated with a certain degree of ambivalence. As a consequence, much research has focused on the development of arthroplasties that obviate the use of bone cements via the placement of an implant in direct contact with bone. It was predicted that uncemented arthroplasties would improve the life-time survival of joint replacements, since the implant is anchored through a press fit fixation, achieved either through micro-/macromechanical interlocking or a bony ingrowth, and no cement used. Despite concerns about the cement, the success rates of cemented hip and knee arthroplasties are exceedingly high and the fact that the clinical outcomes of uncemented joint replacements have identified a new range of problems means that there is a sustained interest in improving the clinical performance of acrylic bone cement. The fact that the composition of acrylic bone cements has primarily remained unchanged over the last 50 years, despite the well-established deleterious effects, can be attributed to four main factors: (a) the adequate performance of the PMMA bone cement; (b) surgeon familiarity; (c) the long and arduous process of launching a new product; and (d) the risk of experimenting with new materials in comparison to a material with an adequate history. Hence, the chemical composition of bone cements has largely been restricted to PMMA-based polymers with MMA as the monomer and an activator–initiator system that uses N,N-dimethyl ptoluidine and benzoyl peroxide. Although it is now known that the tertiary amine, N,N-dimethyl p-toluidine is genotoxic with undesirable side effects, only one commercial cement composition uses an alternative amine.

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Similarly, the radiopaque agents in commercial bone cements, namely zirconia and barium sulphate, both known to cause adverse biological effects, are still being used in commercial cements despite several other alternatives such as bismuth and iodinated derivatives. Nonetheless, additives with therapeutic effects in bone cements are being increasingly researched; this is discussed in detail in Chapter 14. Currently, there are approximately 70 or more commercial bone cements available, marketed by orthopaedic, medical device or specialist companies. The main differences between the cements are the molecular weight of the pre-polymer namely PMMA, inclusion of other copolymers of PMMA, the amount of homopolymer and copolymer, the radiopacifier and additives such as chlorophyll in Palacos®. The cements can also vary in terms of viscosity, inclusion of antibiotics and, more recently, additives such as sodium fluoride, strontium, etc. The use of PMMA cement has been advocated for many diverse medical applications, in addition to use as a grouting agent for fixation of implants in joint replacements, for example, in vertebroplasty and kyphoplasty, as bone defect fillers, for craniofacial and maxillofacial reconstruction; the cements have provided satisfactory functional and cosmetic results. PMMA has been demonstrated to be biocompatible and easy to shape in vivo, allowing its use as a bone substitute in reconstructive surgery. However, one of the principal disadvantages of PMMA cements is the lack of a direct contact between bone and implant, which results in fibrous tissue formation at the interface. This non-adhesiveness causes bone resorption and loosening of the implant. In order to overcome these problems, bioactive cements that are able to integrate with bone have received a great deal of attention in the last decade. The two main types of cements are the bioactive bone cements (see Chapter 15) and calcium phosphate bone cements (see Chapter 10). Bioactive cements, as referred to in the literature, are acrylic-based cements with bioactive fillers such as apatite – wollastonite (AW)-glass ceramic, Bioglass® and hydroxyapatite (Saito et al., 1994; Kobayashi et al., 1997; Shinzato et al., 2000; Deb et al., 2005). These composites were developed in a bid to enhance the biocompatibility, mechanical properties and bone bonding properties of orthopaedic bone cements. Alternative polymer matrices using higher molecular weight monomers, such as Bis-GMA (adduct of bis-phenol A and glycidyl methacrylate), or urethane dimethacrylates with ceramic fillers, have been reported in an attempt to overcome the brittle reinforcement of the PMMA matrix. Although most of the properties of these cements are superior in comparison to PMMA cement, there is very little evidence of their clinical use to date. The lack of clinical application of newer cements can be attributed to two main factors: firstly, the handling properties of composite cements are inferior in comparison to PMMA cements and secondly, the long-established use of PMMA cements

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provides the surgeon with good clinical data, whereas the newer systems have little to show in terms of past clinical history.

8.4

Regulatory perspectives

The Food and Drug Administration (FDA) class medical devices in three categories: (a) Class I (general controls); (b) Class II (special controls), and (c) Class III (premarket approval–special controls, labelling requirements, post-market surveillance and may have training requirements). PMMA bone cements were first regulated as drugs rather than devices as they were in use prior to the Medical Device Amendments 1990. These products were classified as ‘transitional’ devices and subsequently all transitional devices were classified into the Class III category. In 2002 the FDA reclassified PMMA bone cement intended for use in arthroplasty procedures of the hip, knee and other joints for the fixation of prosthetic implants to living bone from Class III to Class II devices (special controls) (Medical Devices; Reclassification of Medical Devices; Reclassification of Polymethylmethacrylate (PMMA) Bone Cement; Federal Register, 2002). PMMA bone cements comprise a number of chemicals, thus manufacturers are required by the FDA to list the detailed composition with the quantities for each formulation (Demian and McDermott, 1998). Nearly all commercial formulations contain PMMA, MMA, benzoyl peroxide, a tertiary amine – generally N,N-dimethyl p-toluidine, hydroquinone as an inhibitor and a radiopaque agent. The formulations may differ in the comonomers of PMMA, the radiopaque agent, additives such as pigments and, most importantly, the powder : liquid ratio or the amounts of each of the components. It is also mandatory that the physical, chemical and mechanical properties of the cements are described and that they are in accordance with the ISO/ASTM standard requirements. Any new or modified cement to be launched also requires approval by the FDA, highlighting the changes in physical, chemical and mechanical performance. The curing kinetics of cements differ and depend on several parameters such as the powder : liquid ratio, particle size, the amount of initiator, initiator : activator ratio, the details of which are described in Chapter 9; these parameters also need to be part of the analysis. The FDA also requires the preclinical mechanical properties of the hardened cement. Thus, to summarize, detailed descriptions of the assembly and application, chemical composition, molecular weight, degree of polymerization, physical, thermal and mechanical properties – along with the stability of components and clinical data – are required for the approval of a new bone cement. However, not all modifications to an existing bone cement require clinical data for obtaining pre-market approval. There are instances when a minor modification of an existing cement need not require detailed clinical data; however, such

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modifications are evaluated by the FDA to ascertain the need for additional data. There are some well-documented risks associated with the use of acrylic bone cements; however, they have been used clinically with a great measure of success over the last few decades. The lack of local biocompatibility during the setting of the cement can cause bone necrosis and regional damage can occur if MMA leaching occurs. Cardiovascular reactions to acrylic bone cement have also been recorded and it is thought that cementation may increase the blood pressure during general or spinal anaesthesia (Svartling et al., 1986; Esemenli et al., 1991). The clinical success of these cements, however, should not preclude the development of safer and more effective cements, but proper evaluation and testing is imperative for any new cement.

8.5

Sterilization of bone cements

As acrylic bone cements are inserted in the body, the components need to be sterile and the mixing should also be conducted under sterile conditions. Acrylic bone cements are dispensed as a liquid and a powder, which are mixed to yield a polymerizing dough that sets with time. The monomer MMA is a polymerizable moiety that is sensitive to heat, light and chemical activation, hence it is expected to be altered if conventional methods of sterilization are used. The sterilization techniques have also evolved over the years and there is a trend to use technologies that are less damaging to the pre-polymers present in the powder component. It is customary to subject the liquid MMA to ultrafiltration, which is then sealed in dark glass ampoules in order to avoid any premature onset of polymerization. However, the pre-polymerized powder of PMMA is typically subjected to γ-sterilization and more recently other methods such as the use of βirradiation or ethylene oxide gas have been used. Irradiation is a convenient technique and requires only a short period of time, and γ-radiation is a well-established and effective technique for sterilization of implants; however, this can also result in alteration of the properties of the polymer. On the other hand, ethylene oxide gas does not cause damage to the polymer but usually the procedure is time consuming and it can take several days for degassing of powder specimens. There is a sustained interest in enhancing the properties of acrylic bone cements and thus its clinical performance; however, the effect of different sterilization methods on PMMA bone cements is relatively less well researched. PMMA is a polymer that is expected to undergo chain scission when subjected to strong radiation such as γ-rays, resulting in a lower molecular weight mix post-irradiation. It is well known that a decrease in the molecular weight of a polymer is detrimental to the mechanical

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properties, hence it is an important factor in the clinical performance of the bone cement. Fatigue failure has been deemed to be one of the causes of implant loosening, therefore, the evaluation of both the static and dynamic mechanical characteristics of bone cements is important. Lewis and Mladsi (1998) reviewed the literature on the effect of the method of sterilization of acrylic bone cements to highlight a rather neglected aspect, but one that has a significant clinical impact. A systematic study (Lewis and Mladsi, 1998) on Palacos® R, a commercially successful bone cement, confirmed that γ-sterilization caused a decrease in the molecular weight of the PMMA powder and significantly lowered the fatigue performance. However, no significant differences in the quasi-static tensile and compressive properties were reported, which may be attributed to the fact that tests were carried after ageing specimens over a short period of time. The fact that a 66% drop in molecular weight was observed is evidence that it can cause detrimental effects on all mechanical properties over time. The fracture toughness of the cements is also affected by the method of sterilization and is found to degrade with γ-sterilization (Lewis and Mladsi, 1999). Another study by Harper et al. (1997) reported a decrease in the molecular weight of the pre-polymers in experimental bone cement powders containing PMMA and PMMA with MMA/Styrene copolymers when subjected to effective doses of γ- and β-irradiation for sterilization. The fatigue life of the cements tested in tension at 2 Hz loads was also reported to decrease significantly when irriadiated with 25 kGy radiation. The authors also studied the rheological time profiles during curing and indicated that the complex viscosity was reduced in cements that used irradiated PMMA powders. Graham et al. (2000) reported the degradation in molecular weight resulting from γ-irradiation, which decreased fracture resistance significantly when compared with ethylene oxide sterilization and non-sterile samples. A corresponding decrease in fatigue resistance was also reported in the cements that were treated by a radiation dose of 10 Mrad, whereas ethylene oxide sterilization did not result in a significantly different fracture resistance when compared with unsterilized controls for vacuum-mixed cement. However, for hand-mixed cements, fracture and fatigue resistance were found to be independent of sterilization method. This observation was related to the higher porosity in hand-mixed cements that compromised the mechanical properties and masked any effect of sterilization. The authors suggest that a combination of non-ionizing sterilization and vacuum mixing is expected to yield the best mechanical properties leading to enhanced longevity in vivo. In a different study, Lewis (1999) also confirmed that the depreciation in molecular weight did significantly decrease the fracture toughness of the cements. Current literature pertaining to the evaluation of mechanical properties is now taking into consideration the effects of sterilization. One of the

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reasons for the superior performance of Palacos® and the fact that it is being considered as the ‘gold standard’ for acrylic bone cements can be attributed to the fact the cement is subjected to ethylene oxide; however, the composition also does play a role!

8.6

Fluoride and other additives in bone cements

Fluorine is known to have an effect on bone metabolism (Rich and Ensinck, 1961; Earnes and Reddi, 1979) and sodium fluoride is known to have a direct stimulating effect on the proliferation of osteoblasts and also enhances bone matrix synthesis (Farley et al., 1983). The bonding of sodium fluoride to bone mineral, namely hydroxyapatite, leads to the formation of fluorapatite (Larsen and Thorsen, 1984), which is more resistant to osteoclastic resorption in comparison with hydroxyapatite, yielding a more stable matrix (Rich and Ensinck, 1961; Posner et al., 1963). Improvement of the fixation of the implant to bone is vital in the success of a cemented implant; thus it has been hypothesized that the addition of sodium fluoride to acrylic bone cement may lead to a more stable interface via the substitution of calcium with fluoride in the hydroxyapatite crystals, which would lead to an improvement in the quality of the bone– cement interface and thereby reduce the risk of loosening. The inclusion of sodium fluoride in acrylic bone cements and its effect have been studied by several researchers and one bone cement (Cemex®) that has undergone clinical evaluation (Digas et al., 2005, 2006) is available commercially. The clinical evaluation of Cemex® (Tecres S.p.A, Italy), a bone cement containing sodium fluoride, was carried out by following the micromotion of the stem using radiostereometry and bone mineral density using manual and automatic dual-energy X-ray absorptiometry (DEXA) at a 2-year follow-up in a group of 97 patients with no significant differences reported in the subsidence and rotation of the stem (Digas et al., 2005, 2006). In a different in vitro and in vivo study, Magnan et al. (1994) also reported similar findings. Fatigue is one of the main causes of failure and additives may alter the fatigue life of bone cements, thus an experimental bone cement containing NaF (12 wt%) was investigated by Minari et al. (2001); these authors reported no significant differences in the fatigue life of the novel bone cement in comparison with the cement without the additive when tested in accordance to current standard requirements. Furthermore, the effect of sodium fluoride in an acrylic bone cement was used to evaluate fixation of titanium implants using a rabbit model with bone deprived of oestrogen (Sundfeldt et al., 2002). The investigators reported that the addition of fluoride in cements resulted in increased area of bone formation around implants in ovariectomized rabbits; however, the

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same effect was not observed in rabbits that were not oestrogen deficient, indicating that the presence of additional fluoride may be beneficial in certain cases. Another bioactive bone cement based on PMMA, consisting of bioactive glass beads as an inorganic filler and high-molecular-weight PMMA (hPMMA) as an organic matrix, was reported by Shinzato et al. (2003). The main purpose was to impart bioactivity to the PMMA cements through incorporation of bioactive glass beads that consisted of fluoridated glass (MgO–CaO–SiO2–P2O5–CaF2). The inclusion of calcium fluoride was expected to contribute to fluoride release and enhance bone formation. The in vivo study reported that histologically, new bone had formed along the surface of both the composite cements containing the bioactive glass within 4 weeks, but the fluoride incorporation did not lead to a significant enhancement in the formation of new bone. Other additives in bone cements include Bioglass, growth hormones and bisphosphonates; these have been assessed in laboratories but have not reached clinical evaluation.

8.7

Other applications of acrylic bone cement

PMMA is a synthetic polymer approved by the FDA for specific human clinical applications such as the bone cement. Bone cements based on PMMA have been widely used in joint replacement surgery and are currently beginning to expand beyond vertebroplasty, cementoplasty of fractures and metastatic lesions in sites ranging from the acetabulum to the shoulder. PMMA particles are also being researched as potential vaccine delivery agents and for skin and soft tissue augmentation. Vertebral body fractures are one of the most common complications of osteoporosis and a large number of patients are affected by this condition. The incidence of vertebral body fractures is higher in women and can cause chronic pain, functional impairment and poor quality of life. Vertebroplasty is indicated in compression fractures and is described in detail in Chapter 4. PMMA cements are used in other surgical procedures and have recently been approved by the FDA for procedures such as vertebroplasty and kyphoplasty. Vertebroplasty and kyphoplasty are minimally invasive surgical procedures developed to manage osteoporotic vertebral compression fractures. The procedures involve injection of the cement and PMMA is one of the cements used, other types of cements used include calcium phosphates. The specific requirements for kyphoplasty and vertebroplasty mean that the cement must be highly radiopaque, easily injectable into the collapsed vertebral body, have a low curing temperature and adequate mechanical properties that allow immediate reinforcement and ensure early

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ambulation of the patient. Thus, as the clinical applications change, the requirements for the cements differ and it is well established that the performance of the cement in providing stabilization in compression fractures is dependent on the flow characteristics, and physical and mechanical properties. Cranioplasty involves the repair of a bony defect or deformity in the skull. It is an area that continues to pose a challenge to surgeons and the quest for an ideal biomaterial continues. The main challenge lies in accurately defining the defect and restoring the bony defect. It is expected that the material used will be easily formable, compatible with bony tissue, be a non-conductor of heat and also have a low density. PMMA has been a favoured material for such purposes. However, materials that act as better osseoconductive bone substitutes are also currently being used. The advantage of using PMMA is its easy formability and fairly inert characteristic with a proven record of use in other surgical procedures. It is also easy to prefabricate and the accurate transfer of data regarding the defect can lead to very effective prostheses. Current methods of charting the defect include computed tomography (CT) scans (Lee et al., 1995; Chiarini et al., 2004), radiographs, radiographic models, stereolithography (Klein et al., 1992) and other computer-aided techniques (Lee et al., 1995). PMMA-based materials find wide use as biomaterials in several clinical applications, for example, sacroplasty (Heron et al., 2007).

8.8

Conclusions

The indications for total hip and knee arthroplasties are expanding in clinical orthopaedics. As a result of the increase in the ageing population and increases in sports-related trauma, the number of patients undergoing joint replacement surgery is on the rise. The majority of the prostheses are being fixed in place using PMMA bone cement, which acts as a grouting agent between the implant and bone. The presence of bone cement aids in the transfer of stress from the prosthesis to the bone and also helps in early mobilization of the patient. The most frequent long-term complication in hip arthroplasties is loosening of the joint prosthesis. However, the exact cause, other than sepsis, is not clear, but in many cases is related to failure of bone cement. Acrylic bone cement is not without its own drawbacks and has often been related to the failure of arthroplasty. Despite the drawbacks, the survival possibilities of recently implanted hips and knees are very high and average at least at 90% survival after 15 years. Factors that contribute to the greater success rate are improved cementing techniques, implantation methods and surgical procedures, and the inclusion of antibiotics.

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8.9

181

References

charnley j, Anchorage of the femoral head prosthesis to the shaft of the femur, J Bone Joint Surg, 42B; 28–30, 1960. chiarini l, figurelli s, pollastri g, torcia e, ferrari f, albanese m, nocini pf, Cranioplasty using acrylic material: a new technical procedure, J CranioMaxillofac Surg, 32; 5–9, 2004. deb s, A review of improvements in acrylic bone cements. J Biomater Appl, 14(1); 16–47, 1999. deb s, aiyathurai l, roether ja, luklinska zb, Development of high-viscosity, twopaste bioactive bone cements, Biomaterials, 26; 3713–3718, 2005. demian hw, mcdermott k, Regulatory perspective on characterization and testing of orthopedic bone cements, Biomaterials, 19; 1607–1618, 1998. digas g, karrholm j, thanner j, Addition of fluoride to acrylic bone cement does not improve fixation of a total hip arthroplasty stem, Clin Orthop Relat Res, 448; 58–66, 2006. digas g, thanner j, anderberg c, kärrholm j, Fluoride-containing acrylic bone cement in total hip arthroplasty. Randomized evaluation of 97 stems using radiostereometry and dual-energy X-ray absorptiometry, J Arthroplasty, 20(6); 784–792, 2005. earnes ed, reddi ah, The effect of fluoride on bone mineral apatite, Metab Bone Dis, 2; 2, 1979. esemenli bt, toker k, lawrence r, Hypotension associated with methylmethacrylate in partial hip arthroplasties. The role of femoral canal size, Orthop Rev, 20; 619– 623, 1991. farley jr, wergedal je, baylink dj, Fluoride directly stimulates proliferation and alkaline phosphatase activity of bone-forming cells, Science, 222; 330, 1983. graham j, pruitt l, ries m, gundiah n, Fracture and fatigue properties of acrylic bone cement: the effects of mixing method, sterilization treatment, and molecular weight. J Arthroplasty, 8; 28–35, 2000. harper ej, braden m, bonfield w, Influence of sterilization upon a range of properties of experimental bone cements, J Mater Sci: Mater in Med, 8; 849–853, 1997. havelin li, engesaeter lb, espehaug b, furnes o, lie sa, vollset se, The Norwegian Arthroplasty Register: 11 years and 73,000 arthroplasties. Acta Orthop Scand, 71; 337–353, 2000. heron j, connell da, james slj, Coccygeal fracture and clinical dentistry. CT-guided sacroplasty for the treatment of sacral insufficiency fractures, Clin Radiol, 62; 1094–1100, 2007. jasty m, davies jp, o’connor do, burke dw, harrigan tp, harris wh, Porosity of various preparations of acrylic bone cements, Clin Orthop Rel Res, 259; 122–129, 1990. klein hm, schneider w, alzen g, voy ed, gunther rw, Pediatric craniofacial surgery: comparison of milling and stereolithography for 3D model manufacturing, Pediatr Radiol, 22; 458–460, 1992. kobayashi m, takashi n, tamuro j, kokubo t, kikutani t, Bioactive bone cement: comparison of AW-GC filler with hydroxyapatite and beta-TCP fillers on mechanical and biological properties, J Biomed Mater Res, 37; 301–313, 1997. larsen mj, thorsen a, A comparison of some effects of fluoride on apatite formation in vitro and in vivo, Calcif Tissue Int, 36; 690, 1984.

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lee c, antonyshyn om, forrest cr, Cranioplasty: indications, technique and early results of autogenous split skull cranial vault reconstruction. J Cranio-Maxillofac Surg, 23; 133–142, 1995. lewis g, Properties of acrylic bone cement: state of the art review. J Biomed Mater Res Appl Biomater, 38; 155–182, 1997. lewis g, Apparent fracture toughness of acrylic bone cement: effect of test specimen configuration and sterilization method, Biomaterials, 20; 69–78, 1999. lewis g, mladsi s, Effect of sterilization method on properties of Palacos® R acrylic bone cement, Biomaterials, 19; 117–114, 1998. magnan b, gabbi c, regis d, Sodium fluoride sustained-release bone cement: an experimental study in vitro and in vivo. Acta Orthop Belg, 60; 72, 1994. medical devices; Reclassification of medical devices; reclassification of polymethylmethacrylate (PMMA) bone cement; Federal Register, July 17, 67(137), 2002. (Rules and Regulations), pp. 46852–46855. From the Federal Register Online via GPO Access, wais.access.gpo.gov (DOCID:fr17jy02-12). minari c, baleani m, cristofolini l, baruffaldi f, The effect on the fatigue strength of bone cement of adding sodium fluoride, Proc Instn Mech Engrs, Part M, J Engng Medicine, 215; 251, 2001. posner as, eanes ed, harper ra, zipkin i, X-ray diffraction analysis of the effect of fluoride on human bone apatite, Arch Oral Biol, 8; 549, 1963. rich c, ensinck j, Effect of sodium fluoride on calcium metabolism of human beings, Nature, 191; 184, 1961. saito m, maruoka a, mori t, sugano n, hino k, Experimental studies on a new bioactive bone–cement–hydroxyapatite composite resin, Biomaterials, 15; 156–160, 1994. shinzato s, kobayashi m, mousa wf, kamimura m, neo m, kitamura y, kokubo t, nakamura t, Bioactive poly(methylmethacrylate) based bone cement: comparison of glass beads, apatite and wollastonite containing glass–ceramic and hydroxyapatite fillers on mechanical and biological properties, J Biomed Mater Res, 51; 258–272, 2000. shinzato s, nakamura t, kawanabe k, kokubo t, PMMA-based bioactive cement: effect of CaF2 on osteoconductivity and histological change with time. J Biomed Mater Res B Appl Biomater, 65; 262–271, 2003. sundfeldt m, persson j, swanpalmer j, wennerverg a, karrholm j, johansson cb, and carlsson lv, Does sodium fluoride in bone cement affect implant fixation. Part II: Evaluation of the effect of sodium fluoride additions to acrylic bone cement and the fixation of titanium implants in ovariectomized rabbits, J Mater Sci: Mater Med, 13; 1045–1050, 2002. svartling n, lehtinen am, tarkkanen l, The effect of anesthesia on changes in blood pressure and plasma cortisol levels induced by cementation with methylmethacrylate, Acta Anaesthesiol Scand, 30; 247–252, 1986. trieu hh, paxson rd, carrol me, bert jm, A comparative study of bone cement preparation using a new centrifugation mixing technique, In Proceedings of the 20th Annual Meeting of the Society for Biomaterials Boston, MA, p. 416, 1994.

9 Poly(methylmethacrylate) bone cement: chemical composition and chemistry B V Á Z Q U E Z L A S A, Institute of Polymer Science and Technology (CSIC) and CIBER-BNN, Spain

Abstract: This chapter describes the chemical composition, setting process and curing parameters of poly(methylmethacrylate) bone cements. The polymerisation reaction is reviewed by using differential scanning calorimetry and electron spin resonance. The rheology of the acrylic bone cement from its early life is discussed. This chapter also discusses polymerisation shrinkage from different points of views and the low molecular weight residuals that remain trapped in the cement after curing. The substitution of the activator with other lower toxicity alternatives and the preparation of cements based on alternative components is also reviewed. Keywords: poly(methylmethacrylate) bone cement, setting process, curing parameters, electron spin resonance, rheology.

9.1

Introduction

For the last 50 years bone cements have been used clinically with encouraging results in orthopaedic surgery for the fixation of artificial joints.1 The cement acts as a grouting agent providing the immobilisation of the prosthesis in joint replacement surgery. The main function of the bone cement is to transfer stress from the implant to the bone and to increase the loadbearing capacity of the system. The main advantages of the cemented technique lie in the excellent primary fixation between bone and implant and, consequently, in the faster recovery of the patient. In addition, this is a technique with low damage rates that is straightforward to use as bone cement pastes are easily moulded and adapted to the complex bone cavities. However, the technique does present some disadvantages, mainly resulting from the high temperature reached during the exothermic polymerisation, which can cause necrosis in the surrounding tissue. Thermal necrosis has been considered to be a factor leading to later failure of the cemented arthroplasties. In addition, release of unreacted monomer, shrinkage of the cement during polymerisation and poor mechanical properties can contribute to final aseptic loosening. It is also important to consider the toxicity of the aromatic tertiary amines used in these formulations to activate the polymerisation process. 183

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9.2

Chemical composition

Commercially available bone cement formulations are based on two components, a solid and a liquid phase, which are packaged separately and are mixed in situ in the operating theatre, prior to use.

9.2.1 Solid phase The solid phase of the early cements mainly comprised spherical particles (beads) of poly(methylmethacrylate) (PMMA); however, currently marketed cements may contain low proportions of acrylic copolymers containing ethyl acrylate or methyl acrylate or even methylmethacrylate (MMA)–styrene copolymers in certain trademark cements.2 The prepolymer is present at a concentration of approximately 80 wt% and it is prepared by suspension polymerisation. In this process, the monomer containing the radical initiator, e.g. benzoyl peroxide (BPO), is dispersed in an aqueous medium in which it forms droplets, which are stabilised by a surfactant agent under mechanical agitation and at a reaction temperature around 60–80°C. The stirring rate strongly affects the particle size distribution and the average diameter of the particles. Usually the particle size of commercial beads ranges from 30 to 150 μm. It has been demonstrated that the particle size distribution and average diameter of the particles have a strong influence on curing parameters.3 Average molecular weights of the beads are reported to be in the range from 44 000 to 1 980 000 and residual monomer content is of the order of 0.28 wt%.4 The solid phase also contains the initiator, BPO, in a concentration ranging from 0.75 to 2.5 wt%. The initiator can be incorporated within the polymer particles or can be physically mixed. Another component of this phase is the radiopaque agent, namely barium sulphate or zirconium dioxide, that is incorporated as an X-ray contrast agent to allow X-ray follow-up of the prosthesis. The concentration of the radiopaque agent is generally around 10 wt%; however, some formulations contain up to 15 wt%, and the average diameter of the radiopaque particles is around 5 μm. It is a very common characteristic of inorganic compounds to form aggregates, and is also observed in bone cements with zirconium dioxide exhibiting more agglomeration compared with barium sulphate.5

9.2.2 Liquid phase The liquid phase consists of MMA monomer at a concentration of 95 wt% for the majority of the formulations, although a few formulations contain small fractions of butylmethacrylate (BMA). It is customary to incorporate N,N-dimethyl-4-toluidine (DMT) as the activator, in a concentration range

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of 0.89–2.7 wt%; however, a recent commercial formulation has an alternative tertiary aromatic amine, namely, 2-(4-dimethylamino)phenyl ethanol. Finally, an inhibitor is added to the liquid phase to avoid any premature polymerisation that may occur in the presence of heat or light during storage. The inhibitor belongs to the family of quinones, of which the most frequently used is hydroquinone; this acts as a radical scavenger forming radicals that are stabilised by resonance. An important parameter of acrylic bone cement formulations is the solid : liquid ratio, which is 2 : 1 in the majority of the cements. In some formulations, however, the solid : liquid ratio varies and may range from 2 : 1 to 2.7 :1. Several studies have been conducted to analyse the influence of this ratio on the cement properties, and, generally speaking, a ratio of 2 : 1 has been accepted as the most suitable for cements used in arthroplasty replacements.4 Recently, the influence of the variation of solid : liquid ratio (from 1 : 1 to 2.5 : 1) and increasing amounts of amine and BPO (1–2 wt%) on polymerisation shrinkage were reported for two commercial bone cements.6 The study was made using the ‘bonding disc’ or ‘Watts’ method and the results clearly indicate the possible adverse effects of changing the solid : liquid ratio in terms of producing either excessive shrinkage or under-polymerisation. On the other hand, it has been shown that the initiator : activator ratio has a significant effect on setting time, polymerisation temperature and strength.7 Bowen and Argentar reported that the maximum polymerisation rate of MMA occurred in monomer solutions containing about 1.5 moles of peroxide per mole of amine.8 However, this ratio can range from 0.5 to 2.5 in commercial formulations. The latest studies on this topic have been performed using the commercial cement Surgical Simplex® P. As BPO is incorporated into the MMA-co-styrene copolymer in the powder, to increase the BPO content of the formulation tested, the method used was to increase the copolymer content at the expense of the prepolymerised PMMA content. They reported an optimum ratio of the MMA-co-styrene copolymer + BPO to that of the activator to be 57.14 (80 wt% copolymer + BPO per 1.4 vol.% DMT).9 Monomers are sterilised by ultrafiltration since they are sensitive to other sterilisation techniques. Solid components are sterilised by ethylene oxide, and β- and γ-irradiation. It has been observed that the irradiation methods significantly decrease the molecular weight of the polymer particles, whereas ethylene oxide has no influence on it. Therefore, the latter is the preferred method for acrylic bone cements since there is no change in the properties of the material.10 This variable is important since it has been demonstrated that the molecular weight of the prepolymer affects other properties of the final set cement. Studies on experimental cements11 have demonstrated that both β- and γ-irradiation of the powder produce a decrease in a number

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of static mechanical properties of the cements, a large reduction in fatigue resistance and a decrease in complex viscosity during curing. These findings have been confirmed for commercial bone cements. γ-Sterilisation of the cement powders lead to a significant reduction in the molecular weight of the prepolymers decreasing the fatigue performance of the cements compared with powders sterilised with ethylene oxide.12

9.3

Setting process

The preparation of the cement starts when both solid and liquid phases are mixed in a bowl or in a proprietary mixing device usually at room temperature. On mixing the two phases, the monomer begins to solvate the surface of the prepolymerised beads producing a paste-like consistency; the viscosity increases with time, so that in a few minutes the mix becomes a viscous mass and then attains a rigid state. This process has been differentiated into four steps, namely: sandy, stringy and doughy followed by setting into a rigid mass.4 More recently, the setting process has been classified into four phases called the mixing phase, the waiting phase, the working phase and the hardening phase.2 The physical processes involved throughout the setting of the cement are wetting of the prepolymerised particles by the monomer, the subsequent swelling and partial dissolution of the PMMA particles into the monomer, diffusion of the liquid into the organic powder and monomer evaporation. From a chemical point of view, when both components are mixed, the reaction between the initiator and the activator starts giving rise to primary free radicals, which initiate the free radical polymerisation of the monomer. It is rather difficult to adjudicate each event to a particulate stage or phase of the setting process. However, one can expect that, during the sandy or mixing step the wetting process predominates, during the stringy or waiting phase the radical polymerisation commences that causes the viscosity to increase, during the working phase the progress of the polymerisation provokes a reduction in the mobility of the macromolecular species and a sudden increase in viscosity and heat generation occurs, and finally, in the setting step the cement hardens with the subsequent cessation of the polymerisation due to vitrification of the material. The final material is a composite formed by PMMA particles partially dissolved and embedded in the polymerised matrix produced during the polymerisation of the monomer MMA. During the polymerisation process of acrylic bone cement formulations there is a density change when the liquid monomer, with a density of 0.937 g/ml, converts to a polymer of higher density (1.18 g/ml). This event leads to what is called ‘volume shrinkage’ and is considered to be of great concern in orthopaedic applications. Pure PMMA exhibits a volume

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shrinkage of approximately 21%; however, the presence of prepolymerised powder in acrylic bone formulations reduces the theoretical value by up to 6–7%. The real shrinkage is lower due to porosity generated by the air bubble inclusions during the dough stage, release of absorbed air during polymerisation, vaporisation of monomer and air absorbed during insertion of the metal stem. In fact, two different types of pores are defined in polymerised bone cements: macropores (with a diameter, d > 1 mm) and micropores (0.1 < d < 1 mm).13 Volumetric water displacement methods were used in the measurement of shrinkage from samples polymerised in water, and values around 2.7% were reported. For samples polymerised in constant pressure moulds, the values oscillated around 5%.14 Other authors considered a diametral shrinkage method where the cement is polymerised about a metal rod simulating a femoral hip prosthesis.15 Recently, Gilbert et al.,16 using a dilatometric method in conjunction with density measurements using Archimedes’ principle, reported values of polymerisation shrinkage between 5 and 6.67% for commercial bone cements. They also carried out a theoretical and experimental analysis of polymerisation shrinkage in a constrained deformation state and demonstrated that, under those conditions, porosity can be developed due to polymerisation shrinkage at the bone cement–implant interface and also in the polymerising cement mantle. Apart from being implicated in loosening of the prosthesis, polymerisation shrinkage has been related to the mechanical performance of the bone cement. A theoretical model has been developed by Orr et al.17 to calculate interference stresses on the basis of thermal and total shrinkages. They concluded that cracks observed around hip prosthesis stems in laboratory specimens were due to shrinkage and the residual stresses were sufficient to cause crack initiation prior to functional loading.

9.4

Polymerisation reaction and kinetics

The main chemical reaction that takes place during the setting of the acrylic bone cements is the free radical polymerisation of the monomers present in the liquid phase. The reaction mechanism is well documented in the literature18 and consists of three well-described stages: initiation, propagation and termination. In the first place, the radical generation starts by the reaction between activator and initiator. The initiator is decomposed by the activator at room temperature to produce the primary free radicals. It is generally accepted that the reaction between DMT and BPO produces benzoate radicals and aminomethyl radicals, both of which can react with the monomer molecules to initiate polymerisation. This reaction mechanism has been studied in-depth by Feng19 in the polymerisation of vinyl monomers and the formation of an intimate pair or cyclic transition between the amine and the peroxide was reported, which eventually gives an aminium

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radical and a benzoate radical. The aminium radical produces an aminomethyl radical and benzoic acid by a proton transfer reaction. This mechanism has been proposed for the initiation of acrylic bone cements.20 Once the primary radicals are formed, they react with monomer molecules in the initiation step. The addition of new monomer molecules enables the propagating species to grow in the propagation step. It should be considered that each macromolecular chain is completely formed within a small fraction of a second. Finally, during the termination step, two growing radicals react by combination or disproportionation through β-hydrogen atom transfer, to give inactive macromolecular chains. The polymerisation mechanism can be described as shown below. •

Radical generation d I + A ⎯K⎯ → R•

•

Initiation i → RM• R•+ M ⎯K⎯

•

Propagation K

p → RM n+1• RM n •+ M ⎯⎯

•

Termination by coupling or disproportionation i RM n•+ RM m• ⎯K⎯ → Inactive polymer chains

Kd is the decomposition rate constant; Ki is the initiation rate constant; Kp is the propagation rate constant; and Kt is the termination rate constant for coupling and disproportionation. The rate of polymerisation is given by: Rp = Kp(fKd[I]/Kt)½[M]

or

Rp = K[M]

If a steady state is assumed, then: −d[M]/dt = K[M] where [I] is the initiator concentration, f is the initiator efficiency, [M] is the monomer concentration, K is the overall rate constant and Rp is the overall polymerisation rate. The bulk polymerisation of the bone cement is an exothermic reaction and the addition of new monomer molecules provides 130 cal/g of heat released, then a fast and highly non-isothermal polymerisation takes place. Bulk polymerisation of MMA is affected by diffusion mainly at high conversions as a consequence of the increase in viscosity. Then the reduction in mobility of the growing chains results in a drastic increase in the polymerisation rate providing the ‘autoacceleration effect’ or ‘gel effect’. It has

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been demonstrated that the addition of prepolymerised PMMA and fillers accelerates the onset of the gel effect.21 The polymerisation reaction of acrylic bone cements has been extensively studied by differential scanning calorimetry (DSC).22 The use of this technique allows the study of the polymerisation kinetics. The polymerisation conversion (αt) can be expressed as a function of time as follows: αt(%) = 100([M]0 − [M]/[M]0 = 100Qt/QT where [M]0 is the initial monomer concentration, Qt is the reaction heat corresponding to the partial area under the curve at time t, and QT is the total heat of polymerisation. Assuming first-order kinetics with respect to the monomer concentration, the polymerisation rate can be calculated from the slope of the linear curve that results when plotting ln ([M]0/[M]) versus reaction time. An Arrhenius relationship can be assumed between the rate constant for the polymerisation of the cement (K) and its temperature during the polymerisation, T, so that the following relation can be equated as: K = K0 exp(−Ea/RT) where K0 is the frequency factor, Ea is the activation energy and R is the molar gas constant (8.314 J/mol K). Some researchers have found that the polymerisation reaction of acrylic bone cement under non-isothermal conditions follows approximately firstorder reaction kinetics and the plot of ln K versus 1/T is a straight line with values of activation energy between 257 and 140 KJ/mol and ln K0 between 90 and 49 s−1 when the heating rate oscillates between 5 and 20°C/min.23 Recently, the polymerisation kinetics of commercial bone cements modified in terms of reducing the activator concentration and the prepolymerised powder content, have been analysed by calorimetry under non-isothermal conditions, reporting values of activation energy in the range from 263 to 211 kJ/mol, values of ln K0 between 92 and 75 s−1, and polymerisation rates between 5.3 × 10−2 and 7.1 × 10−6 s−1.24 However, polymerisation kinetics of acrylic bone cements deviates from a first-order reaction after initial conversion due to the ‘autoaccelaration effect’ and attainment of a maximum conversion of MMA, beyond which no significant reaction occurs despite the fact that some unreacted monomer remains. Thus, a number of models have been proposed, taking into account all these facts. Maffezzoli25 studied the isothermal and non-isothermal polymerisation of commercial bone cement and proposed a simple phenomenological model which considers the ‘autoacceleration effect’, the diffusion controlled termination mechanism and the reaction between inhibitor and initiator. Values of ln K0 of 18.8 s−1 and Ea/R of 6600 K were reported for Simplex® P in this work. Another kinetic model has been

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reported considering the material thicknesses, the initial temperature between the components and the ability of the different materials to dissipate heat.26 In this work values of ln K0 and Ea/R of 9.15 s−1 and 3763 K respectively were reported for CMW 1 (DePuy©). Other kinetics studies on the free radical polymerisation of MMA, using the amine N,N dimethyl-4-aminophenethyl alcohol (DMPOH) as an activator,27 have been performed with BPO and polymerisation kinetics were determined for varying BPO/amine molar ratios and reaction temperatures. The highest polymerisation rate was observed for a BPO/amine ratio of 1.5, but final conversion was found to be independent of the molar ratio. The activation energy of polymerisation reported for the new BPO/DMPOH system was 51.8 kJ/mol, which compared well with the value of 47.1 kJ/mol obtained for the traditional BPO/DMT system.

9.5

Free radical studies on acrylic bone cements

Electron spin resonance (ESR) or electron paramagnetic resonance (EPR) has been used to study the free radical polymerisation of MMA during the curing process of acrylic bone cements. When MMA was polymerised using the DMT/BPO system in the presence of PMMA powder,28 the conventional nine-line ESR spectrum for the growing polymer radical was detected at the gel state of polymerisation. The optimum free radical concentration was observed near the equimolar amine/BPO concentration, but an excess amount of activator with respect to BPO was found to inhibit the polymerisation process. The use of OH-modified accelerators produced a significant increase in the free radical concentration, indicating its higher efficiency in the MMA polymerisation than the currently used DMT. ESR studies performed on radiolucent and radiopaque surgical PMMA revealed comparable free radical concentration curves for both types of cements, as reported by Park et al.29 The free radicals present in the radiopaque bone cement powder produced in the radiation sterilisation process were unaffected by the polymerisation process and were stable for temperatures lower than approximately 100°C.30 Turner31 has investigated the kinetics of decay in PMMA bone cements, performing thermal annealing at various temperatures (70–88°C), and found that the logarithm of the concentration of radicals varied linearly with time, but with non-zero intercept. The analysis of the results by a first-order decay process provided an activation energy value of 39 kcal/mol, probably due to a diffusion-limited termination. The analysis based on a second-order process gave a value of 36 kcal/mol and it was attributed to bimolecular termination. Other authors32 found that samples of cement cured in saline at 37°C exhibited first-order decay kinetics for the polymerising radicals for approximately 1 week after mixing, and they attributed the decay to a transfer process. A decrease in

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the temperature of the components prior to mixing resulted in smaller decay rates, but still first-order decay kinetics were observed. Other variables such as the solid : liquid ratio have been studied, showing that larger solid : liquid ratios resulted in shorter characteristic growth time of the free radicals associated with the polymerisation33 as a result of the greater concentration of initiator present in the polymerising mass. The use of low-frequency EPR spectrometers, in the range of megahertz,34 permits an increased wave penetration (several centimetres) into tissues which enabled the observation of free radical production during the curing of PMMA bone cement in vivo.35 The amount of polymerisation radicals was measured non-invasively over days on the same animals. The decay rates of the radicals were significantly lower for the curing process occurring in vivo compared with the in vitro process. The decay period can last up to 4 weeks for in vivo experiments. These results confirmed other previous studies on the same subject.36 The use of this technique could be applied to larger samples or even to humans.

9.6

Curing parameters: standards

The polymerisation reaction of MMA is an exothermic reaction resulting in a temperature increase during curing. The ASTM F451-99a37 and ISO 583338 standards recommend monitoring the exotherm by recording the temperature–time profile using a specific procedure. Briefly, the time is recorded from the onset of mixing of components and when the mass reaches its dough stage it is inserted in a Teflon mould and the temperature recorded by a thermocouple placed in the centre of the mould at a height of 3.0 mm in the internal cavity. The curing parameters that characterise the polymerisation reaction are: doughing time, maximum temperature and setting time. Maximum temperature, frequently called peak temperature (Tmax), is obtained from the maximum of the exotherm, and corresponds to the maximum temperature attained by the bulk during polymerisation. The setting time (Tset) is defined as the time taken to reach a temperature midway between ambient and maximum, and doughing time is the time elapsed between the beginning of mixing until the mixture is able to separate cleanly from a gloved finger. Another important parameter that describes the handling characteristics of the cement is the working time which has been defined as the difference between both setting and doughing times. This parameter is not defined in the standards but it is very useful in acrylic bone cement literature.39 The standards recommend that dough time be measured using a ‘nonpowder surgical glove (latex)-covered finger’. However, it has been recently reported that the brand of the surgical glove used for the measurement has enormous influence on the final results.40 A study using four different brands

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of gloves with different surface morphology as determined by scanning electron micrographs showed that the dough time varied depending on the brand and the difference was attributed to the different surface finish. The brand that gave the shortest value had a fairly smooth surface, whereas the highest value of dough time was recorded for the roughest surface. This study makes it evident that there is a need to describe the type and brand of the glove used when dough time is reported. This finding may also have repercussions when the cement is applied clinically. For the determination of curing parameters, the ASTM F451-99a37 and ISO 583338 standards state that the mixing of both components is carried out by the traditional method that uses a spatula and a bowl (first generation). However, in practice, the mixing devices vary from the simplest traditional method to more sophisticated ones; many of them are patented and commercialised (third generation). They can be classified41 into different categories – namely, manual mixing, centrifugation mixing, vacuum mixing and combined mechanical mixing – and the mixing method may have a marked influence, mainly on the mechanical properties of the acrylic bone cements.42 Recently, the variation in the curing characteristics as a function of the mixing technique employed to prepare the bone cement has been assessed.43 In this study, two new curing parameters were introduced, cure temperature and cure time, which were defined as the values at which the cement reaches about 90–100% cure. The main conclusion obtained was that not all types of bone cement are suited to a particular mixing system and the authors warn of the need of a full assessment on a mixing system before it is introduced in the market. Other investigations elucidate the influence that ambient temperature has on the determination of setting time.44 It seems that there is a complex relationship between both parameters coming from the polymerisation process but also from the swelling and dissolution of the polymer particles in the liquid monomer. From values for the activation energies of both swelling and dissolution processes, it was concluded that the sensitivity of setting time to ambient temperature depends more on swelling and dissolution than on the polymerisation process. The mould for recording the exotherms recommended by ASTM F45199a37 and ISO 583338 standards has also been questioned in a recent paper.45 Due to the heat transfer from cement to the surrounding mould, such tests might underestimate the exothermic temperature of the bone cement. The time–temperature profiles were observed to be sensitive to the thickness of the cement mass and the mould material. The study highlights the effects of mould material and the geometry on curing parameters. It was found that the measured peak temperature using polytetrafluoroethylene (PTFE) moulds varied by about 75°C for different mould heights but only by 28°C with polyurethane (PU) foam moulds. The measured setting time with

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PTFE moulds varied by about 740 s for different mould heights, while only by about 130 s for PU moulds. The authors conclude that poor conductivity materials, such as PU foam, were more suitable for characterisation of the thermal behaviour of bone cements. The maximum value admitted by both standards for maximum temperature (Tmax) measured in vitro is 90°C; however, it has been observed that temperature measurements recorded in vivo at the cement–bone interface are lower.46 A wide variety of temperatures at the interface between the bone and bone cement have been reported by Planell et al.47 In a recent paper, Quarini et al.48 developed a numerical model that predicts the temperature–time history in an idealised prosthetic–cement–bone system. The model is used to identify the critical parameters controlling thermal and unreacted monomer distributions. They reported that the degree of polymerisation was very dependent on the thermal history of the setting process. Another outstanding finding from this work was that relatively minor changes in the cement thickness produced very dramatic effects on thermal damage and burn necrosis. Therefore, from this study the surgeons are advised to be very careful with the thickness of the cement mass, and they should attempt to ensure that it is as close to 3 mm as possible, but not higher than 4.5 mm.

9.7

Rheology of acrylic bone cements

The viscosity of the bone cement in its early life is a crucial property since it represents the capacity of the system to flow into interstices within cancellous bone and enable effective mechanical interlock at the bone cement– bone interface. Typically, at early time periods the cement presents its lowest viscosity and, as the reaction proceeds, viscosity increases, thereby decreasing the cement’s flowability. Bone cements can be applied in two ways: the simplest method consists of mixing the cement and when it reaches its dough stage, with a viscosity high enough to be manually manipulated, the cement is inserted into the bone. The other way involves the use of a syringe or cement gun to inject the paste and has the advantages of reducing porosity and reaching higher penetration into the bone cavity. Obviously, for each method, bone cements with different flow characteristics are required. According to this property, commercially available acrylic bone cements can be found as low-, medium- and high-viscosity formulations. The studies carried out on the performance of the cements have found some correlation between the different durations of the stages of hardening and the type of cement.2 The differences in viscosity for high-viscosity and lowviscosity bone cements have been reported for various commercial brands.49 Examples of commercial brands of high-viscosity bone cements that are in the market are Genta C-ment® 1 (EMCM BV), Osteopal® HA (Merck),

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Osteopal® VS (Merck), Palacos® R (Merck), Palacos® R with Gentamicin (Merck), Palamed® (Merck) and Subiton (Prothoplast). Typical examples of low-viscosity bone cements in the market are Endurance® (DePuy/ Johnson & Johnson), Osteobond® (Zimmer), Osteopal® (Merck), Palacos® LV/E Flow (Schering Plough) and Zimmer® dough-type radiopaque and radiolucent (Zimmer). The clinical benefits of using one type of formulation rather than another seem to be influenced by the femoral stem applied; however, the topic is still under debate. A recent paper has reported an in vivo study on the migration of an Exeter-type stem using röntgen stereophotogrammetric analysis (RSA) and the authors conclude that cement viscosity does not influence the 1-year migration, and it is therefore unlikely to affect longterm outcome.50 The method described in the ASTM F451 and ISO 5833 standards to determine a viscosity-like parameter for cements that are applied as a dough, is an intrusion test; for cements applied through a syringe, the standards describe an extrusion viscosity test performed with a capillary rheometer. However, there are few studies that have analysed the rheological behaviour of acrylic bone cements by using the recommended techniques.41 Noble51 has reported studies on the viscosity changes produced during the polymerisation of several commercial acrylic bone cements; Noble developed an extrusion technique in which the cement was extruded from a syringe, and the variation in extrusion force with working time was recorded, as indicated by the ASTM F451 standard. A recent investigation has been conducted using a capillary rheometer to study the influence of initial component temperature on the apparent viscosity and handlings characteristics of PMMA bone cements.52 It was found that, by adjusting the initial temperature of the components, high-viscosity cement is able to mimic the flow characteristics of medium-viscosity cement. In other studies, the rheological properties of several bone cements have been evaluated by using a rotational cone and plate rheometer.53 A low-viscosity brand was found to be nearly one-half as viscous as the conventional formulations during the working time. Other rheological studies54 have been performed using an oscillating parallel plate rheometer, which is adequate to characterise the viscoelastic properties of the material. The evolution of dynamic viscosity and viscoelastic parameters with time provides the complete characterisation of the cement, which in its early life behaves as a viscous material and finally behaves as an elastic solid as the cement sets. By application of this technique, the rheological behaviour of suspensions of PMMA on a mixture of MMA and BMA monomers was investigated and the effects of concentration of powder, temperature and addition of inorganic components as radiopaque agents were evaluated.55 Viscosity increased with temperature and

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in the presence of a radiopaque agent. Different models in the literature were applied to results obtained for formulations prepared with different particle concentrations and only the model proposed by Chong et al.56 gave the best fit at low shear stress; however, two adjustable parameters were introduced to improve the correlation. In an investigation on the variables that affect the rheological parameters of bone cements, Lewis et al.57 found that they are strongly dependent on the characteristics of their powder particles, specifically the relative amounts of small-sized PMMA beads (mean diameter between 0 and 40 μm) and large-size PMMA beads (mean diameter higher than 75 μm).

9.8

Low molecular weight residuals after curing

Due to the fact that the setting reaction terminates in vivo, all unreacted products remain within the cement mass. Among them, residual monomer, activator and initiator will be the most significant agents in terms of producing any adverse biological response, especially those that have sufficient solubility in tissue fluids to leach from the materials into adjacent tissues and the systemic circulation. Residual monomer is always present in the cured cement due to vitrification phenomenon of the system. As a consequence of the high viscosity of the reacting mass, the polymerisation process evolves with difficulty and stops after a certain time due to the macromolecular rigidity of the vitreous state attained during the process. During the curing process the average temperature after reaching the maximum decreases to about 40°C which is very low compared with the glass transition temperature of PMMA, around 110°C, and this fact prevents the total conversion of monomer into polymer. Therefore, the non-reacted monomer remains trapped within the cement mass. At the time the cement is inserted into the body, a substantial amount of monomer is present in the polymerising mass.58 Immediately after insertion into human bone marrow, part of the monomer is released whereas the rest continues polymerising. Willert et al.59 analysed the content of monomer released into the surrounding tissues during the polymerisation process and reported values ranging from 0.4 to 1.2 wt% in the fat, 0.2–0.4 wt% in bone marrow fibres and cells, and 0.03–0.3 wt% in red blood cells. Assuming cement thickness in the range of 1.0–2.0 cm, estimations on the typical clinical ratios of release surface area to volume provided values from 0.5 to 4.0 cm2/cm3 calculated using several total hip components.60 Based on this assumption and using an experimental model, it has been concluded that mixing and insertion technique variations available to the surgeon might reduce the patient exposure to monomer from 3.5 wt% down to 1.2 wt%; however, it is still unclear whether such a reduction is sufficient to be clinically helpful in eliminating patient complications. Residual monomer

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entrapped in commercialised acrylic bone cements cured in vitro has been measured by different techniques, such as nuclear magnetic resonance,61 gas chromatography62 and high-performance liquid chromatography.63 The content of monomer has been reported to be in the range of 2–6 wt%.49 However, this amount decreases over time due to the slow postpolymerisation. Brauer et al.62 analysed the content of monomer in cured samples after storage in air and found contents around 1 wt% in a 3-month period and, later on, Kuhn et al.2 reported values lower than 1% in a 4-week period. On the other hand, a part of the monomer is released through diffusion processes into the physiological fluid and it is later metabolised to carbon dioxide and water in the citric acid cycle in the human body.64 The amount of monomer leaching out under clinical conditions will depend mainly on the volume and geometry of the bone–cement interface and on the diffusion rate of the monomer from the cement to the surrounding medium. Values of 0.43 wt% have been reported62 for the amount of monomer leached out from the cement after 6 days of immersion in water at 37°C. Other authors also support that the monomer released into aqueous medium is always less than 3% of the total monomer weight, and that mainly, the release takes places within the first 15 min after immersion.65 MMA is the only monomer present in the majority of commercial formulations. However, a few brands contain BMA in the liquid phase. BMA is less soluble in water than MMA, so it may remain in the cured system for a longer period of time. In addition, BMA is metabolised at a much slower rate than MMA because of its more hydrophobic structure.63 Residual monomer in acrylic bone cements has been investigated under different curing conditions66 and the maximum conversion attained by the monomer during polymerisation of a commercial PMMA bone cement was characterised by thermal analysis. In this paper a phenomenological kinetic model was also proposed which was able to reproduce experimental results obtained in dynamic test. Other studies reported by the same authors support the findings that the cure temperature and mould dimensions affect the amount of residual monomer in the hardened material which, in turn, influences mechanical properties67 such as flexural modulus, compressive yield stress and fracture toughness. Little information is available in the literature on the amount of residual activator and initiator in bone cement matrices. These compounds, although present in small concentrations in the set cement, can compromise the biocompatibility of the final material.68 Trap et al.69 reported concentrations of DMT of between 0.2 and 0.3 wt% in commercial PMMA cements postcuring. The concentrations of tertiary aromatic amine in eluates of bone cement pellets after 72 h of extraction were 38–48 ppm. Stea et al.70 reported concentrations of DMT at different times from mixing and after 72 h extraction and found that DMT concentration in aqueous extracts decreases with

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time. A high release was observed at 6 h after mixing and the release strongly decreased after 48 h, to almost negligible amounts 7 days after curing. Among the different commercial brands, they found a correlation between the initial content of amine in the liquid phase and the final concentration. Regarding residual initiator, even less information is available in the literature. However, the presence of residual peroxide in the cured cement was confirmed and a theoretical amount of BPO of 3.20 × 10−3 mol was reported for a dosage of cement.66 Residual BPO can leach from the cement and it converts within a few seconds to benzoic acid when in contact with blood, serum and saliva.68 On the other hand, residual BPO can react very slowly with the residual DMT.62 Some studies on the toxicity of commercial acrylic bone cements have found correlations between toxicity and initial concentration of BPO in the solid phase. The toxicity could be mediated by free radicals remaining in the set mass and it is reasonable to assume that they will be more abundant in cements with high BPO content.71

9.9

Future trends

9.9.1 Research on new activators of reduced toxicity Studies on the toxicity of DMT have revealed that this amine is a chromosome-damaging agent and has a significant clastogenic effect.72,73 DMT produces a delay in the cell replication cycle in vitro,70 it is a clear inhibitor of protein synthesis and interferes with the mineralisation processes.74 This activator has been associated with an adverse reaction at the bone–cement interface, mainly observed in hypersensitive patients.75 In order to reduce the adverse biological effects related to DMT, other more biocompatible amines based on 4-N,N-dialkylaminophenalkanoic acids and their corresponding methyl esters were introduced by Brauer et al.76 in the curing of commercial bone cements. The amines 4-N,N-(dimethylamino) phenethanol and 4-N,N-(dimethylamino) phenyl acetic acid yield negative Ames test and thus appear to be more biocompatible than DMT. In fact, the amine 4-N,N-(dimethylamino) phenethanol has been incorporated in a commercial low-viscosity bone cement.77 Trap et al.69 tried to substitute partially the activator DMT with the more biocompatible N,N-dihydroxypropyl-4toluidine in the curing of the new bone cement Boneloc. The results indicated that a more optimal activator/initiator system can be achieved, which produces a faster and more complete curing of the cement than that of DMT/PMMA. Other investigations have focused on the preparation of polymerisable tertiary amines susceptible to reacting with the growing macro-radicals and being incorporated into the polymeric matrix.78 In addition, due to the

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copolymerisation of the polymerisable amine with MMA, a slightly crosslinked polymer is obtained. Polymerisable amines based on unsaturated tertiary-aryl-amine accelerators such as acryloyl- and methacryloyl-Nphenylpiperazine have been synthesised and tested as activators of the amine/BPO system by Tanzy et al.79,80 These amines were used to cure CMW bone cement,81 the acryloyl derivatives giving rise to good results in terms of mechanical properties and suggesting that they could be considered as good alternatives in the preparation of biomedical acrylic resins. Activators of reduced toxicity82 such as N,N-dimethylaminobenzyl alcohol and the polymerisable amine 4-N,N-dimethylaminobenzyl methacrylate have been shown to be efficient in the curing of experimental acrylic bone cements83 and provided a lowering of exotherm temperature and a rise of the setting time.84 Less cytotoxic activators based on a long chain, naturally occurring fatty acid85 have been synthesised86,87 in order to obtain hydrophobic and high molecular weight activators, which are more biocompatible and less prone to leaching out from materials. N,N-dimethylamino-4-benzyl laurate and N,N-dimethylamino-4-benzyl oleate have been successfully used in the curing of PMMA bone cements as well as the standard commercial Palacos®R. These activators provided cements with lower exotherms, improved fracture toughness and fatigue life88 and fulfil an array of parameters that make them suitable alternatives to DMT.89 Recently, activators with antiseptic properties have been applied. 4,4′-Dimethylamino benzydrol with a chemical structure resembling that of crystal violet has been shown to be four times less cytotoxic than DMT and more active against different microorganisms, showing similar efficiency in the amine/ BPO redox initiation system.90 The in vivo response to the intraosseous implantation of the dough cement revealed an early osseous neoformation and less connective tissue at the bone–cement interface.91

9.9.2 Research on alternative formulations versus the traditional powder/liquid cements An alternative to the traditional formulation based on the two solid and liquid components has been reported by Moseley et al.92 The system consists of a toughened prepolymer and a gel polymerisation method. The toughened PMMA prepolymer (pellets) has a particle size distribution and particle sizes ranging from 74 to 500 μm. The gel phase consists of a solution of the prepolymer in MMA with 0.75 wt% DMT, at a ratio of 1.3 to 1. Polymerisation of the gel is carried out by mixing with enough 12 wt% BPO solution to give a final concentration of 0.2 wt%. These cements presented a higher residual monomer content after 2 h of polymerisation (around 7 wt%) compared with the traditional cement (around 3 wt%), which can

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be attributed to the absence of a dispersed solid phase. Residual monomer content decreased to around 3 wt% after 4 months of exposure to an in vitro medium, due to diffusion processes. Mechanically, the cements so prepared possessed significantly reduced values of compressive strength compared with the traditional formulations, even lower than the value of 70 MPa established in the standards. Elastic modulus also decreased, but the modulus of rupture of the toughened materials was slightly higher. However, the most striking properties that improved in the new materials are the manner in which failure occurs, which is not brittle, and the fracture toughness, showing a large increase in the KIC (the value of KI, the mode I stress intensity factor range generated during a single loading cycle, at fracture, which corresponds to the maximum load applied) attributed to the presence of the toughening component, and revealing that a material with improved resistance to fracture can be obtained by this method. Another alternative to traditional powder/liquid formulations has been investigated, but its true composition remains the same as current formulations. It is based on a high-viscosity PMMA solution system.93 Two separate solutions are prepared by dissolving PMMA powder into MMA monomer and adding the initiator to one solution and the activator to the other. Then, the system is based on solutions rather than admixtures. This fact produces notable changes in the rheological behaviour of the cement, especially in its early life. The solutions have a much higher initial viscosity, but it remains relatively constant throughout the early stages of polymerisation, which allows for more consistent handling characteristics throughout the working stage of the cement. Another advantage of this new system is that it eliminates the porosity associated with the spatulation of the mass. In addition, the dependence of material properties on the surgical technique is decreased because the two solutions can be mixed and delivered via a single, closed system, without the attendant issues related to mixing powders and liquids. Behind this rationale, there are different variables, which have to be optimised in order to obtain desirable properties in the final material. These are the polymer to monomer ratio, polymer molecular weight, and initiator and activator concentrations. Of these, the concentration and molecular weight of the polymer will have a strong influence on the polymerisation exotherm and on the mechanical properties of the set cement, whereas BPO and DMT concentrations will influence setting time. A range of setting times between 5 and 16 min are obtained by decreasing the DMT concentration from 1.4 ml to 0.2 ml for a constant BPO concentration of 1.25 g/100 ml of MMA. Mechanically, the set cement possesses higher flexural strength and modulus compared with Simplex® P cement, and this was attributed to the lack of porosity but also to the higher molecular weights of the set cement that can be achieved using the solution cement.98 Fracture toughness and fatigue strength for the solution formulations are

200

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comparable with Simplex® P and are not affected by the BPO concentration or the BPO : DMT ratio. Residual monomer content is the weak point of this system, since it is significantly higher compared with commercial cement; this is attributed to the higher initial monomer concentration rather than to a lower degree of conversion.94 In conclusion, the latest efforts dedicated to improve the characteristics of acrylic bone cement formulations have been directed at improving their biocompatibility and their processability. The first improvement has been pursued by changing the activator of the redox system for another of reduced toxicity. Advances in this topic have been successful and recently the less toxic amine 2-(4-dimethylamino)phenyl ethanol has been incorporated into a commercial formulation. The second improvement has attempted to change the traditional solid and liquid phases by using two solutions in order to improve handling characteristics, reduce porosity and avoid mixing instruments. Mechanically, these formulations are good, but they have the drawback of increased residual monomer. More studies must be carried out on this topic.

9.10

References

1 charnley j. Anchorage of the femoral head prosthesis to shaft of the femur. J Bone Joint Surg 1960;42B:28–35. 2 kuhn kd. Bone Cements. Up-To-Date Comparison of Physical and Chemical Properties of Commercial Materials. Springer, Germany, 2000. 3 pascual b, vázquez b, gurruchaga m, goñi i, ginebra mp, gil fj, planell ja, levenfeld b, san román j. New aspects of the effect of size and size distribution on the setting parameters and mechanical properties of acrylic bone cements. Biomaterials 1996;17:509–516. 4 lautenschlager ep, stupp si, keller jc. Structure and properties of acrylic bone cements. In Functional Behaviour of Orthopaedic Biomaterials, Eds P Ducheyne, G Hastings. CRC Press, Boca Raton, Florida, 1987. 5 rudigier j, kirschner p, richter ie, schweikert ch. Influence of different X-ray contrast materials on structure and strength of bone cements. In Mechanical Properties of Biomaterials, Eds GW Hastings, DF Williams. John Wiley and Sons, Chichester, UK, 1980, pp. 289–297. 6 silikas n, al-kheeeraif a, watts dc. Influence of P/L ratio and peroxide/amine concentrations on shrinkage-strain kinetics during setting of PMMA/MMA biomaterial formulations. Biomaterials 2005;26:197–204. 7 nussbaum da, gailloud p, murphy k. The chemistry of acrylic bone cements and implications for clinical use in image-guided therapy. J Vasc Interv Radiol 2004;15:121–126. 8 bowen rl, argentar h. A method for determining the optimum peroxide-toamine ratio for self-curing resins. J Appl Polym Sci 1973;17:2213–2222. 9 madigan s, towler, lewis g. Optimisation of the composition of an acrylic bone cement: application to relative amounts of the initiator and the activator/ co-initiator in Surgical Simplex® P. J Mater Sci Mater Med 2006;17:307–311.

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10 lewis g, mladsi s. Effect of sterilisation method on properties of Palacos R acrylic bone cement. Biomaterials 1998;19:117–124. 11 harper ej, braden m, bonfield w, dingeldein e, wahlig h. Influence of sterilization upon a range of properties of experimental bone cements. J Mater Sci Mater Med 1997;8:849–853. 12 graham j, pruitt l, ries m, gundiah n. Fracture and fatigue properties of acrylic bone cement. The effects of mixing method, sterilisation treatment and molecular weight. J Arthroplasty 2000;15:1028–1035. 13 wang js, franzen h, jonsson e, lidgren l. Porosity of bone cement reduced by mixing and collecting under vacuum. Acta Orthop Scand 1993;64:143–146. 14 haas ss, brauer gm, dickson g. A characterisation of polymethylmethacrylate bone cement. J Bone Joint Surg 1975;57A:380–391. 15 davies jp, harris wh. Comparison of diametral shrinkage of centrifuged Simplex P bone cement. J Appl Biomater 1995;6:209–211. 16 gilbert jl, hasenwinkel jm, wilson rl, lautenschlager ep. A theoretical and experimental analysis of polymerisation shrinkage of bone cement: a potential major source of porosity. J Biomed Mater Res 2000;52:210–218. 17 orr jf, dunne nj, quinn jc. Shrinkage stresses in bone cement. Biomaterials 2003;24:2933–2940. 18 elias hg. Macromolecules. Structures and Synthesis, Vol. 1. Wiley VCH Verlag GmbH & Co. KGaA, Weinheim, Germany, 2005. 19 feng xd. The role of amine in vinyl radical polymerisation. Macromol Chem Macromol Symp 1992;63:1–18. 20 vázquez b, deb s, bonfield w, san román j. Characterization of new acrylic bone cements prepared with oleic acid derivatives. J Biomed Mater Res, Part B, Appl Biomater, 2002;63:88–97. 21 sack r, schlulz gv, meyerhoff g. Free radical polymerisation of methyl methacrylate up to the glassy state. Rates of propagation and termination. Macromolecules 1988;21:3345–3352. 22 migliaresi c, fambri l, kolarik j. Polymerisation kinetics, glass transition temperature and creep of acrylic bone cements. Biomaterials 1994;15:875–881. 23 yang jm, shyu js, chen hl. Polymerysation of acrylic bone cement investigated by differential scanning calorimetry: Effects of heating rate and TCP content. Polym Engng Sci 1997;37:1182–1187. 24 lewis g, xu j, madigan s, towler mr. Influence of two changes in the composition of an acrylic bone cement on its handling, thermal, physical and mechanical properties. J Mater Sci Mater Med 2007;18:1649–1658. 25 maffezzoli a. Polymerisation kinetics of acrylic bone cements by differential scanning calorimetry. J Therm Anal 1996;47:35–49. 26 borzachiello a, ambrosio l, nicolais l, harper ej, tanner ke, bonfield w. Comparison between the polymerisation behaviour of a new bone cement and a commercial one: modelling and in vitro analysis. J Mater Sci Mater Med 1998;9:835–838. 27 achilias ds, sideridou i. Study of the effect of two BPO/amine initiation systems on the free radical polymerisation of MMA used in dental resins and bone cements. J Macromol Sci Pure & Appl Chem 2002;A39:1435–1450. 28 oldfield ff, yasuda hk. ESR study of MMA polymerisation by a peroxide/ amine system: Bone cement formation. J Biomed Mater Res 1999;44:436–445.

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29 park jb, ackley ma, turner rc. EPR studies of PMMA bone cement. Adv Biomater 1982;3:303–307. 30 park jb, turner rc, atkins pe. EPR study of free radicals in PMMA bone cement: a feasibility study. Biomater Med Devices Artif Organs 1980;8:23–33. 31 turner rc. Free radical decay kinetics in PMMA bone cement. J Biomed Mater Res 1984;18:467–473. 32 turner rc, white fb, park jb. The effect of initial temperature on free radical decay in PMMA bone cement. J Biomed Mater Res 1982;16:639–646. 33 turner rc, atkins pe, ackley ma, park jb. Molecular and macroscopic properties of PMMA bone cement: free-radical generation and temperature change versus mixing ratio. J Biomed Mater Res 1981;15:425–432. 34 halpern hj, bowman mk. Low frequency EPR spectrometers: MHz range. In EPR Imaging and In Vivo EPR, Eds GR Eaton, S Eaton, K Ohno. CRC Press, Boca Raton, Florida, 1991, pp. 5–63. 35 gallez b, beghein n. Non-invasive in vivo ERP monitoring of the methyl methacrylate polymerisation during the bone cement formation. Biomaterials 2002;23:4701–4704. 36 looney ma, park jb. Molecular and mechanical property changes during aging of bone cement in vitro and in vivo. J Biomed Mater Res 1986;20:555–563. 37 american society for testing materials (astm). ASTM Standard F451-99a: Standard specification for acrylic bone cement. In Annual Book of ASTM Standards, Vol 13.01. ASTM, West Conshohocken, PA, United States. 38 international organisation for standardisation (iso). ISO 5833:2002(E): Implants for surgery – Acrylic resin cements. ISO, Geneva, Switzerland. 39 belkoff sm, sanders jc, jasper le. The effect of the monomer-to-powder ratio on the material properties of acrylic bone cement. J Biomed Mater Res, Part B, Appl Biomater 2002;63:396–399. 40 he s, scott c, de luise m, edwards b, higham p. Effect of choice of surgical gloves on dough time measurements of acrylic bone cement. Biomaterials 2003;24:235–237. 41 lewis g. Properties of acrylic bone cement: state of the art review. J Biomed Mater Res, Part B, Appl Biomater 1997;38:155–182. 42 deb s. A review of improvements in acrylic bone cements. J Biomater Appl 1999;14:16–47. 43 dunne nj, orr jf. Curing characteristics of acrylic bone cement. J Mater Sci Mater Med 2002;13:17–22. 44 milner r. The development of theoretical relationships between some handling parameters (setting time and setting temperature), composition (relative amounts of initiator and activator) and ambient temperature for acrylic bone cement. J Biomed Mater Res, Part B, Appl Biomater 2004;68B:180–185. 45 li c, mason j, yakimicki. Thermal characterisation of PMMA-based bone cement curing. J Mater Sci Mater Med 2004;15:85–89. 46 huiskes r. Some fundamental aspects of human joint replacement. Analysis of stresses and heat conduction in bone–prosthesis structures. Acta Orthop Scand 1980;51(suppl 185): 43–105. 47 planell ja, vila mm, gil fj, driessens fcm. Acrylic bone cements. In Encyclopaedic Handbook of Biomaterials and Bioengineering. Part B: Applications, Volume 2, Eds DL Wise, DJ Trantolo, DE Altobelli, MJ Yaszemski, JD Gresser, ER Schwartz. Marcel Dekker, Inc., New York, 1995, pp. 879–921.

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48 quarini gl, learmonth id, gheduzzi s. Numerical predictions of the thermal behaviour and resultant effects of grouting cements while setting prosthetic components in bone. Proc. IMechE, Part H: J Engng Med 2006;220:625–634. 49 kuehn kd, ege w, gopp u. Acrylic bone cements: composition and properties. Orthop Clin Am 2005;36:17–28. 50 glyn-jones s, hicks j, alfaro-adrian j, gill hs, mclardy-smith p, murray dw. The influence of cement viscosity on the early migration of a tapered polished femoral stem. Int Orthop 2003;27:362–365. 51 noble pc. Selection of bone cements for use in joint replacement. Biomaterials 1983;4:94–100. 52 sullivan sjl, topoleski ldt. Influence of initial component temperature on the apparent viscosity and handling characteristics of acrylic (PMMA) bone cement. J Biomed Mater Res, Part B, Appl Biomater 2007;81B:224–230. 53 ferracane jl, greener eh. Rheology of acrylic bone cements. Biomater Med Devices Artif Organs 1981;9:213–224. 54 farrar df, rose j. Rheological properties of PMMA bone cments during curing. Biomaterials 2001;22:3005–3013. 55 nzihou a, attias l, sharrock p, ricard a. A rheological, thermal and mechanical study of bone cement – from a suspension to a solid biomaterial. Powder Technol 1998;99:60–69. 56 chong js, christiansen eb, baer ad. Flow of viscous fluid through a circular aperture. J Appl Polym Sci 1971;15:369–379. 57 lewis g, carroll m. Rheological properties of acrylic bone cement during curing and the role of the size of the powder particles. J Biomed Mater Res, Part B, Appl Biomater 2002;63:191–199. 58 sheinin eb, benson rw, brannon w. Determination of methyl methacrylate in surgical acrylic cement. J Pharm Sci 1976;65:280–283. 59 willert hg, frech ha, bechtel a. Measurements of the quantity of monomer leaching out of acrylic bone cement into the surrounding tissues during the process of polymerisation. In Biomedical Applications of Polymers, Polymer Science and Technology, Vol 7, Ed. HP Gregor. Plenum Press, New York, 1975. 60 bayne sc, lautenschalager ep, greener eh, mayer pr. Clinical influences on bone cement monomer release. J Biomed Mater Res 1977;11:859–869. 61 vázquez b, deb s, bonfield w. Optimization of benzoyl peroxide concentration in an experimental bone cement based on poly(methyl methacrylate). J Mater Sci Mater Med 1997;8:455–460. 62 brauer gm, termini dj, dickson g. Analysis of the ingredients and determination of residual components of acrylic bone cements. J Biomed Mater Res 1977; 11:577–607. 63 davy kwm, braden m. Residual monomer in acrylic polymers. Biomaterials 1991;12:540–544. 64 wenda k, scheuermann h, weitzel e, rudigier j. Pharmacokinetics of methylmethacrylate monomer during total hip replacement in man. Arch Orthop Trauma Surg 1988;107:316–321. 65 schoenfeld cm, conard gj, lautenschlager ep. Monomer release from methacrylate bone cements during simulated in vivo polymerisation. J Biomed Mater Res 1979;13:135–147. 66 vallo ci. Residual monomer content in bone cements based on poly(methyl methacrylate). Polym Int 2000;49:831–838.

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67 vallo ci, montemartini, cuadrado tr. Effect of residual monomer content on some properties of a poly(methyl methacrylate)-based bone cement. J Appl Polym Sci 1998;68:1367–1383. 68 shintani h, tsuchiya t, hata y, nakamura a. Solid phase extraction and HPLC analysis of toxic components eluted from methyl methacrylate dental materials. J Anal Toxicol 1993;17:73–78. 69 trap b, wolff p, jense js. Acrylic bone cements: Residuals and extractability of methacrylate monomers and aromatic amines. J Appl Biomater 1992;3:51–57. 70 stea s, granchi d, zolezzi c, ciapetti g, visentin m, cavedagna d, pizzoferrato a. High-performance liquid chromatography assay of N,N-dimethyl-p-toluidine released from bone cements: evidence for toxicity. Biomaterials 1997;18:243– 246. 71 ciapetti g, granchi d, stea s, cervellati m, pizzoferrato a, toni a. In vitro testing of ten bone cements after different time intervals from polymerisation. J Biomater Sci Polymer Edn 2000;11:481–493. 72 taningher m, pasqini r, bonatti s. Genotoxicity analysis of N,N-dimethyl-ptoluidine. Environ Mol Mutagen 1993;21:349–356. 73 bigatti mp, lamberti l, rizzi fp, cannas m, allasia g. In vitro micronucleus induction by polymethyl methacrylate bone cement in cultured human lymphocytes. Mutat Res 1994;321:133–137. 74 bosch p, harms h, lintner f. Toxicity of bone cement component. Aktuelle Probl Chir Orthop 1987;31:87–89. 75 haddad fs, levell nj, dowd pm, cobb ag, bentley g. Cement hypersensitivity: a cause of aseptic loosening? J Bone Joint Surg 1995;77B:329–330. 76 brauer gm, steinberger dr, stansbury jw. Dependence of curing time, peak temperature and mechanical properties on the composition of bone cement. J Biomed Mater Res 1986;20:839–852. 77 fritsch ew. Static and fatigue properties of two new low-viscosity PMMA bone cements improved by vacuum mixing. J Biomed Mater Res 1996;31:451–456. 78 dnebosky j, hynkova v, hrabak f. Polymerizable amines as promoters of coldcuring resins and composites. J Dent Res 1975;54:772–776. 79 tanzi mc, levi m, danusso f. Amides from N-phenylpiperazine as low-toxicity activators in radical polymerisations. Polymer 1990;31:1735–1738. 80 sandner b, baudach s, knoth p, tanzi mc. Copolymerisation and activation of peroxide decomposition with acrylic derivatives of tertiary aromatic amines. Polymer 1994;15:3285–3289. 81 tanzi mc, sket i, gatti am, monari e. Physical characterization of acrylic bone cement cured with new accelerator systems. Clin Mater 1991;8:131–136. 82 liso pa, vázquez b, rebuelta m, hernáez ml, rotger r, san román j. Analysis of the leaching and toxicity of new amine activators for the curing of acrylic bone cements and composites. Biomaterials 1997;18:15–20. 83 vázquez b, elvira c, levenfeld b, pascual b, goñi i, gurruchaga m, ginebra mp, gil fx, planell ja, liso pa, rebuelta m, san román j. Application of tertiary amines with reduced toxicity to the curing process of acrylic bone cements. J Biomed Mater Res 1997;34:129–136. 84 vázquez b, levenfeld b, san román j. Role of amine activators on the curing parameters, properties and toxicity of acrylic bone cements. Polym Int 1998;46: 241–250.

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10 Calcium phosphate bone cements M. P. G I N E B R A, Technical University of Catalonia (UPC), Spain

Abstract: Calcium phosphate cements (CPCs) are self-setting bioactive materials with unique properties for bone regeneration applications. The chapter first describes CPCs and compares them to other orthopaedic cements widely used in clinics, such as acrylic cements. Subsequently, a brief overview of the chemistry of calcium phosphates is presented, together with a classification of the main types of CPCs. The basic physico-chemical, mechanical and biological properties of CPCs are also reviewed and the technological issues most relevant to improving the clinical performance of CPC are described. Finally, the present and potential applications of these materials are discussed. Key words: calcium phosphate cements, bone cements, bone regeneration, hydroxyapatite, bioceramics.

10.1

Introduction

This chapter focuses on calcium phosphate cements (CPCs), a family of materials that has great potential in bone regeneration, due to their unique properties of bioactivity, injectability and in vivo setting ability. In Section 10.2, a historical perspective is presented, and CPCs are defined and compared with other orthopaedic cements widely used in clinics, such as acrylic cements. Subsequently, a brief overview of the chemistry of calcium phosphates is presented, necessary in order to understand the nature of the cementation reaction, together with a classification of the main types of CPCs. Section 10.4 presents the basic properties of the CPCs, and these are compared with conventional high-temperature calcium phosphate ceramics, stressing the advantages of this family of materials. The technological issues relevant to improving the clinical performance of CPCs are the subject of Section 10.5. Another aspect that is dealt with is the present and potential applications of these materials, both as injectable or pre-set materials – such as bone cavity filling, vertebroplasty, scaffolds for tissue engineering or drug delivery. Finally, the new trends and research lines that these materials instigate are presented. 206

Calcium phosphate bone cements

10.2

207

Historical overview: calcium phosphate cements versus acrylic cements

When a surgeon uses the term ‘bone cement’ he often refers to acrylic bone cements, widely used in orthopaedic surgery since the 1960s, especially for arthroplasty fixation.1,2 However, at the beginning of the 1980s a new family of bone cement emerged, namely the CPCs. In common with the acrylic bone cements, this new family could self-set inside the body, which allowed its implantation in a paste form. However, their chemical nature, properties and applications are very different, as summarised in Table 10.1, and described in the following sections. CPCs are hydraulic cements, which means that water is used as the liquid phase of the cement, and their hardening is not due to a polymerisation reaction, but to a dissolution and precipitation process. Unlike acrylic bone Table 10.1 Comparison of the nature and properties of acrylic bone cements versus calcium phosphate bone cements

Acrylic bone cements

Calcium phosphate bone cements

Material type

Polymer

Ceramic

Liquid phase

Mainly methyl methacrylate

Water or aqueous solutions

Powder component

Polymer beads (PMMA/copolymers) Some inorganic non-reacting phase can be added as radiopaque agent

Calcium phosphate powders

Setting reaction mechanism

Polymerisation

Dissolution and precipitation reaction

Reaction products

Mainly polymethyl methacrylate

Calcium phosphates, usually hydroxyapatite or brushite

Exothermic peak temperature during setting (ISO 5833)

50–90°C

37°C

Stability

Non-resorbable

Resorbable (low or high resorption rate depending on composition and microstructure)

Bioactivity

Non-bioactive

Bioactive

Applications

Moderate load-bearing applications: arthroplasty fixation, vertebroplasty

Bone regeneration Non load-bearing applications

208

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cements, this reaction is not exothermic and therefore it does not evoke problems of necrosis by hyperthermia.3,4 Some of the salient features of CPCs are their excellent biocompatibility, bioactivity with the ability to form a direct bonding with bone, and osteoconductivity. In addition, they can be resorbable, with a resorption rate that depends on their composition and microstructural features. These features confer high bone regeneration potential. On the other hand, they have some limitations related to their poor mechanical properties and solubility. As with most ceramic materials, they are brittle and, in addition, due to the fact that they are intrinsically porous materials, their strength in general is lower than that of acrylic cements. This fact limits their use either to non-load-bearing applications, such as the treatment of maxillofacial defects or in load-bearing applications in combination with metal implants, for example, in the treatment of some fracture defects such as wrist fractures. CPCs were discovered by Legeros5, and Brown and Chow6 in the early 1980s. They demonstrated the formation of hydroxyapatite in a monolithic form at room or body temperature by means of a cementitious reaction. This was an important breakthrough in the field of bioceramics research, since it supplied a material that was mouldable, and therefore could adapt to the bone cavity, presenting a good fixation and an optimum tissue– biomaterial contact, necessary for stimulating the bone ingrowth. Since then, CPCs have attracted much attention and different formulations have been put forward.6–10 Currently, many commercial products exist on the market.11

10.3

Chemistry of calcium phosphate cements

In general, all CPCs are formed by a combination of one or more calcium orthophosphates, which upon mixing with a liquid phase, usually water or an aqueous solution, form a paste that is able to set and harden after being implanted within the body. The cement sets as a result of a dissolution and precipitation process, as represented in Fig. 10.1. The entanglement of the precipitated crystals is responsible for cement hardening. Calcium orthophosphates are the calcium salts derived from orthophosphoric acid. Their names, abbreviations, chemical formulae and Ca/P molar ratio are summarised in Table 10.2.12 Some of these calcium orthophosphates can be obtained by precipitation from an aqueous solution at low temperature, while others can only be obtained at high temperature. All of them can be used as reactants for CPCs, and only those calcium orthophosphates that can precipitate at low temperature in aqueous systems can be obtained theoretically as a result of the CPC setting reaction. However, despite the large number of possible formulations, the CPCs developed up to now have only two different end products, precipitated hydroxyapatite

Calcium phosphate bone cements

209

Cement Powder

Liquid

Dissolution +

Plastic paste Setting Rigid paste Hardening

Precipitation at physiological temperature

Solid body

10.1 Rationale of a calcium phosphate cement.

(PHA) or brushite (DCPD). This in fact is a predictable situation since hydroxyapatite is the most stable calcium phosphate at pH > 4.2 and brushite the most stable one at pH < 4.2. The thermodynamics of calcium phosphate salts in an aqueous solution at room or body temperature is the basis for understanding the manufacturing technology involved in CPCs for clinical applications. The solubility of the different compounds in the ternary system Ca(OH)2-H3PO4-H2O at 37°C is normally represented through solubility diagrams, where the isotherms of different calcium phosphate salts in equilibrium with their saturated solution are plotted.13,14 In these diagrams either the calcium or the phosphate concentration in the saturated solution is represented in a logarithmic scale versus pH. It has to be taken into account that these diagrams are applicable to dilute or weakly supersaturated solutions, and CPCs consist of heavily supersaturated systems, far from equilibrium conditions. However, the application of these general concepts allows an understanding of the driving forces controlling dissolution and precipitation reactions, which are related to respective super- or under-saturation levels defined with regard to the thermodynamic solubility product. In fact, in addition to thermodynamic factors,13,14 kinetic factors can control both phase dissolution and the precipitation of PHA, and can determine the final products obtained in a CPC setting reaction.15,16 This means that the thermodynamic conclusions that can be derived from the solubility and relative stability diagrams of the different calcium phosphates must be taken as a first approximation, but never as an exhaustive explanation of what is actually happening during the setting reaction.

10.3.1 Apatite calcium phosphate cements The relevance of hydroxyapatite as a bone substitute arises from the fact that the mineral phase of bone is precisely a hydroxyapatite. In this sense, it has to be clarified that, even if stoichiometric hydroxyapatite has a fixed

Table 10.2 Calcium orthophosphate compounds Compound

Abbreviation

Formula

Ca/P ratio

Compounds that can precipitate at room temperature in aqueous systems Monocalcium phosphate monohydrate MCPM Ca(H2PO4)2·H2O Dicalcium phosphate dihydrate DCPD CaHPO4·2H2O Octocalcium phosphate OCP Ca8H2(PO4)6·5H2O Precipitated hydroxyapatite PHA Ca10(PO4)6(OH)2 Calcium-deficient hydroxyapatite CDHA Ca10-x(HPO4)x(PO4)6-x(OH)2-x Amorphous calcium phosphate ACP —

0.50 1.00 1.33 1.67 1.50–1.67 1.35–1.5

Compounds obtained at high temperature Monocalcium phosphate anhydrous Dicalcium phosphate α-Tricalcium phosphate β-Tricalcium phosphate Sintered hydroxyapatite Oxyapatite Tetracalcium phosphate

0.50 1.00 1.50 1.50 1.67 1.67 2.00

MCPA DCP α-TCP β-TCP SHA OHA TTCP

Ca(H2PO4)2 CaHPO4 α-Ca3(PO4)2 β-Ca3(PO4)2 Ca10(PO4)6(OH)2 Ca10(PO4)6O Ca4(PO4)2O

Mineral name

Brushite

Monetite

Hydroxyapatite

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composition, it is well known that the apatitic structure can exist in a range of compositions. Non-stoichiometric or calcium-deficient hydroxyapatite (CDHA) can be obtained at low temperatures, with a composition that can be expressed as Ca10−x(HPO4)(PO4)6−x(OH)2−x, where x ranges from 0 to 1, being 0 for stoichiometric hydroxyapatite and 1 for fully calcium-deficient hydroxyapatite. In fact, biological apatite is a carbonate containing calciumdeficient hydroxyapatite, which in addition contains several other ionic substitutions such as Na+, K+, Mg2+, F− and Cl−. Apatitic CPCs can form either PHA or CDHA through a precipitation reaction. Their fabrication process allows the incorporation of different ions in its lattice depending on the composition of the starting materials. In general it can be stated that the formation of hydroxyapatite through a cement-type reaction is a biomimetic process, in the sense that it takes place at body temperature and in a physiological environment, as happens when bone is formed or remodelled. This can account for the fact that the hydroxyapatite formed in the setting of CPCs is much more similar to biological apatites than that obtained when high-temperature sintering processes are applied to fabricate ceramic hydroxyapatite. The CPCs leading to the formation of PHA or CDHA can be classified in three groups, taking into account the number and type of calcium phosphates used in the powder mixture.17 1

Monocomponent CPCs, in which a single calcium phosphate compound hydrolyses to form PHA or CDHA. Since hydroxyapatite is the least soluble phase at pH > 4.2, this means that any other calcium phosphate present in an aqueous solution at that pH range will tend to dissolve, and PHA will tend to precipitate. As a result, H3PO4 or CaOH2 can be released into the solution as a by-product of the hydrolysis reaction. However, in most cases the formation of PHA from the hydrolysis of one calcium phosphate is kinetically very slow, due to a decrease of the super-saturation level, as the reaction proceeds.13 The only cement system that contains a single calcium phosphate was first reported by Monma et al.18,19 and was further optimised and characterised by Ginebra el al.20–25 This system is based on the hydrolysis of α-tricalcium phosphate (TCP) to CDHA according to equation [10.1]: 3α-Ca3(PO4)2 + H2O → Ca9(HPO4)(PO4)5(OH)

[10.1]

Since the Ca/P ratio of the initial and final calcium phosphates is the same, no acid or base is released as a by-product. 2

CPCs formed by two calcium phosphates, one acidic and the other one basic, which set following an acid–base reaction. The basic component is normally tetracalcium phosphate (TTCP), since it is the only calcium phosphate having a Ca/P ratio higher than PHA. Therefore, TTCP can

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be combined with one or more calcium phosphates with lower Ca/P ratio to obtain either PHA or CDHA, without the formation of acids or bases as by-products. From a theoretical point of view, any calcium phosphate more acidic than PHA can react directly with TTCP to form PHA or CDHA. The most widely studied combinations are in fact the TTCP + DCPD and TTCP + dicalcium phosphate (DCP) mixtures, which were first developed by Brown and Chow6,26 and have been the object of extensive research.15,16,27–30 These mixtures produce cements that set at body temperature in a pH range around neutral, according to equations [10.2] and [10.3].

3

Ca4(PO4)2O + CaHPO4 → Ca5(PO4)3OH

[10.2]

Ca4(PO4)2O + CaHPO4·2H2O → Ca5(PO4)3OH + 2H2O

[10.3]

Systems formed by more than two compounds, including calcium phosphates, and other salts, for example calcium or strontium carbonate, magnesium phosphates among others. An example of this group of CPCs is the product developed by Norian Corporation (Norian SRSTM, Skeletal Repair System),31 where mixtures of calcium phosphates with a Ca/P ratio lower than PHA are used and CaCO3 is added as an additional source of calcium ions. Specifically this system is formed by using a mixture of α-TCP, monocalcium phosphate monohyarate (MCPM) and CaCO3. The initial setting process involves the formation of DCPD, while the final setting product is dahllite, a carbonated hydroxyapatite with a Ca/P ratio between 1.67 and 1.69, and with a carbonate ion content similar to bone mineral.31

10.3.2 Brushite calcium phosphate cements Brushite (DCPD) is an acidic calcium phosphate that has been detected in some physiological sites, for instance in bone,32 fracture callus33 and kidney stones.34 In contrast to hydroxyapatite, brushite is metastable under physiological conditions,15 and for this reason brushite CPCs resorb much faster than apatite CPCs, although it has been shown that in vivo DCPD tends to convert into PHA.35 Some CPCs have been designed that give brushite (DCPD) as the end-product. All brushite CPCs are obtained as a result of an acid–base reaction. Several compositions have been proposed for brushite cements, e.g. β-TCP + MCPM,36 β-TCP + H3PO4,37,38 and TTCP + MCPM + CaO.35 In the first of these (β-TCP + MCPM), the reaction responsible for the setting of the cement is β-Ca3(PO4)2 + Ca(H2PO4)2·H2O + 7H2O → 4CaHPO4·2H2O [10.4] The paste of brushite CPC is acidic during setting because brushite can only precipitate at a pH values lower than 6.12 After setting, the pH of the cement

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paste slowly changes towards the equilibrium pH, which depends on the singular points of the phases present in the cement.39

10.4

Basic properties of calcium phosphate cements

In the late 1960s and early 1970s, the quest for improved biocompatibility of implant materials resulted in a new concept of bioceramic materials that would mimic natural bone tissue.40,41 Because hydroxyapatite, a naturally occurring ceramic mineral, is also the mineral component of bone, it was believed that synthetic hydroxyapatite used for bone replacement would be entirely compatible with the body. It was subsequently shown that these materials were bioactive materials, i.e. they were able to develop a direct bonding with bone, without the formation of a fibrous capsule. Therefore, they were osteoconductive materials, able to guide bone ingrowth on their own surface. Since then, ceramic calcium phosphates have been used, mainly in the form of sintered hydroxyapatite (SHA) or β-TCP ceramics, as materials for cavity filling for bone regeneration, or even as coatings for metallic prostheses. However, their fabrication method, usually sintering at high temperature, represented a significant drawback since it limited the formability of their shape and size. This often caused problems of adaptation and fixation to the bone cavity where they had to be placed. The development of CPCs at the beginning of the 1980s brought to light several advantages in comparison with the use of ceramic calcium phosphates. These can be summarised as follows: (a) in vivo self-setting ability; (b) injectability, which allows cement implantation by means of minimally invasive surgical techniques, less aggressive than the traditional surgical techniques; (c) perfect fit at the implant site, which assures good bone–material contact, even in geometrically complex defects; this allows for an optimum tissue–biomaterial contact, necessary for stimulating bone ingrowth; (d) reaction products chemically and structurally more similar to the biological hydroxyapatite (in the case of apatite cements) due to the fact that the setting reaction is a low-temperature dissolution–precipitation process; this contributes to an increased reactivity of CPCs as compared with calcium phosphate ceramics; (e) the possibility of incorporating different drugs – given the fact that the setting reaction takes place at low temperature, from antibiotics and anti-inflammatory drugs to growth factors that are able to stimulate certain biological responses.

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10.4.1 Processing parameters The properties of a cement system depend on, and therefore can be tailored by, several processing parameters, which are summarised in Table 10.3. The composition of the solid phase varies with the setting reactions desired. Normally, the solid phase is chosen from the orthophosphate family, listed in Table 10.2. On the other hand, those calcium orthophosphates containing biocompatible components such as Na+, K+, Mg2+, Zn2+, CO32−, SO42− or Cl− are also suitable as constituents of CPC powders, and calcium carbonate is also added in some formulations. The particle size of the starting powder plays an important role in the setting and final properties of the cement. As mentioned before, the setting reaction occurs through a dissolution–precipitation process. Hence, fineness of the powder will increase the rate of hardening since smaller particles will dissolve faster than bigger particles and the precipitation of a new phase will begin earlier.25 Another strategy that can be used to produce cement with good setting characteristics is the use of seed crystals to act as a ‘nucleator’ for the precipitation reaction. Several parameters can be modified, such as the amount of seed added, its crystallinity and crystal size. Although the effects of seeds under different conditions have not been entirely clarified, it seems that their main effect is to reduce the setting time of the cement. The primary role of the liquid phase is to function as a vehicle for dissolution of the reactants and precipitation of products. The liquid phase in CPCs is always water or an aqueous solution. Water solutions ranging from plain water to simulated body fluid have been used. In some cases, some soluble phosphate salts such as NaH2PO4 and/or Na2HPO4 are added as a source of phosphate ions in solution, because it is known that common ions can

Table 10.3 Processing parameters that affect the properties of a calcium phosphate cement Powder phase

Chemical composition Relative proportion of the constituents Additives (seeds, acelerants, retarders, etc.) Particle size distribution of the powder

Liquid phase

Additives (acelerants, retarders, cohesion promoters) pH

Mixing parameters

Liquid/powder ratio Mixing protocol (time, speed, etc.)

Environmental factors

Temperature Humidity pH

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have an accelerating effect on the setting reaction.42 Therefore, neutral salts dissolved in the liquid phase can be used to shorten the setting time exhibited by the cement, for instance in apatitic cements.23 In some cases, watersoluble polymers can be added with the scope of modifying the cohesion of the cement paste or its rheological properties. The liquid/powder ratio is a factor that affects the initial plasticity of the paste and consequently its injectability and setting times. In addition, the final strength is affected by this parameter since the porosity of the set specimen is directly correlated to the liquid/powder ratio used. Therefore, reducing the liquid/powder ratio within the limits of workability would be a means of improving the strength of CPCs. Several properties must be taken into account in relation to the applicability of the CPCs as bone substitutes. Among them, we can mention the setting and cohesion times, the injectability, the hardening rate, the mechanical strength and the pH evolution during setting. All of these properties depend on the composition and the processing parameters that are chosen for each formulation.

10.4.2 The setting time of calcium phosphate cements The time required for the initial setting of the cement paste, which is reflected in a loss of plasticity, is called the setting time. Usually it is measured following mechanical methods, as a quick way of determining whether a reaction occurs upon making a paste of the mixture of reactants with water. The most common methods are based on the assessment of the ability of the cement paste to resist a mechanical load applied to its surface. Two examples are the Vicat needle and the two Gillmore needles. In the first, a single needle is applied on the cement surface. The rationale for the two Gillmore needles, as proposed by Ginebra et al.22 is that with the lightand-wide needle one can measure the initial setting time, which indicates the end of mouldability, without serious damage to the cement structure, whereas with the heavy-and-fine needle one can measure the final setting time, beyond which it is possible to touch the cement without causing serious damage. As far as clinical applications are concerned, the proposed ranges were 4 min < I < 8 min for the initial setting time, I, and 10 min < F < 15 min for the final setting time, F.22 In general, setting times of apatite CPCs are too long and several strategies have to be applied to reach the clinical requirements.43 Among the parameters that can be adjusted to accelerate the setting of apatitic cements are: (a) the liquid/powder ratio (a smaller amount of liquid reduces the setting time); (b) the reduction of the powder size (smaller size leads to shorter setting time); (c) the addition of calcium or phosphate ions either pre-dissolved in the liquid phase or as highly soluble salt (common ion

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effect: the higher the concentration, the shorter the setting time); (d) the addition of seed materials, which act as crystal nuclei (the greater the number of nuclei, the shorter the setting time). On the other hand, brushite CPCs tend to set too fast. The setting time of brushite CPCs is controlled by the solubility of the basic phase: the higher the solubility, the faster the setting time.36 In brushite CPCs setting retarders are often used, and a common approach to increase the setting time is the addition of inhibitors of DCPD crystal growth.44

10.4.3 Cohesion time CPCs are materials designed to be implanted while in a paste state. This means that the paste is in contact with blood or other physiological fluids. Cohesion can be defined as the capacity of a CPC to set in a fluid without disintegrating. It has to be clarified that, in fact, several terms have been used to describe this property – such as non-decay ability, anti-washout, compliance, swelling or stability – and some studies have been devoted to this topic.45–51 In general, this property has been evaluated by an immersion test in water, Ringer’s solution or simulated body fluid. The addition of some water-soluble polymers has been proven to be very effective in enhancing the cohesion of CPC pastes.48,49 However, the approach is very empirical, and in fact there is a need for more understanding of the underlying mechanisms concerning cohesion. It is indeed an important topic since if a CPC has no cohesion at all it will not be able to form a solid body when implanted, or, in the case of poor cohesion, calcium phosphate particles can be released, which can elicit harmful reactions such as inflammation or blood clotting.50,51

10.4.4 Injectability The ability to inject the cement in the surgical site is an important property since it can minimise surgical invasion and allow for complex-shaped defects to be filled adequately. Although all CPCs are mouldable materials, not all of them can be injected, since in some formulations the viscosity of the paste is too high. The injectability of a CPC paste can be defined as its ability to be extruded through a needle without demixing. This will, of course, depend on the diameter and length of the needle (2 mm diameter and 10 mm length can be illustrative values).52 A common problem with CPC injection is demixing or filter pressing, which occurs when the mixing liquid separates from the powder phase, i.e. it is expelled through the needle without the CaP particles. Khairoun et al.53 defined an injectability coefficient as the percentage by weight of the amount of a CPC paste that could be extruded from a syringe with respect to the total mass of the cement introduced in the syringe, when it was extruded at a compression rate of 15 mm min−1 up

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to a maximum force of 100 N. Injectability depends on the rheological properties of the paste. The characterisation of the rheological properties of CPCs is a complex issue, since they have transient rheological properties due to the setting reaction that is taking place from the beginning of the mixing of the liquid and the powder.54 There are several ways of modifying the viscosity of a CPC paste, which in turn will affect its injectability. The most relevant are the liquid/powder ratio, the particle size of the powder phase and the addition of ionic and non-ionic additives. For example, some studies have been performed on the effect of citric acid and its salts on the rheological properties of CPCs.55,56 Citric acid and sodium citrate have been shown to make the surface charge of the particles more negative, acting as a dispersant of the paste and acting as a liquefying agent. This allows a reduction in the amount of water used in the cement, and therefore significantly decreases the porosity and improves the mechanical properties. Another approach is based on the addition of soluble polymers, such as polysaccharides – i.e. sodium alginate, sodium hyaluronate or chondroitin sulphate11,50,57 – or even some polymeric drugs.58

10.4.5 Microstructure and porosity Chemically, the setting reaction of a CPC consists of firstly, the dissolution of one or more constituents of the cement powder and secondly, the precipitation of a different calcium phosphate. Physically, it takes place by the entanglement of the crystals of the precipitating calcium phosphate. A precipitation reaction will only lead to considerable strength in these materials on the following two conditions: (a) the precipitate grows in the form of clusters of crystals that have a fair degree of rigidity and (b) the morphology of the crystals of the precipitate enable the entanglement of the clusters.43 It is interesting to note that, in fact, as previously mentioned, many apatitic cements involve a reaction in water between acidic dicalcium phosphates (DCP or DCPD) and basic TTCP or α-TCP. No water is consumed during the setting of these CPCs, as can be seen in equations [10.2] and [10.3], and liquid is required only to make the reactants workable and to allow a homogeneous reaction. In other cases, when a hydration reaction takes place (see, for instance, equation [10.1]), some water is consumed, but much less than the total amount added to make a workable paste. Hence, water is a major contributor to the origin of porosity in this system, and therefore CPCs are intrinsically porous materials. Figure 10.2 shows the microstructure of an apatitic cement after setting and it can be clearly seen that CPCs develop a highly micro-/nanoporous structure. The porosity of the set CPCs is closely related to the liquid/powder ratio used, and it normally varies between 30% and 50%, although even higher values can be reached. The pores are normally micro- or nanometric in size, and the par-

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10 μm

10.2 Scanning electron microscopy image showing the microstructure of an apatitic cement after setting. Scale bar corresponds to 10 μm.

0.6 dV/d log D

0.5 0.4 0.3 0.2 0.1 0 0.001

0.01

0.1

1

10

100

1000

Entrance pore diameter (μm)

10.3 Hg-porosimetry diagram of an apatitic CPC after setting. The starting powder consists of α-TCP and the liquid/powder ratio is 0.35 ml/g, showing open porosity centred around 10 nm, and pore size below 0.5 μm. This CPC had a total porosity of 32%. dV/d log D is the log differential intrusion volume versus diameter.

ticle size distribution of the starting powder can modify the size of the precipitated crystals, and also the pore size distribution.25 A typical pore size distribution diagram of an apatitic CPC obtained by Hg-porosimetry is shown in Fig. 10.3.

10.4.6 Mechanical strength For adequate performance of CPCs as biomaterials, some physical properties such as compressive strength, diametral tensile strength and fracture

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toughness may be important. The main mechanical property has been described to assess the quality of CPCs is the compressive strength.59 Since these materials are conceived as bone substitutes, it is important to keep in mind as reference values that, according to different studies, the compressive strength of human cortical bone ranges between 90 and 209 MPa,60,61 and that of the cancellous bone ranges between 1.5 and 45 MPa.62 Owing to the intrinsic porosity of CPCs, their strength is lower than that of calcium phosphate ceramics.13 The liquid/powder ratio, the particle size of the reactants, the crystallinity and amount of seed, and the use of liquid accelerators are factors that affect the strength of the CPCs. A wide range of values can be found in the literature, depending on the composition and processing parameters, and it is difficult to make comparisons between them due to a lack of consistency in specimen dimensions, testing protocols and sample pre-treatments. As indicative values, the compressive strength of apatite cements normally ranges between 20 and 50 MPa,8,20,25,46–49 although lower and higher values for some formulations have also been reported.63 The compressive strength of some commercial formulations of apatitic CPCs are summarised in Table 10.4. Brushite CPCs are, in general, weaker than apatite CPCs and compressive strengths of 25 MPa have been reported.64 Different models have been proposed to describe the strength dependence on porosity, in ceramics. Among them, Ryshkewitch65 claimed that strength varies as an exponential function of porosity, as described by s = s0 exp(−bP)

[10.5]

Table 10.4 Compressive strength of some commercial apatitic calcium phosphate cements

Cement name

Company

Powder composition

Compressive strength (MPa)

α-BSM Biopex

ACP, DCPD α-TCP, TTCP, DCPD, HA

12 60–90

Cementek

ETEX Mitsubishi materials Tecknimed

15–25

Calcibon Mimix Mimix QS Norian SRS Norian CRS

Biomet Biomet Biomet Synthes-Norian Synthes-Norian

α-TCP, TTCP, Na-glycerophosphate α-TCP, DCP, CaCO3, PHA TTCP, α-TCP, C6H5O7Na3·2H2O — α-TCP, CaCO3, MCPM —

60 24 22 50 30

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where s is strength, s0 is strength at zero porosity, P is porosity and b is a parameter characteristic of each material. This model was fitted to the compressive strength of apatitic CPCs with different liquid/powder ratios,66 and a value of 408 MPa was obtained for s0, which is in good agreement with the compressive strength reported in the literature for dense, sintered hydroxyapatite.67 The reduction of porosity in CPCs has been explored as a way of increasing their strength. Reducing the amount of added water and improving particle packaging can reduce the porosity of the cements. Compaction of the cement paste during setting has been demonstrated to increase the compressive strength of apatitic CPCs,68–70 and values as high as 118 MPa have been reported.70 The addition of water-reducing or liquefying agents, such as sodium citrate, allows for a further densification of the paste and compressive strength values of 180 MPa in wet conditions have been reported.70 However, it has to be considered that this compaction cannot be applied if the cement is implanted or injected within the bone tissue, and therefore at present it can be used only to fabricate pre-set substrates or scaffolds for bone regeneration. A two-step protocol, including precompaction of a paste followed by a conventional application has been suggested as an alternative for potential clinical use.70 The evolution of the strength of the CPC after implantation has also been studied. The mechanical properties of apatite CPCs are reported to increase,71 while those of brushite CPCs tend to decrease72 due to the higher solubility of DCPD compared with that of PHA. Only after a few weeks of implantation, when bone growth is significant, do the mechanical properties of brushite CPCs increase.72

10.4.7 In vivo behaviour Since their discovery, numerous studies have been devoted to the characterisation of the in vivo behaviour of CPCs, showing that they are highly biocompatible and osteoconductive materials, which can stimulate tissue regeneration.8,35,73–82 Most of the apatitic cements are resorbed via cellmediated processes. In these processes osteoclastic cells degrade the materials layer by layer, starting at the bone–cement interphase throughout its inner core. The biodegradability of apatite CPCs is greater than that of sintered hydroxyapatite, but it is still slow. It has been shown, for instance, that some CPCs could remain for as long as 78 weeks when implanted in dog femurs.78 There are several factors that affect the degradation rate of apatitic CPCs, with the crystallinity of the PHA being especially relevant, and also the specific surface area and the porosity of the set cement. One strategy proposed in recent years in a bid to accelerate resorption of apatite

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CPCs is the incorporation of macroporosity, and this will be described in the next section. Brushite CPCs are resorbed much faster than apatitic cements,79–81 owing to the fact that brushite is metastable in physiological conditions. However, it has been reported that brushite CPCs tend to transform to PHA in vivo, this transformation reducing its overall degradation rate. The addition of magnesium salts can be used to avoid, or at least delay, this transformation.82

10.5

Applications of calcium phosphate cements: present and future perspectives

10.5.1 Present applications The first commercial apatitic CPCs were introduced in the market in the 1990s, and subsequently brushite cements were also commercialised. They are used for different bone regeneration applications, such as: (a) maxillofacial and craniofacial reconstruction (cranioplasty, cranial recontouring, cranial flap augmentation, augmentation genioplasty, on-lay grafting, skull base defect repair);83 (b) treatment of several fracture defects – such as distal radius, proximal and distal tibia, calcaneus, proximal and distal femur, proximal humerus, acetabulum;84 (c) treatment of surgically or traumatically created osseous defects, filling of cystic lesions and augmentation of screws; (d) more recently, they are being used for the treatment of spinal fractures and vertebroplasty (with or without the aid of a kyphoplasty balloon).85–87

10.5.2 Macroporous calcium phosphate cements As mentioned in Section 10.4.7, although apatitic cements have the ability to slowly be replaced by bone, one of the weaknesses in their clinical performance is their slow rate of resorption. The introduction of macroporosity in CPCs is envisaged as a method to facilitate bone ingrowth not only from the external surface but throughout the whole bulk of the material. This would accelerate its resorption and its transformation in newly formed bone tissue. Presently, two different strategies have been adopted to introduce macroporosity in CPCs. The first approach aims to produce the macropores after the setting of the cement. Different porogenic agents have been suggested, which are added within the CPC paste and, after the setting, degrade faster than the cement itself, creating the macroporosity; for example, sugars,88,89 PLA fibres or particles90,91 or frozen sodium phosphate solution particles.92 However, it is necessary to add a large amount of porogenic

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1 mm

10.4 Scanning electron microscopy image of a macroporous CPC. Scale bar corresponds to 1 mm.

agent to guarantee interconnectivity of the porosity, thus compromising the excellent bioactivity and biocompatibility of CPCs. In a second approach, macroporosity is created before the setting of the cement. The cement paste is foamed while it has a viscous consistency, and on setting a solid macroporous construct is created. The macroporosity can be produced by two main routes: (a) the addition of some gas-generating compounds, such as hydrogen peroxide93 or sodium bicarbonate,94–96 although it has to be taken into account that the liberation of gas after the implantation of the cement paste could have harmful effects for the organism; (b) the use of biocompatible foaming agents.97,98 This last approach has allowed the development of injectable macroporous CPCs, which maintain the macroporosity after injection.97 Figure 10.4 shows the macroporous structure of a foamed CPC. In vivo studies have shown that this strategy can be effective in accelerating the resorption of CPCs.98

10.5.3 Pre-set granules and bone tissue engineering scaffolds CPCs can be used to fabricate pre-set granules and blocks, which can have some advantages in comparison with ceramic granules or blocks. The advantages arise from the low-temperature and wet processing method intrinsic to CPCs. Indeed, the apatite CPCs products are micro-/nanocrystalline and have high specific surface area, and therefore are much more reactive than sintered ceramics. In addition, CPCs enable the fabrication of lowtemperature calcium phosphates – such as DCPD, octacalcium phosphate

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(OCP) or CDHA – which cannot be obtained by high-temperature sintering, and which are much more close to the calcium phosphates found in living tissues. In this context, CPCs have been used to fabricate pre-set hydroxyapatite porous blocks to act as scaffolds for guiding in vitro or in vivo tissue regeneration, which in fact is one of the main goals of tissue engineering.92,93 These three-dimensional macroporous constructs satisfy several requirements, such as osteoconductivity, adequate mechanical properties, formability and high interconnected macroporosity to ensure cell colonisation and flow transport of nutrients and metabolic waste. In addition, the apatite foams combine the interconnected macroporosity with the intrinsically high micro/nanoporosity of CPCs.93 Recently it has also been shown that CPCs can be used in rapid prototyping techniques, and specifically low-temperature three-dimensional powder printing can be used to fabricate calcium phosphate structures with simultaneous control of geometry and organic molecule incorporation in three dimensions.99

10.5.4 Drug delivery Although most drug-delivery materials are polymeric in nature, calcium phosphate-based materials have an added value owing to their bioactive character in the specific field of the pharmacological treatment of skeletal disorders. Moreover, another important property of calcium phosphates is their unique ability to adsorb different chemical species on their surfaces. This property has been exploited in hydroxyapatite chromatography, which has proved to be very efficient for the purification and separation of proteins, enzymes, nucleic acids and other macromolecules. This great affinity of hydroxyapatite for these various active molecules can extend the application of CPCs not only as bone substitutes, but as carriers for local and controlled supply of drugs in the treatment of different skeletal diseases – such as bone tumours, osteoporosis or osteomyelitis – which normally require long and painful therapies.100 Unlike calcium phosphate ceramics employed as drug-delivery systems, where the drugs are usually absorbed on the surface, in CPCs the drugs can be incorporated throughout the whole material volume, by adding them into one of the two cement phases. This fact can facilitate the release of drugs for a more prolonged period. Certain factors need to be taken into consideration with reference to the incorporation of drugs in CPC cements. In the first instance, it is necessary to verify that the addition of the drug (either to the liquid or the solid phases of the cement) does not interfere with the setting reaction, modifying the physico-chemical properties, not only in terms of the setting and hardening mechanisms, but also with respect to the rheological behaviour. Secondly, it is necessary to characterise the

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kinetics of drug release in vitro. Subsequently, the effectiveness of the cement in acting as a carrier for drug delivery in vivo must be assessed. Finally, the clinical performance of the drug-delivery system must be evaluated. To date, the first three factors have been extensively studied, but the application of CPCs as drug-delivery systems has not yet reached clinical applications. Over the last decade, several studies related to the application of both commercial and experimental CPCs as drug carriers have been published.101 Major attention has been paid to antibiotics, due to their wide area of application: either as prophylactics to prevent infections produced during surgical interventions, or in general in the treatment of bone infections. Other types of drugs incorporated in cements include anti-inflammatory drugs and anticancer drugs, and even hormones have been studied. In recent years the inclusion of growth factors that are able to stimulate bone regeneration, such as bone morphogenetic proteins (BMPs) or transforming growth factor β (TGF-β) have been considered for controlled delivery from CPC cements. The research that has been carried out in the past, especially the in vitro studies, has highlighted the great potential of CPCs as carriers for controlled release and vectoring of drugs in the skeletal system. However, the industrial use of CPCs for drug delivery is not yet a reality, and in fact two main underlying problems can be identified. First, implant companies selling CPCs do not generally have the expertise to deal with drugs, and pharmaceutical companies do not have any expertise with CPCs. Secondly, infections are not always produced by the same microorganisms and, therefore, it would be necessary to design versatile systems that could combine a given CPC with many different drugs, in such a way that the surgeon could choose the drug just before implantation. However, as mentioned previously, various drugs have various effects on CPC properties, and this represents a serious drawback for the implementation of the technology. Therefore, a lot of work has still to be done to be able to adjust the use of CPCs to different therapeutical needs and to obtain reproducible and predictable drug-delivery systems.

10.6

References

1 ginebra mp, gil fj, planell ja. Acrylic bone cements, in R Barbucci (Ed.), Integrated Biomaterial Science. Kluwer Academic/Plenum Publishers, New York (2002), pp. 569–588. 2 kühn kd. Bone Cements. Springer Verlag, Berlin, Heidelberg (2000). 3 driessens fcm, fernández e, ginebra mp, boltong mg, planell ja. Calcium phosphate and ceramic bone cements vs. acrylic cements. Anal Quím Int Ed 1997 93, 538–543.

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Part III Properties of bone cement

11 Mechanical properties of bone cements N. D U N N E, Queen’s University Belfast, UK

Abstract: This chapter discusses the mechanical properties of poly(methylmethacrylate) (PMMA) bone cements as used for joint replacement surgery. The chapter first reviews the structure of PMMA and the different test standards that are available to determine the mechanical properties of the cured bone cement. The chapter then discusses the relationship between the PMMA cement microstructure and its mechanical properties. The chapter concludes by identifying ways of modifying the mechanical performance of conventional PMMA bone cements. Key words: acrylic bone cement, mechanical properties, joint replacement surgery, standard test techniques.

11.1

Introduction

Poly(methylmethacrylate) (PMMA) bone cement has been primarily used for the fixation of joint prostheses in humans for more than 50 years. The self-curing PMMA bone cement fills the free space between the prosthesis and the bone and constitutes a crucial interface (Fig. 11.1). Owing to their optimal rigidity, the bone cements can evenly cushion the forces acting against the bone. The close association between the cement and the bone leads to optimal distribution of the stresses and interface strain energy. The transfer of the forces bone-to-implant and implant-to-bone is the primary function of the bone cement. In order to carry out this function, bone cement must be compatible with the tissue it contacts and have adequate strength. Since the forces transmitted through the hip are high (Bergmann et al. 1993, Berme & Paul 1979) – c. three times body weight when walking, increasing to eight times body weight when falling – bone cement is subjected to high stresses and has to perform in the aggressive environment of the body. The ability to do this reliably for a long time is critical for long-term survival of the implant. An adequate cement interdigitation and reinforcement of the cancellous bone are of utmost importance. If the continuous stress from outside exceeds the capability of the bone cement to transfer and absorb forces, a fatigue break is possible. The mechanical properties of PMMA bone cement can be divided into two parts: short-term and long-term properties. Short-term properties 233

234

Orthopaedic bone cements Acetabular component Femoral prosthesis

Reamed out socket of hip joint

Bone cement

Cement in tension Cement in compression

11.1 Schematic diagram of prostheses and PMMA bone cement in an acetabular socket and femur.

demonstrate that the PMMA bone cement is weak in tension and strong in compression. Long-term properties (e.g. fatigue, creep and stressrelaxation) are considered in greater detail as these properties can significantly influence the transfer of loads into the bone over the normal working life of a total joint replacement (TJR). The effects of the physiological environment in which the cement performs, and the design characteristics of the prosthesis being implanted, on the performance of the bone cement are described in this chapter.

11.2

Nature and structure of polymethylmethacrylate bone cement

In order to appreciate the physical behaviour of PMMA bone cement in TJR, it necessary to understand the basic nature and structure of the material (Ashby & Jones 1988). PMMA bone cements are two-component systems, comprising a polymer powder (Fig. 11.2) and a liquid monomer (Fig. 11.3), the cured cement is produced by a free radical polymerisation (Fig. 11.4) lasting 6–14 minutes depending on the bone cement formulation, mixing conditions and the environment. The resultant PMMA bone cement is a cured polymer comprising long-chain molecules made up of many thousands of units called monomers. These long-chain molecules are twisted around each other in a random amorphous structure and consist of strongly bonded covalent chains with side groups that are linked by relatively weaker secondary bonds. PMMA bone cement is a thermoplastic polymer; that is, its behaviour and properties change as the temperature of the polymer varies. The temperature most frequently cited is the glass transition temperature. Above this temperature (which is c. 95°C for PMMA bone cement),

Mechanical properties of bone cements Methyl methacrylate (MMA) CH3 H2C

C COOCH3

Clear, colourless liquid of strong odour Boiling point: 100 °C Density: 0.943 gcm−3 at 20 °C Vapour pressure: 38 hPa at 20 °C Molecular weight: 100 gmol−1 Odour threshold: 0.2 ppm

11.2 Chemical composition and properties of methyl methacrylate (MMA).

Polymethyl methacrylate (PMMA) CH3 R

(CH2

CH3 CH2

C

COOCH3

C

CH3 CH2

COOCH3

C)n

R

COOCH3

Fine powder, polymer beads Bead diameter: 1–120 μm Soluble in liquid monomer Density: 1.18 g cm−3 Molecular weight: 0.8 g mol−1

11.3 Chemical composition and properties of PMMA.

CH3

CH3 Addition of many monomers

CH2

C

CH2 C

C

O

O

n C

O CH3

11.4 Polymerisation reaction of PMMA.

CH3

O

235

236

Orthopaedic bone cements (a)

(b)

Three-point bending

Three-point bending

11.5 (a) Bone cement specimen under a constant load for 24 hours at 37°C. (b) Bone cement specimen under the same load for 24 hours at 18°C (Lee 2000).

the polymer has a significantly lower modulus, with the long-chain molecules more free to move over each other, giving the polymer leathery properties. There are other transitions (secondary transitions) that are important, and for PMMA bone cement one occurs at c. 25°C – above room temperature but below body temperature. The secondary transition is a consequence of the relaxation of the weaker bonds between the side groups, allowing the side groups to rotate over each other. The result of this secondary transition is that the PMMA bone cement is more flexible under body conditions than it is at room temperature (Fig. 11.5). The importance of the temperature dependency of PMMA bone cement is that any tests that are designed to demonstrate how bone cement functions in an implant system must be conducted at body temperature (37°C) or slightly higher.

11.3

Test standards

The International Organization for Standardization (ISO) standard 5833 (ISO 2002), which was first published in 1979 and later revised in 2002, is a standard that describes tests and minimal requirements for acrylic-based bone cements. The ISO standard specifies the requirements for the liquid monomer, the polymer powder and the dough. There are some special

Mechanical properties of bone cements ISO 5833

3500

Cross-head speed 19.8–25.6 mm min−1

3000 2500 Load (N)

6 mm

237

2000 1500 1000 500

12 mm

0 0.00

0.50

1.00 1.50 2.00 2.50 Displacement (mm)

3.00

3.50

Minimum ISO requirement Compressive strength >70 MPa

11.6 ISO 5833 test method to determine the compressive properties for PMMA bone cement.

requirements concerning the packaging and criteria for the cured bone cement. There are three requirements for the cured bone cement: compressive strength (minimum requirement of 70 MPa), bending strength (minimum requirement of 50 MPa) and bending modulus (minimum requirement of 1800 MPa). The compressive strength is tested on cylindrical bone cement specimens 24 ± 2 hours after curing and storage in dry air at 23°C (Fig. 11.6). The test is carried out under a constant cross-head speed of between 19.8 and 25.6 mm min−1. The strength is determined from maximum force applied to cause fracture, 2% offset or upper yield point load, whichever occurs first. The bending strength and bending modulus are determined using a four-point bending test arrangement on beam bone cement specimens 24 ± 2 hours after curing and storage in dry air at 23°C (Fig. 11.7). The test is carried out under a constant cross-head speed of 5 ± 1 mm min−1. Formulae are given in the standard for the calculation of the bending strength and bending modulus. All commercially available bone cements must satisfy the requirements in this standard. Patients, operating theatre staff and surgeons need to be confident that compliance with the requirements of the international standard is adequate to guarantee that the bone cement being used clinically will have acceptable long-term performance. Unfortunately, the test procedures and minimum criteria for acceptance are set at a low level and can be easily fulfilled. Therefore, it is difficult for ISO 5833 to distinguish if a particular type of bone cement is wholly suitable for clinical use or not. The Swedish National Hip Arthroplasty Register in 2002 reported that the most prevalent cause of revision of a total hip joint was aseptic loosening of the implant component (Malchau et al. 2002). Given that c. 93% of total hip

238

Orthopaedic bone cements ISO 5833 Cross-head speed 5.0 mm min−1 75 mm

Load (N)

140 120 100 80 60 40 20 0 0

3.3 mm

10 mm

Minimum ISO requirement Bending strength >50 MPa Bending modulus >1800 MPa

1

2 3 4 5 Displacement (mm)

6

7

11.7 ISO 5833 test method to determine the bending properties for PMMA bone cement.

replacements performed in Sweden are cemented, it can be presumed that aseptic loosening is the most common cause of failure of cemented implant components. Aseptic loosening of a cemented implant most likely indicates failure by shear or tension of the bone–bone cement interface, or failure after the generation of wear debris, together with lubrication effects within the joint, causing osteolysis. Hardly ever, is failure of cemented implant components as a result of compression or bending of bone cement. Therefore, the requirements of the ISO 5833 standard, while being practical for qualitative purposes, are not adequate to make sure that PMMA bone cements are suitable for use. Currently, a simple tensile test is not included in ISO 5833. PMMA bone cement is particularly weak in tension, but relatively strong in compression. In the 2002 version of ISO 5833 a test to determine the bending properties was included. Bending does include a tensile element, but all commercially available bone cements can easily achieve the minimum requirements of this test. Most notably, there is no test for the determination of the fatigue properties of bone cement in the ISO standard. This type of testing has recently been included in the American Society for Testing and Materials (ASTM) standard F2118-2001 (ASTM 2001). This ASTM standard documents a procedure for a fully reversed tensile compression cyclic loading test of PMMA bone cement. However, the standard does not indicate a minimum requirement. A fatigue test is a very time-consuming and expensive test. The test results will be highly dependent on the bone cement composition,

Mechanical properties of bone cements

239

mixing conditions and ultimately the porosity of the cured bone cement specimens.

11.4

Mechanical properties: short-term strength of polymethylmethacrylate bone cement

11.4.1 Compressive properties An overview of several bone cements and their compression properties is given in Table 11.1. Bone cements based on more brittle PMMA polymers appear to have a higher compressive strength and modulus of elasticity than bone cements based on more flexible molecules such as Palacos® R. Bone cements with low liquid monomer concentration (e.g. polymer powder : liquid ratio of 3) such as CEMEX RX demonstrate lower compressive strength after polymerisation than bone cements with conventional liquid monomer concentrations (e.g. polymer powder : liquid ratio of 2 : 1), such as CMW 3. In all cases, the compressive strength is significantly greater than the physiological compressive stress levels, which are of the order of 5 MPa. Commercial bone cements loaded with antibiotics exhibit similar compressive properties to plain bone cements.

11.4.2 Bending properties The bending strength and bending modulus for several commercial bone cements are shown in Table 11.2. According to ISO 5833 the minimum Table 11.1 Mean (with standard deviation (SD) in brackets) compressive properties of bone cements derived from compression tests according to ISO 5833

Cement type

Failure stress (MPa)

Failure strain (%)

Modulus of elasticity (MPa)

CEMEX RX CMW 1 CMW 1 + G CMW 3 CMW 3 + G CMW Endurance Palacos® R Palacos® R + G Palamed Palamed + G SmartSet HV Surgical Simplex P

102.30 96.52 94.67 98.80 95.10 93.20 86.91 78.91 95.29 88.45 86.54 90.32

6.95 7.13 7.05 6.89 6.83 7.03 6.89 6.40 6.97 6.70 6.83 6.94

2689 2791 2287 2573 2261 2853 2891 2401 2820 2541 2691 2851

(5.6) (3.8) (4.8) (3.9) (4.9) (3.9) (3.9) (4.8) (4.2) (3.8) (3.5) (4.1)

(0.7) (0.8) (0.9) (0.6) (0.5) (0.7) (0.9) (0.7) (0.5) (0.8) (0.8) (0.9)

(110) (178) (209) (156) (95) (209) (267) (186) (340) (203) (259) (302)

Compression test at 24 hours after mixing under atmospheric conditions.

240

Orthopaedic bone cements

Table 11.2 Mean (SD) bending properties of bone cements derived from bending tests according to ISO 5833 Cement type

Bending strength (MPa)

Bending modulus (MPa)

CEMEX RX CMW 1 CMW 1 + G CMW 3 CMW 3 + G CMW Endurance Palacos® R Palacos® R + G Palamed Palamed + G SmartSet HV Surgical Simplex P

56.71 67.81 65.81 71.90 69.83 74.81 75.67 71.56 71.23 64.12 64.32 68.45

2481 2687 2503 2791 2751 2891 2761 2691 2687 2553 3010 2681

(3.8) (4.5) (4.5) (3.8) (3.6) (2.2) (3.9) (4.0) (3.9) (4.6) (2.1) (2.9)

(178) (259) (264) (178) (109) (248) (197) (253) (226) (178) (332) (206)

Bending test at 24 hours after mixing under atmospheric conditions.

requirement for the bending strength is 50 MPa and for the bending modulus it is 1800 MPa. The addition of antibiotics reduces the bending strength, but the differences between antibiotic-loaded bone cement and plain bone cement are not always statistically significant. The bending modulus represents the ratio of stress to corresponding strain of the material within the elastic region. Stiff materials demonstrate a high modulus and ductile materials exhibit a low modulus. Within the elastic region the stress and strain are directly proportional following Hooke’s Law, and, if the load is released, the material regains its original dimensions. The elastic region is limited by a stress limit, the proportional limit at which the physical properties of the material actually change and the material might not recover its initial shape after releasing the load. As already mentioned, the bone cement acts as a mechanical cushion. For this function the modulus of elasticity of the bone cement must be lower than the moduli of the metallic implant components and the bone. The modulus must not decrease to below a minimum value, therefore a lower limit for the modulus of bone cement has been determined in ISO 5833. The elastic modulus varies with temperature, which means the higher the temperature, the lower the modulus. Testing bone cement at 22°C is not really a suitable way to get relevant results for applications associated with the human body.

11.4.3 Tensile properties Although bone cement demonstrates good compressive properties, it is susceptible to fracture that might be caused by tensile loading (Harper &

Mechanical properties of bone cements

241

ISO 527 Cross-head speed 20.0 mm min−1

4 mm

150 mm

10 mm

Minimum ISO requirement Tensile strength – no minimum value

11.8 ISO 527 test method to determine the tensile properties for PMMA bone cement.

Bonfield 2000). There is no explicit standard for static tensile testing of PMMA bone cement. However, general standards for tensile testing of plastics do exist: DIN (Deutsches Institut für Normung) 53455 (DIN 1981), ISO 527-1 (ISO 1993) and ASTM D636 (ASTM 2000). These standards describe a static test method applicable for all general-purpose plastics (Fig. 11.8). As the preparation and conditioning of the specimens is not described in detail, it is often difficult to compare test results from different studies. Table 11.3 shows the results of tensile tests of bone cement specimens made in accordance with DIN 53455 and ISO 527-1, stored in air at 22°C. The tensile strength ranged between 32 and 56 MPa. There is a strong correlation between the chemical formulation of the bone cement and the mechanical properties (Table 11.4). The bone cements based on copolymers with ductile methacrylate (MA) – molecules (for example, CMW SmartSet, Palacos® R and Palamed) demonstrate the highest tensile strength, while the more brittle bone cements based on co-polymers with styrene molecules and PMMA homopolymers exhibit the lowest tensile strength. The more ductile MA co-polymers have strain to failure of ±2.2%, which is around double the failure strain of the weaker and more brittle styrene co-polymers and MMA homopolymers.

242

Orthopaedic bone cements

Table 11.3 Mean (SD) tensile properties of bone cements derived from tensile tests according to ISO 527-1 and DIN 53455 Cement type

Tensile strength (MPa)

Tensile modulus (MPa)

CEMEX RX CMW 1 CMW 1 + G CMW 3 CMW 3 + G CMW Endurance Palacos® R Palacos® R + G Palamed Palamed + G SmartSet HV Surgical Simplex P

32.0 42.1 38.2 43.8 35.9 39.4 51.4 47.9 56.0 48.5 51.6 52.3

3098 2956 2748 3420 3061 3028 3217 3020 3412 3281 3103 3379

(2.0) (2.6) (3.8) (3.8) (2.9) (3.9) (3.6) (3.1) (2.7) (3.9) (3.7) (2.6)

(180) (260) (170) (172) (270) (210) (241) (218) (289) (161) (238) (267)

Bending test at 24 hours after mixing under atmospheric conditions. Table 11.4 Glass transition temperature of different types of acrylic molecules Elasticity

Molecule

Glass transition temperature (°C)

Brittle

Styrene Methyl methacrylate Ethyl methacrylate Butyl methacrylate Methylacrylate

120 105 65 20 6

Ductile

11.4.4 Other static mechanical test methods There are definite differences in the resulting data of the three static test methods described. The compressive strength of PMMA bone cement is greater than the bending strength and that is higher than the tensile strength. This order is observed for all general-purpose plastics. This means that tensile loading may be a greater risk factor for failure than compressive loading. In vivo, simple tensile loading, however, does not play a major role; complex combinations of different modes of loading are more appropriate. From a physical point of view, bending combines tensile and compressive loading; therefore, the bending test is the most pragmatic test. Another static test applied to bone cements is a shear strength test according to ASTM D732 (ASTM 2002). The shear strength is determined as the ratio of shear stress to shear strain. The cement test specimen, which can be either a 50 mm diameter disc or a 50 mm square, is placed in a clamp such that its upper and lower surfaces are supported (Fig. 11.9). A punch-

Mechanical properties of bone cements

243

ASTM D732

0.127–12.7 mm 11 mm 50 mm

11.9 ASTM D732 test method to determine the shear properties of plastics.

Table 11.5 Mean shear strength of bone cements derived from shear tests according to ASTM D732 Cement type

Shear strength (MPa)

CMW 1 (hand mixed) CMW 1 (vacuum mixed) Palacos® R (hand mixed) Palacos® R (vacuum mixed) Surgical Simplex P (hand mixed) Surgical Simplex P (vacuum mixed)

32 62 33 50 32 63

type shear tool with a 25.4 mm diameter is bolted to the specimen and a load is applied to the punch. The shear strength is calculated as the maximum force encountered during the test divided by the area of the sheared edge (circumference of the punched circle multiplied by the specimen thickness). Table 11.5 shows the results of shear tests of bone cement specimens made in accordance with ASTM D732, stored in water at 37°C for at least 24 hours (Kindt-Larsen et al. 1995). The shear strength ranged between 32 and 63 MPa depending on the method of cement preparation, that is hand or vacuum mixing. Shear strength is an important mechanical parameter because debonding of the implant–cement interface has been known to initiate the failure of cement femoral prostheses. The interfacial static shear strength is influenced by surface roughness, cement type and porosity. Surface finish or texture has the strongest influence on the interfacial strength at the implant–cement interface. However, increasing the surface

244

Orthopaedic bone cements

ASTM D5045-99: Single edge notched three-point bend(SENB)

ASTM B771-87: Chevron notched short rod (CNSR)

B W

a

t

S 2.2W 22W Nominal dimensions W = 16 mm, a = 8 mm, S = 64 mm, ASTM D5045-99: Rectangular compact tension (RCT) 0.5W

∅ W/4

12W

0.275W

Nominal dimensions B = 4 mm W = 8 mm a = 2.83 mm t = 0.12 mm q = 55°

End view a q

W

B Cross-section Side view

a W

Nominal dimensions W = 32 mm a = 16 mm

11.10 ASTM D5045 and ASTM B771 test methods to determine the fracture properties of materials.

roughness of an implant component beyond a threshold value has no additional effect. Bone cement type and porosity have only a negligible effect on the static interfacial strength at the implant–cement interface (Wang et al. 2003). Moreover, there are test methods to determine the fracture properties, such as fracture toughness (ASTM B771-87 (ASTM 1997); ASTM D504599 (ASTM 2007)) (Fig. 11.10) and impact strength (ISO 179 (ISO 1997); ISO 180 (ISO 2000); DIN 53435 (DIN 1983)) (Fig. 11.11). Bone cements are accepted as linear elastic solids, since PMMA behaves as a brittle material, the fracture properties can be easily determined (Sih & Bernam 1980). There is a strong relationship between fracture toughness and impact strength of PMMA bone cement (Lewis & Mladsi 2000). For a given specimen geometry, γ-irradiation produced a statistically significantly reduction in fracture toughness because of the related decrease in molecular weight (Lewis 1999a). Lewis (1997) evaluated the fracture toughness of PMMA bone cements using different types of specimen. He concluded that the bone cement type, method of mixing, moulding conditions and type of test performed play a significant role in the fracture results obtained (Lewis 1997). Impact strength is an assessment of the energy needed to cause a

Mechanical properties of bone cements ISO 179: Charpy impact

ISO 180: Izod impact h1 Impact direction

b

Impact direction b

a h2

S2 W2

245

W1

a

Nominal dimensions W2 = 80 mm S2 = 62 mm a = 1.2–3.2 mm b = 4 mm h2 = 10 mm

S1

Nominal dimensions W1 = 50 mm S1 = 19 mm a = 1.2–3.2 mm b = 4 mm h1 = 6 mm

11.11 ISO 179 and ISO 180 test methods to determine the impact properties of plastics.

Table 11.6 Mean (SD) fracture toughness properties of bone cements derived from fracture tests according to ASTM B771-87 and ASTM D5045-99 Cement type

Test specimen type

Fracture Toughness (Mpa m½)

CMW 1 CMW 3 CMW Endurance Palacos® R Palacos® R Palacos® R SmartSet HV Surgical Simplex P Surgical Simplex P Surgical Simplex P

CNSR CNSR CNSR SENB CNSR RCT CNSR SENB CNSR RCT

1.50 1.65 1.21 1.76 1.39 1.85 1.30 1.52 1.23 1.50

Fracture test at 24 hours after mixing under atmospheric conditions. CNSR, chevron notch short rod; SENB, single edge notched, three-point bend; RCT, rectangular compact tension.

material to fracture when struck by an abrupt blow. The addition of additives (radiopacifiers and antibiotics), together with pores in the bone cement, may have a destructive influence on the performance of the bone cement under sudden impact loads (De Wijn et al. 1975, Kühn 2000). Table 11.6 shows the results of fracture toughness tests of bone cement specimens carried out in accordance with ASTM B771 and ASTM 5045, stored in air

246

Orthopaedic bone cements

Table 11.7 Mean (SD) impact properties of bone cements derived from impact tests according to DIN 53435: 1993 Cement type

Impact strength (kJ m−2)

CEMEX RX CMW 1 CMW 1 + G CMW 3 CMW 3 + G CMW Endurance Palacos® R Palacos® R + G Palamed Palamed + G SmartSet HV Surgical Simplex P

5.37 3.68 3.42 2.87 3.10 3.26 6.81 4.64 5.50 4.61 6.93 4.10

(0.13) (0.11) (0.08) (0.10) (0.17) (0.16) (0.21) (0.13) (0.18) (0.12) (0.19) (0.15)

Impact test at 24 hours after mixing under atmospheric conditions.

at 22°C for at least 24 hours. Table 11.7 shows the results of impact strength tests of bone cement specimens carried out in accordance with DIN 53435, stored in air at 22°C for at least 24 hours.

11.4.5 Creep behaviour PMMA bone cements demonstrate a combination of elastic and viscous behaviour called viscoelasticity. When a polymer is subjected to a constant load, the resulting deformation can be divided into two components: the immediate elastic deformation and the time-dependent, continuous deformation. The immediate elastic deformation happens directly on application of a load. It is a recoverable deformation that is independent of time. Subsequent to this fast deformation, there is a deferred continuous deformation resulting from stress. One element of this deformation is recoverable in time after releasing of load. This component is called primary creep. The second part of this continuous deformation is a non-recoverable permanent deformation called secondary creep. Different test methods are described in ASTM D2990 for the measurement of creep (ASTM 2001). The test specimen is loaded either in compression, tension or bending. All plastics materials including PMMA bone cement creep to some degree. The level of creep seems to correlate with the molecular weight distribution of PMMA bone cement. It has been reported that creep resistance increases with density and that large PMMA powder particle size, residual MMA monomer, radiopaque agents and plasticising environments such as water uptake decrease it (Treharne & Brown 1975). Creep will also depend upon the polymerisation process of the bone

Mechanical properties of bone cements

247

cement and the temperature at which the test is conducted (Oysaed & Ruyter 1989). Different experimental models to represent the creep behaviour of bone cement have been recommended (Norman et al. 1975, Verdonschot & Huiskes 1995). It has been postulated that creep of PMMA bone cement may be a causal factor in the loosening of cemented implant components. However, bone cement creep relaxes cement stresses and creates a more beneficial stress distribution at the interfaces (Verdonschot & Huiskes 1997, 2000). Belated injection of the PMMA bone cement increased creep compared with bone cement applied according to standard injection protocols (Norman et al. 1997). Consequently, creep not only depends on the material properties, but also on the cement management and delivery by the theatre scrubs nurses and orthopaedic surgeons.

11.4.6 Fatigue behaviour To ensure the survival of PMMA bone cement in the human body, the cement must be able to withstand the varying loads it bears. The fatigue properties of the PMMA bone cement are therefore of significant importance, and they may determine when a properly used bone cement will fail. Many studies have investigated the fatigue properties of PMMA bone cement in different ways (Krause & Mathis 1988, Lewis 1997, 1999b, 1999c, Eveleigh et al. 2002, Dunne et al. 2003). Currently, there are three testing protocols (Fig. 11.12) used to characterise the fatigue behaviour: the Four-point bending

Tension Tension–compression 4 mm 10 mm

75 mm

5 mm

62 mm

150 mm

3.3 mm 10 mm

10 mm

ISO 5833

Tests normally conducted at 2–5 Hz in buffered saline solutions at 37°C

ISO 527

ASTM F2118

11.12 ISO 583, ISO 527-1 and ASTM F2218 test methods used to characterise the fatigue behaviour for PMMA bone cement.

248

Orthopaedic bone cements

four-point bending arrangement recommended by ISO 5833 (ISO 2002); the uniaxial pure tensile test according to ISO 527 (ISO 1993); the uniaxial compression–tension test according to ASTM F2118 (ASTM 2001). The first test method is comparable with the bending test method according to ISO 5833. Fatigue testing is a dynamic test and is carried out with a sinusoidal cyclic loading under stress control. The test continues until failure or run-out (usually between 5 and 10 million cycles (Kühn 2000). The second test technique is akin to the tensile testing according to ISO 527. Again, the test is conducted with a sinusoidal cyclic loading under stress control until failure or run-out, whichever comes first (Harper & Bonfield 2000). The third test method is according to ASTM F2118, where the test specimens are subjected to fully reversed compressive and tensile loading in a sinusoidal way until failure or when the run-out limit is reached (Davies et al. 1987). Most fatigue tests are carried out at a particular frequency (for example, 2–5 Hz) in a buffered saline solution at 37°C. Testing is normally conducted at several stress levels to determine the fatigue behaviour over a range of stresses. The different testing methods result in different stress levels for failure under quasi-static conditions and under loading (Kühn et al. 2005). The trend is well documented and can be explained by the dissimilarities in the load applications and the consequential stress distributions in the cross-section of the test specimen. For tension–compression, the initial test results demonstrate a steeper reduction that may be caused by a stronger breakdown from the additional compressive loading. The materials perform in a similar way under bending and uniaxial tension. The most basic test configuration is the standardised four-point bending test in accordance with ISO 5833. Furthermore, the preparation of the test specimens for the tension–compression test is more complicated than the preparation technique for the four-point bending test samples. Therefore, the four-point bending technique is the favoured method for fatigue testing. The environmental conditions in which the tests are carried out have a significant effect on the fatigue performance. PMMA bone cements demonstrate significantly different fatigue properties when tested in dry or in wet conditions (Kühn et al. 2005). Fatigue tests conducted in air at room temperature result in a stress versus number of cycles to failure curve (S–N curve) of a noticeably greater slope. In order to simulate the physiological environment, fatigue tests should be conducted in a suitable aqueous solution, such as simulated body fluid or Ringer solution. The results of fatigue tests that are carried out in air at room temperature should be graded carefully (Freitag & Cannon 1977, Johnson et al. 1989). Additionally, the method of sterilisation of the PMMA polymer powder has a strong influence on the fatigue properties of the cured bone cement.

Mechanical properties of bone cements

249

Sterilisation by γ-sterilisation or β-sterilisation significantly lowers the molecular weight of the PMMA polymer powder and the resulting cured bone cement, while sterilisation by ethylene oxide (EtO) has no influence on the molecular weight of the PMMA powder. Bone cements with high molecular weight exhibit better fatigue properties than cements with low molecular weights (Tepic & Soltesz 1996, Harper et al. 1997, Lewis & Mladsi 1998, Lewis 2000). Porosity is a key reason for reduced fatigue performance of PMMA bone cement (Dunne et al. 2003). Pores or other inclusions act as stress concentrations in the bone cement and often initiate fatigue cracks within the microstructure of the bone cement. The cracks will ultimately result in failure of the cemented implant. The primary sources of porosity are air initially surrounding the polymer powder that is trapped during wetting of the powder, air trapped in the bone cement during mixing and air trapped in the cement during transfer from the mixing vessel to the application device. Mixing bone cement under atmospheric conditions in an open bowl results in a significantly higher number of pores than mixing cement in a vacuum mixing system (Wang et al. 1993, 1996, Dunne & Orr 2001). Modern cement mixing systems reduce cement porosity and improve bone cement strength by removing the likelihood of air entrapment in the cement (Wixson et al. 1987, Lewis 1999b, 1999c, Eveleigh et al. 2002, Dunne et al. 2003). Another method for the characterisation of the fatigue behaviour of PMMA bone cements is the fatigue crack propagation test (Nguyen et al. 1997). The test studies the fatigue crack propagation rate in compacttension bone cement specimens. During testing, the crack length and the number of fatigue cycles are measured. The fatigue crack propagation rate and the matching stress intensity factor range are subsequently determined.

11.5

Factors affecting the microstructure–mechanical properties relationship

The microstructure of PMMA bone cements can be understood as multiphase materials consisting of PMMA beads, polymerised MMA monomer and radiopacifying particles. Furthermore, the following factors will all influence the bulk of the interfacial microstructure of the cements and therefore their mechanical performance: the mixing method; the presence of blood or body fluids; the laminations produced when injecting bone cement into the bone cavity; the interface with existing cement, with the implant component, or with the cancellous bone; the thickness of the cement mantle; and the water sorption.

250

Orthopaedic bone cements

11.5.1 Porosity The debate about the real impact of porosity on the mechanical behaviour continues; however, is obvious that it is always suitable to decrease the porosity of PMMA bone cements. As a result, the improvement in cement mixing methods has undergone a major development. When bone cement was first used in arthroplasty, it was hand mixed in a bowl in the operating room and then inserted by hand or transferred and injected into the desired location. Because PMMA comes as a powder composed of pre-polymerised particles to be mixed with the liquid monomer, monomer fumes are released into the air. Furthermore, with hand mixing, a certain amount of porosity in the final material is unavoidable due to air entrapment, even in lower viscosity cements. During the 1980s different techniques were introduced in the hope of improving mixing and thereby bone cement properties (Burke et al. 1984, Lindén 1991). The results, however, were not convincing. Lidgren et al. (1984) introduced vacuum mixing of bone cement; the quality of the bone cement was improved. Today, vacuum mixing is widely accepted as the method of choice for achieving homogenous cement, reducing porosity and increasing cement strength, which is why it is an integral part of modern cementing techniques (Malchau & Herberts 1996). Vacuum mixing systems reduce the exposure of the operating theatre staff to monomer by 50–70% (Schlegel et al. 2004) and eliminate contact with bone cement during delivery (Darre et al. 1988, Buchhorn et al. 1992, Bettencourt et al. 2001, Eveleigh et al. 2002). The working environment for the theatre staff is improved, and the risk of fume-induced headaches, respiratory irritation and allergic reactions becomes minimal. Conventional mixing of bone cement produces porosity of 5–16% (Dunne & Orr 2001). Vacuum mixing produces porosity of 0.1–1% (Lindén & Gillquist 1989, Wang & Kjellson 2001). Porosity has been found to be the major cause of decreased mechanical performance of bone cement. To ensure its in vivo survival, the cement must be able to withstand the varying loads it endures. Thus the fatigue property, which is directly affected by porosity, is as important in determining the long-term survival of a joint replacement as static strength. Fatigue failure occurs when cement cracks are initiated as a result of defects in the cement mantle. It is known that vacuum mixing of cement increases mechanical properties (Lidgren et al. 1984, Alkire et al. 1987, Wixson et al. 1987, Schreurs et al. 1988, Lindén 1989, Askew et al. 1990, Davies & Harris 1990, Mau et al. 2004) largely as a result of decreasing micro- and macropores (Wang et al. 1993, 1996). Numerous studies have confirmed that vacuum mixing enhances the fatigue life of the bone cement (Lewis 1997, Harper & Bonfield 2000, Wilkinson et al. 2000, Dunne & Orr 2001, Schelling &

Mechanical properties of bone cements

251

11.13 Void on a fracture surface. Many partially unpolymerised PMMA particles and zirconium dioxide particles are seen in the voids (Wang et al. 1994) Arrows indicate free unbonded cement particles.

Breusch 2001, Yau et al. 2001, Murphy & Prendergast 2002, Dunne et al. 2003). Incomplete mixing of the monomer and polymer may lead to partially united and, in some cases, free unbonded cement particles (Figure 11.13). Vacuum mixing of bone cement not only decreases the number of voids, but also improves the microscopic homogeneity of bone cement (Wang et al. 1994). When cement fracture occurs, inhomogeneous cement may release PMMA and contrast media particles to the bone–cement interface. These particles may evoke a foreign body response or stimulate osteoclast activity (Sabokbar et al. 1997, 2001, Wimhurst et al. 2001), leading to osteolysis of the surrounding bone. Extensive porosity at the cement–stem interface has been found in retrieved cement mantles and in laboratory-prepared specimens (James et al. 1993, Bishop et al. 1996). This interface porosity is caused by entrapment of air at the stem surface during stem insertion and by residual porosity in the cement. When cement is mixed under vacuum, cement porosity is significantly reduced, thus producing less porosity at the cement–prosthesis interface (Bishop et al. 1996, Wang et al. 1998) (Fig. 11.14). Various studies have shown that interface porosity weakens the resistance of the cement to torsional load (Davies & Harris 1995) and decreases the fatigue life of the cement–metal interface (Iesaka et al. 2003). Interface porosity has also been linked to the initiation of cement cracks (Jasty et al. 1991, James et al. 1993, Verdonschot 1995). The evidence is convincing that reduction of interface porosity improves the strength of the interface, thereby increasing the longevity of cemented implants. The variation of cement porosity from mixing systems is still large (Dunne et al. 2004, Mau et al. 2004, Wang 2005). Various studies indicate

252

Orthopaedic bone cements

(a)

(b)

BC

M M

2 mm

BC

2 mm

11.14 Samples from a cemented implant. The cement was mixed at atmospheric pressure (a), and under vacuum (b). M, metal; BC, bone cement (Wang 2005).

that macropores increase the risk of fatigue failure, and the current opinion is that efforts should be made to minimise the number and size of macropores. The development and use of a pre-packed bone cement mixing and powerful delivery system to further minimise PMMA exposure, reduce porosity and make handling easier seems warranted.

11.5.2 In vivo environment The bone cement has to perform its function in the human body, therefore the influence of body fluids and body temperature (c. 37°C) on its properties is important. Uptake of liquids has been reported to affect some of the mechanical properties of the cement post-implantation. Sorption of water generally reduces the mechanical properties, resulting in a lower modulus of elasticity and in a less stiff material (Bargar et al. 1986, Looney & Park 1986). The decreasing stiffness may even be beneficial for fracture resistance and long-term stability of the implant component. However, fracture mechanics studies show that the crack velocity is less in water than in air, and that fracture toughness is c. 15–20% greater in water than in air (Beaumount & Young 1975, Buckley et al. 2003). Other results demonstrate that the work of fracture increases with the storage of the bone cement in liquids (for example water, Ringer solution and lipids), although in each case the work of fracture after storage at 22°C is greater than that after storage at 37°C (Hailey et al. 1994). Water in PMMA bone cements acts as a plasticiser. The water sorption of commercial plain bone cements is c. 1– 2% and is slightly higher for antibiotic-loaded cements. Kühn et al. (2005) reported a high initial water uptake during the first few days of incubation at 37°C and water saturation was reached after 4–8 weeks’ incubation.

Mechanical properties of bone cements (a)

253

(b)

11.15 Scanning electron micrographs of PMMA bone cement containing radiopacifier. (a) Palacos® R: 1, polymer bead; 2, ZrO 2. (b) Simplex P: 1, polymer bead; 2, BaSO4.

11.5.3 Additives Radiopacifier particles PMMA bone cement is not a radiopaque material, therefore it is impossible to determine the boundaries of bone cement applied during surgery by ordinary X-ray imaging techniques. Until 1972, PMMA cement did not contain any radiopaque materials and was, therefore, radiolucent. It is important that the orthopaedic surgeon can easily monitor and evaluate the healing and loosening processes to differentiate between bone, bone cement and osteolysis after a joint replacement, so well-defined opacity is required. For this reason, radiopaque media are added to PMMA bone cements. Barium sulphate (BaSO4) or zirconium dioxide (ZrO2) are used as radiopacifiers in all commercially available bone cements (Fig. 11.15). These radiopacifiers are not a part of the polymer chain. They are dispersed uniformly in the polymer powder and in the resulting hardened bone cement. All PMMA bone cements contain 8–15% X-ray contrasting agent within the polymer powder (Kühn 2000). Compared with bone cements using BaSO4, those containing ZrO2 have a significantly higher opacity. Bone cements with 15% ZrO2 in the powder component have the highest opacity. The addition of radiopacifiers to bone cement has many disadvantages. Animal experiments with different cell cultures showed greater differences in bone resorption around the bone cement application area when using BaSO4, in comparison to ZrO2 (Sabokbar et al. 2001). Notwithstanding the low solubility of BaSO4, toxic barium ions can be released. In comparison, the abrasive properties of ZrO2 appear to be their primary disadvantage. The disadvantage is apparent only in cases of loosening of the prosthesis, or if free cement particles migrate to the joint articulation. The radiopaque media remain in the body for decades as components of

254

Orthopaedic bone cements

PMMA bone cement. When the static and dynamic mechanical properties of PMMA bone cement that included BaSO4 and ZrO2 were examined, it was observed that the radiopacifying agents have a significant influence on the mechanical properties of PMMA cement depending on their size and morphology. It has been reported that, although the addition of BaSO4 produces a decrease in tensile strength of c. 10% (Haas et al. 1975, Kusy 1978, Ginebra et al. 1999), the addition of ZrO2 does not influence this property. However, it has been shown that the fracture toughness, which is unchanged by the addition of BaSO4, is increased c. 20% by the inclusion of ZrO2 (Ginebra et al. 2000a). Furthermore, it has been reported that both inorganic radiopacifying agents improve the fatigue crack propagation resistance of PMMA bone cement (Molino & Topoleski 1996, Ginebra et al. 2000a, 2000b). This behaviour can be explained in terms of the microstructure of the radiopacifier agents and their interaction with the polymer matrix (Fig. 11.15). As there is no chemical reaction between the inorganic particles and the PMMA bone cement, it can be understood that the filler particles act like pores when a tensile stress is applied to the bone cement. Nevertheless, the morphology of the particles plays an important function with respect to the mechanical properties of the bone cement. In fact, the ZrO2 particles, which have a cauliflower-like shape, can secure mechanically to the polymer matrix, reinforcing it to a certain degree. This situation does not occur in the case of the smaller and regular-shaped BaSO4 particles. Presently, some iodine-containing monomers are being investigated as potential substitutes for the inorganic radiopacifying agents with improved biological and mechanical characteristics (Vázquez et al. 1999, Ginebra et al. 2000a, 2000b, van Hooy-Corstjens et al. 2004). Antibiotics Surgical operating theatres have sterile requirements, but even under these constraints some bacteria can pass through all of the protective barriers and contaminate the open body tissues during surgery. Since the early 1970s, the use of antibiotics in bone cement in both primary and aseptic revision joint replacement surgery has been widespread in an attempt to prevent postoperative infections and biofilm formation (Buchholz et al. 1984). The antibiotics are released from the bone cement into the tissues surrounding the total joint, this local concentration of antibiotics is sufficient to kill the antibiotic-sensitive bacteria left in the operative wound (Walenkamp & Murray 2001, Hendriks 2003, Breusch & Malchau 2005). Gentamicin sulphate is used in many commercial brands of PMMA bone cements worldwide. Apart from gentamicin sulphate, other antibiotics have also been used effectively in bone cement, such as tobramycin, vancomycin,

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clindamycin and fusidic acid. Combinations of these antibiotics, for example, clindamycin and gentamicin in bone cement, have been used for revision surgery (Breusch & Malchau 2005, Walenkamp & Murray 2001, Penner et al. 1996, Koo et al. 2001, Konig et al. 2001). It has been reported that the presence of small doses of antibiotic (less than 1 g per 40 g pouch of PMMA polymer powder) in bone cement has no detrimental influence on the mechanical behaviour of the cured PMMA cement, and does not change the thermal or viscosity characteristics (Marks et al. 1976, Murray 1984) but the contrary has also been reported (Lautenschlager et al. 1976, Davies et al. 1988, Klekamp et al. 1999, Persson et al. 2006). Dunne et al. (2007) showed that large doses of antibiotics (greater than 1 g per 40 g pouch of polymer powder) decreased the compressive, bending, tensile and fatigue properties of bone cements. Moreover, He et al. (2002) observed that the addition of large doses of antibiotic to PMMA cement increased the amount of unreacted MMA liquid monomer, thereby implying that antibiotic accelerated the polymerisation reaction. If surgeons hand mix antibiotics into bone cement at the time of surgery, the quality of the antibiotic-loaded cement may be affected (Deluise & Scott 2004, Lewis et al. 2005). The disadvantages of this approach are poor mechanical properties and a deleterious effect on elution kinetics. Fibres PMMA bone cement was first used for orthopaedic implant fixation in total joint prostheses in the 1960s and since then it has been adopted to support knee, shoulder, elbow and other prostheses. Nevertheless, in the first applications many prostheses were revised because fractures of the surrounding PMMA cement were significant (Charnley 1960). Therefore, many investigations have been conducted to improve the mechanical performance of PMMA bone cement. Early work tried to improve the mechanical properties by additives, and different mechanical tests were carried out – for example, tensile, compressive, three-point bending, fracture toughness and fatigue tests. One of the methods of improving the performance of the bone cement involves dispersing small quantities (typically 1–2 vol%) of reinforcing materials such as carbon (Pal & Saha 1982, Friss et al. 1996), graphite (Knoell et al. 1975, Saha & Pal 1984), aramid (Pourdeyhimi et al. 1986), bone particle (Park et al. 1986, Liu et al. 1987), polyethylene (Wagner & Cohn 1994; Friss et al. 1995), titanium (Topoleski et al. 1992), ultra-high molecular weight polyethylene (Pourdeyhimi & Wagner 1989), or PMMA (Buckley et al. 1992, Gilbert et al. 1994) fibres in the bone cement matrix. Although most of the results from the mechanical tests have been encouraging, the biocompatibility issues relating to some of these fibres are not yet known

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(Wright & Trent 1979). Therefore, none of these reinforced cements have been approved for clinical application.

11.6

Modification of acrylic bone cements

Currently, a vast range of possible modifications of conventional PMMA bone cements are being investigated, focusing on the improvement of either the mechanical performance or biological properties. Researchers have attempted to enhance the mechanical properties of conventional bone cements by way of improving the preparation protocols for cement mixing, such as porosity reduction, or by improving surgical techniques, for example, reaming and cleaning of the bone cavity or pressurisation during cement delivery and implantation of the prosthesis. Moreover, augmentation of the mechanical properties of PMMA bone cement can be achieved through the reinforcement of the cement matrix either using particles or with fibres (Deb 1999, 2006). With regard to particle addition, three main directions can be illustrated: (a) reinforcement with hard particles such as glass beads (Beaumount 1977) or glass ceramic particles (Henning et al. 1979); (b) reinforcement with tough or rubber-toughened particles (Murakami et al. 1988, Vila et al. 1999a, 1999b); and (c) reinforcement with bioactive particles, for example, inorganic bone and demineralised bone matrix (Liu et al. 1987) and hydroxyapatite (Low et al. 1993), which could improve both the mechanical performance of the cement and the strength at the bone–cement interface. Fibre reinforcement of bone cements can be described as reinforcement with metallic fibres (Topoleski et al. 1992) and reinforcement with polymeric fibres, including carbon fibres (Saha & Pal 1984, Pourdeyhimi & Wagner 1989, Buckley et al. 1991, 1992). Alternative chemical modifications of the bone cement matrix – for example, substitution of initiators, accelerators (Brauer et al. 1986, Vázquez et al. 1997) or radiopacifying agents (Ginebra et al. 1999, 2000a, 2000b, Vázquez et al. 1999, van Hooy-Corstjens et al. 2004) with more biocompatible compounds, or the addition of other monomers to the liquid phase (Brauer et al. 1986, Oysaed 1990, Davies & Harris 1992, Pascual et al. 1999) – improve other physical, chemical and biological properties.

11.7

Summary

As stable fixation of implant prostheses is the principal application of PMMA bone cements, their mechanical properties are of noteworthy importance. In the current standard for PMMA bone cements, two static tests are recommended. These two tests are not adequate for an accurate mechanical appraisal of PMMA bone cements. Many other mechanical test techniques are presented in the literature reporting the mechanical behav-

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iour of PMMA bone cements. The simulation of the in vivo scenario is very complex, as the mechanism of loading in the clinical situation is difficult, especially for a total hip replacement. Obviously, the mechanical property data are influenced by factors such as bone cement type, mixing method, test specimen geometry, test technique, temperature and test environment. One of the most important test techniques is the fatigue test. At present, there are three well-known test methods with differing considerations. As all of them provide similar results, the four-point bending method should be the test of choice. All PMMA bone cements that are commercially available satisfy the minimum mechanical requirements as stipulated by the standards, but there are important variations in their mechanical behaviour. The orthopaedic surgeon, therefore, has to select the optimum bone cement to achieve the best outcome for the patient.

11.8

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verdonschot n, huiskes r (2000), ‘Creep properties of three low temperaturecuring bone cements: a preclinical assessment’, J Biomed Mater Res, 53, 498–504. vila m, ginebra m, gil, planell j (1999a), ‘Effect of porosity and environment on the mechanical properties of acrylic bone cement modified with acrylonitrilebutadiene-styrene particles: Fracture toughness’, J Biomed Mater Res, 48, 121–7. vila m, ginebra m, gil, planell j (1999b), ‘Effect of porosity and environment on the mechanical properties of acrylic bone cement modified with acrylonitrilebutadiene-styrene particles: Fatigue crack propagation’, J Biomed Mater Res, 48, 128–34. wagner h, cohn d (1994), ‘Use of high performance polyethylene fibers as a reinforcing phase in poly(-methylmethacrylate) bone cement’, Biomaterials, 16, 139–41. walenkamp g, murray d (Eds) (2001), Bone Cement and Cementing Technique, Berlin, Heidelberg, New York, Springer-Verlag. wang js (2005), ‘Mixing: The benefit of vacuum mixing’, In The Well-Cemented Total Hip Arthroplasty. Theory and Practice, Eds S Breusch, H Malchau, Berlin, Heidelberg, New York, Springer-Verlag, pp. 107–12. wang js, franzén h, jonsson e, lidgren l (1993), ‘Porosity of bone cement reduced by mixing and collecting under vacuum’, Acta Orthop Scand, 64, 143–6. wang js, goodman s, franzén h, aspenberg p, lidgren l (1994), ‘The effects of vacuum mixing on the microscopic homogenicity of bone cement’, Eur J Exper Musculoskeletal Res, 2, 159–65. wang js, toksvig-larsen s, müller-wille p, franzén h (1996), ‘Is there any difference between vacuum mixing systems in reducing bone cement porosity?’, J Biomed Mater Res (Appl Biomater), 33, 115–9. wang js, aspenberg p, goodman s, lidgren l (1998), ‘Interface porosity in cemented implants in vitro study’, In Proceedings of the 8th European Research Orthopedics Society Meeting, p. 2. wang js, kjellson f (2001), ‘Bone cement porosity in vacuum mixing system’, In Bone Cements and Cementing Technique, Eds G Walenkamp, D Murray, Berlin, Heidelberg, New York, Tokyo, Springer, pp. 81–95. wang js, taylor m, flivik g and lidgren l (2003), ‘Factors affecting the static shear strength of the prosthetic stem–bone cement interface’, J Mater Sci Med, 53, 55–61. wilkinson j, eveleigh r, hamer a, milne a, miles a, stockely i (2000), ‘Effect of mixing technique on the properties of acrylic bone cement’, J Arthroplasty, 15, 663–7. wimhurst j, brooks r, rushton n (2001), ‘The effects of particulate bone cements at the bone–implant interface’, J Bone Joint Surg, 83-B, 88–92. wixson r, lautenschlager e, novak m (1987), ‘Vacuum mixing of acrylic bone cement’, J Arthroplasty, 2, 141–9. wright t, trent p (1979), ‘Mechanical properties of aramid fiber reinforced acrylic bone cement’, J Mater Sci, 14, 502–5. yau wp, ng tp, chiu ky, poon kc, ho wy, luk dk (2001), ‘The performance of three vacuum mixing cement guns- a comparison of the fatigue properties of Simplex P cement’, Int Orthop, 25, 290–3.

12 Fracture toughness and fatigue characteristics of bone cements A. B. L E N N O N, Trinity College Dublin, Ireland

Abstract: This chapter describes fundamental characteristics of bone cement fatigue and fracture and some of the factors affecting them. It begins with bulk cement behaviour and later moves to factors that are relevant to fatigue and fracture of cement as used in many clinical interventions. Finally, some properties for fracture toughness of two commercial cements are reviewed and tabulated from the literature for three common mixing methods. Key words: acrylic bone cement, poly(methylmethacrylate) (PMMA), fatigue, fracture, fracture toughness.

12.1

Introduction

Acrylic bone cement is used widely in orthopaedics and is a dominant fixation material in joint arthroplasty in particular. Since bone cement is a form of poly(methylmethacrylate) (PMMA), it might be expected to have similar fracture and fatigue properties to industrial Perspex/Plexiglas. However, due to the requirement to fill and conform to the bone cavity and prosthesis, it is prepared during surgery as a self-curing resin. This polymerisation process is very sensitive to environmental conditions and this can cause considerable variability in the mechanical properties of the final cement and thus in the longevity of the fixation. Air entrapment, in particular, can considerably reduce the fatigue life of the final cement by introducing stress-concentrating pores that act as sites of crack initiation. Bone cement is thus considerably weaker than the other components typically used in joint replacement, e.g. Co–Cr or ultra-high molecular weight polyethylene (UHMWPE). Because of this, along with its widespread use, bone cement has been the focus of much attention in failure analyses of joint replacement.1–3

12.2

Factors affecting fracture toughness and fatigue resistance of bone cement

Fracture toughness and fatigue resistance are notoriously variable phenomena, especially fatigue. This variability can arise from many sources and, in 265

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the case of bone cement, as used in the clinical setting, is further complicated by the need to prepare the material intra-operatively. This tends to produce compositional variability and stochastic porosity distributions. Brandspecific additives can cause both crack initiation and arrest. In addition to these compositional and structural factors, the type of loading can also influence fracture behaviour, e.g. static versus fatigue loading, crack opening versus sliding versus torsional loads, and uniaxial versus multiaxial loading. The following sections will discuss each of these factors in more detail to give a better understanding of bone cement fracture behaviour. Fractography, i.e. visual analysis of fracture surfaces, will form the basis for much of the discussion to highlight the microstructural features relevant to a particular fracture behaviour.

12.3

The effect of loading mode on fracture and fatigue

Bone cement is typically subjected to complex time-varying loads that are repeated for many millions of cycles. Some activities can induce one-off loads of high magnitude, e.g. due to stumbling, while other loads may be less severe but repeated for millions of cycles during the service life, e.g. walking. In addition to this, the loading complexity can also vary considerably due to the type of device or application. For example, consider the high degree of bending and wedging action occurring in the cement surrounding a femoral prosthesis during walking versus the pressurisation of the cement layer of an acetabular component or the cement body of a vertebroplasty. Understanding how bone cement responds and fails due to these different loading modes is therefore critical before deciding if it is suitable for use in a given clinical application.

12.3.1 Monotonic loading Typical static bench tests subject simple bone cement specimens to uniaxial loads in either tension or compression. Static tensile testing is particularly useful for classifying the characteristic fracture behaviour of a material, typically specified as either brittle (low toughness) or ductile (greater toughness due to plastic deformation) depending on the type of fracture surface observed and the amount of energy required to fracture the specimen.4 In an early low magnification study Kusy1 divided the fracture surface of statically tested bone cement specimens into three regions: (a) mirror – a reflective, comparatively featureless, region near the initiation site; (b) mist – a rougher region characterised by conical markings;

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Bulk polymer matrix Conical void profile Fracture path

Craze zone (A) Mirror

(B) Mist

(C) Hackle

12.1 Schematic of crack propagation in a glassy polymer. Region A is relatively featureless as the crack propagates mainly in the mid-plane of the craze zone. Region B becomes rougher as the crack begins to accelerate and alternates between craze bundles leaving patches of mating craze bundles (mist). In region C the crack tip has outpaced the main crazed region and secondary craze bundles form ahead of the tip. After propagating through a given bundle, a new bundle is formed and the process repeats, leading to a banded structure (hackles).

(c) hackle – similar to the mist region but modifying the conical markings by the generation of ribs. This type of fracture surface is indicative of a predominantly brittle behaviour, typical of glassy polymers. Hertzberg4 describes the mirror phase as the progression of fracture through a pre-crazed region by the nucleation and growth of conical voids along the mid-plane of the crazed region (region A of Fig. 12.1). Additionally, the mirror region also exhibits coloured patterns caused by the different refractive index of the layer of craze matter on the fracture surface. As the crack velocity increases, fracture tends to progress alternately from one craze-matrix boundary to another, forming mating islands of craze matter on each side of the fracture surface (region B of Fig. 12.1). During the terminal phase of fast fracture, the crack front outpaces the tip of the craze region. This causes secondary craze formation at the crack tip and crack opening begins to occur through bundles of these secondary crazes. The crack propagates through one bundle of crazes until a new bundle is formed and the process is repeated (region C of Fig. 12.1). Recent studies, e.g. by Vallo et al.5 and Liu et al.6, have confirmed these features for clinical preparations of bone cement, in addition to stresswhitening, arising from the crazing prior to crack propagation.

12.3.2 Fatigue/cyclic loading Fatigue failure differs from monotonic fracture mainly in the early crack propagation phase and is more likely to exhibit multiple crack nucleation sites. Crazing, as for the monotonic case, is the dominant mechanism of

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Crack tip

with each cycle

Crack tip

⎫ ⎬ ⎭

⎫ ⎬ ⎭

extends

Striations (high ΔK)

Craze zone

⎫ ⎬ ⎭

⎫ ⎬ ⎭

268

⎫ ⎬ ⎭

Craze zone grows continuously Crack extends as craze reaches critical length

Hackle region

River/ratchet line

Beach markings

Discontinuous growth bands (low ΔK)

Crack initiation site

12.2 Fatigue markings and associated crack propagation processes from microscopic to macroscopic length scales. Severed PMMA bead

River lines Discontinuous crack growth bands

100 μm

12.3 Crack initiation site in a hand-mixed specimen. Several features typical of fatigue fracture are demonstrated: discontinuous growth bands, river lines and severed PMMA beads. Courtesy of Dr B. P. Murphy, National University of Galway.167

crack propagation.7,8 Two sets of microscopic striations can sometimes be observed in both industrial4 as well as clinical9 preparations of PMMA. A first set corresponds to incremental crack advance during one load cycle that tends to occur for relatively high crack-tip stress ranges.4,9 A second, more widely spaced, set of striations are caused by repeating periods of continuous craze zone growth – each period corresponds to growth up to a critical length, at which point the crack advances through the crazed region and arrests again. This process results in periodic markings, sometimes referred to as discontinuous growth bands4 (a schematic describing the various fatigue crack propagation phases is presented in Fig. 12.2). At the macroscopic level, discontinuous crack bands, or beach markings, may occur as a result of altering crack propagation characteristics in different amplitude loading blocks (Figs 12.2 to 12.4). Finally, lines emanating radially

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(a)

2 mm (b)

500 μm

12.4 Fracture surface of a vacuum-mixed uniaxial fatigue specimen. (a) Crack nucleation site (box) surrounded by pre-critical fracture zone and remaining hackled fast fracture region. (b) Close-up of nucleation site showing river lines originating from a pore. Images courtesy of Dr B. P. Murphy.167

from the crack nucleation site are frequently observed, often referred to as ‘ratchet’ or ‘river’ lines.4,10 These lines represent the junction of adjacent crack fronts that may propagate out from a defect, e.g. a pore, and disappear when these crack fronts link together (see Figs 12.2 to 12.4). All these features occur in the region of crack propagation prior to fast fracture. Once the critical crack length is achieved for a given loading condition, fast fracture will occur. Thus a hackle region is usually observed beyond the initial fatigue crack (Fig. 12.4(a)). Another difference that tends to occur with fatigue of bone cement is a tendency for multiple cracks to initiate and propagate in a specimen before

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one becomes critical. In an early fatigue study, Freitag and Cannon11 observed textured and rough fracture surfaces at lower stresses until a crack of sufficient length was created to initiate fast fracture. Small cracks were also observed at locations away from the failure crack. At higher stresses, smooth and planar fracture surfaces were observed with little cracking occurring away from the main crack. Murphy and Prendergast12 also observed multiple crack initiation in fatigue specimens but found that greater crack initiation occurred for higher stresses. Studies of ex vivo retrieved cement mantles have also found multiple cracks and, along with fractographic analysis, demonstrated that fatigue failure is the dominant failure mode of bone cement in vivo.13–15 Finally, load sequence, or variable-amplitude loading, is an event that can affect the time to failure for many materials. Palmgren–Miner’s rule,16,17 i.e. a linear sum of the life fractions for each block, is often applied in such cases. This assumes that, at failure, the life fractions of all loading blocks will equal unity (i.e. Σni/Nfi = 1, where ni is the actual number of cycles for a given block ‘i’ and Nfi is the fatigue life for the stress level applied for that block). However, experimental observation frequently shows that this is not the case.18 For example, in metals, high–low sequences often result in summation to less than unity (i.e. n1/Nf1 + n2/Nf2 < 1) and low–high sequences result in summation to greater than unity (i.e. n1/Nf1 + n2/Nf2 > 1). Although this is well established in metal fatigue,19 it has received little attention in studies of bone cement fatigue. One of the first studies to address non-linear accumulation in bone cement was carried out by Murphy and Prendergast.20 They developed a non-linear fatigue damage accumulation model based on experimental measurement of microcracks in wasted bone cement specimens subjected to different stress levels. Damage was expressed as D = (n1/Nf1)α(s), where α is an exponent that is a function of applied stress, s. This stress dependence of the exponent results in nonlinear accumulation between blocks of different stress levels.19,21 Unlike metals, the stress-dependent exponent causes their model to predict that high–low sequences result in n1/Nf1 + n2/Nf2 > 1 and low–high sequences in n1/Nf1 + n2/Nf2 < 1. Unfortunately, they did not test this prediction. However, this type of behaviour is known to occur frequently in notched specimens4 so it is possible that the stress concentrations occurring from porosity may be causing a similar effect in bone cement. In 2007, Evans22 performed variable-amplitude testing of fatigue crack propagation in pure PMMA and Palacos R bone cement. This study showed that a variety of effects could be demonstrated under different overloading scenarios (i.e. low–high sequences): acceleration of fatigue crack growth rate for individual overloads, no effect for repeated overloads, and retardation of crack growth rate for spectrum overloading. Given the variety of effects that can occur and the potential to influence fatigue life, it seems that this remains a relatively

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unexplored area of bone cement failure analysis that merits further research.

12.3.3 Multiaxial loading Much of the fracture and fatigue data that exist for bone cement are for uniaxial tension and thus do not completely represent the more complex loading that tends to occur in clinical use. For example, in the cement mantle of a total hip replacement there can be a combination of stress states due to different loading modes, for example:23,24 • •

axial tension and compression due to bending loads; symmetric hoop and radial stresses due to distal displacement of the prosthesis; • asymmetric hoop and radial stresses due to varus/valgus rotation and/or anteversion/retroversion of the prosthesis; • torsional stresses due to rotation of the prosthesis about its long axis. Very few studies have investigated the effect of multiaxial loading on bone cement failure. A number of studies by Leevers et al.25–27 investigated both monotonic fracture and fatigue crack propagation of industrial PMMA under biaxial stress fields. They found that, under monotonic conditions, increasing the stress parallel to the crack with respect to the normal stress could cause crack path deviation with respect to the traditional uniaxial case (i.e. stress normal to crack).26 Furthermore, they observed a decrease in fatigue crack propagation rates with increases in the parallel stress with respect to the normal stress.27 They suggested that a crack closure mechanism due to the parallel stress caused this phenomenon. Silvestre et al.28 investigated fracture using pressurised cylindrical specimens subjected to axial loads, resulting in triaxial stresses, and found that failure could be predicted using Mohr–Coulomb theory. This cylindrical test specimen design was later adopted by Murphy and Prendergast29 to investigate fatigue failure by changing the monotonic axial load to a cyclic load. They found that the probability of survival of specimens under multiaxial loading was reduced compared with corresponding uniaxial specimens. Furthermore, the variability in fatigue life was also increased for multiaxial specimens. Retrieval studies have highlighted the relevance of analysing multiaxial failure. For example, Topoleski et al.13 observed orthogonal groups of cracks in some regions of the cement mantle, indicating that biaxial tension was acting in such regions. Jasty et al.14 also observed different types of cracks, e.g. radial and circumferential cracks near the cement–prosthesis interface.

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12.4

The effect of porosity on fracture and fatigue

The importance of the role of porosity in initiating fatigue failure of bone cement has been demonstrated in many fractographic studies.12–14,30 Pore production has several causes arising from the preparation of bone cement:1,31–34 (a) air entrapment during mixing; (b) heating of the cement can cause the liquid monomer to boil; (c) shrinkage during polymerisation and the initiation of this process at the warmer cement–bone interface, which causes bone cement to shrink away from the stem;35,36 (d) the rheological behaviour of the cement as it comes in contact with the inserted prosthesis – large numbers of pores have been observed forming at this interface during implant insertion and the phenomenon has been related to the shear rate experienced by the doughy cement as the prosthesis is inserted.37 Pores decrease fatigue life and fracture strength via increased loading due to loss of net cross-section that can resist applied loads, the geometrically induced stress concentration due to their approximately spherical shape, and stress intensity interaction effects when pores cluster in groups. Furthermore, in the already raised stress state occurring around a pore, a further stress concentration occurs between beads in the bead–matrix microstructure, making these primary sites for crack initiation from pores (Fig. 12.5).29,30 Clear evidence of the link between pores and fatigue damage accumulation has been reported in a time-lapse study of damage accumulation in uniaxial specimens,12,20 in which microcracks initiated almost exclusively from pores. In particular, there was a positive correlation of damage accumulation per pore per cycle with applied stress. This link has been confirmed using synchrotron X-ray microtomography by Sinnett-Jones et al.30 Since pores have obvious detrimental effects on fatigue and fracture resistance, there has been much emphasis on reducing porosity by decreasing air entrapment during the mixing process. This has been achieved mainly by centrifuging and vacuum mixing.33,38–40 A large body of fatigue and fracture literature on bone cement has focused on demonstrating differences between bone cements prepared using so-called first-, second-, and thirdgeneration mixing techniques.10,41–48 These studies have demonstrated improvements in both fatigue and fracture resistance for both centrifuged and vacuum-mixed bone cement test specimens. However, the fatigue lives for such specimens are highly variable – centrifuged specimens still contain small pores and are prone to density variation while vacuum-mixed samples are susceptible to occasional large pores. Consequently, there has been

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273

(b)

12.5 Cracks occurring around pores. (a) Example of crack initiation around a pore. (b) Stress concentrations between beads initiate cracks in the matrix phase. Scanning electron micrographs courtesy of Dr B. P. Murphy.29

some speculation as to how well laboratory studies of relatively simple test specimens translate to the more complex environment of clinical application. Davies and Harris49 speculated that vacuum mixing would not eliminate early failures, in spite of increases in average strength, because of the presence of such large pores. Controversially, Ling and Lee50 have suggested that clinical evidence does not support the goal of porosity reduction in hip replacement. On the other hand, results from the Swedish National Hip Arthroplasty Register clearly demonstrate that improvements in cementing technique have led to improved survival of cemented hip replacements.51 This controversy has also spilled over into the field of laboratory testing, specifically in the area of specimen selection. On the one hand, Prendergast et al.52 have argued that discarding specimens with large pores yields results with unrealistic variability, particularly for vacuum mixing which is susceptible to such porosity. On the other hand, Cristofolini et al.53,54 argued that discarding specimens with pores of greater than 1 mm diameter provides a rational approach to specimen selection that allows statistical comparison of the material resistance to fatigue between different cements. Both approaches have their merits, but one benefit of not discarding samples is the ability to identify problems, and emphasise their consequences, in the mixing technique. For example, by including all samples Dunne et al.55

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demonstrated that, for equivalent vacuum levels, the mixing method could influence the susceptibility to generate large pores, and hence increase the variability of vacuum-mixed samples.

12.5

The effect of inclusions on fracture and fatigue

There are two main sources of inclusions to be considered when analysing the fatigue and fracture of bone cement: (a) biological inclusions due to the the intra-operative and in vivo environment and (b) manufacturer’s additives. Biological inclusions, e.g. blood and fat, reduce mechanical strength. PMMA is also hydrophilic, absorbing up to several weight per cent water. Absorbed water acts as a plasticising agent and has been shown to increase fatigue life. Mechanical properties can therefore be expected to vary in clinical usage, especially considering the application environment of surgery.1,11,31,56,57 Manufacturer’s additives, such as radiopacifiers and antibiotic agents, that are not directly involved in the polymerisation of the cement can act as both crack initiation sites or, conversely, crack arresters. For instance, Ginebra et al.58 found toughening mechanisms for inorganic radiopacifiers barium sulphate and zirconium dioxide that were dependent on the size and morphology of the particles. In particular, barium sulphate tended to form smaller agglomerated inclusions while zirconium dioxide had a larger cauliflower-like morphology that enabled better mechanical interlock with the cement matrix; a similar result was also observed by Harper and Bonfield.59 The effect of particle size may explain some conflicting results in the literature for barium sulphate. For example, some authors have found increases in fatigue strength for addition of barium sulphate,11 while others have measured a decrease.60 However, the general trend would appear to be that addition of barium sulphate has a positive effect on fatigue life.61 A general trend of no significant effect appears to apply for antibiotic additions, such as gentamicin and vancomycin, with most showing no change or no significant reduction in fatigue life60,62–65 but a small number of studies finding reductions in fatigue life for some combinations of cement brand, mixing system and antibiotic.66,67 Some experimental inclusions have also been used in the past, most notably fibre reinforcement. For example, Robinson et al.68 compared the fracture surfaces of graphite fibre-reinforced low-viscosity cements and observed toughening mechanisms in the form of fibre pull-out. Other reinforcement methods have included metal meshes69 and fibres70 and even PMMA fibres.71 Although increases for in vitro strength have been demonstrated in many studies, reinforcement has not achieved clinical use, primarily because of outstanding issues with (a) interfacial adhesion between fibre

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and matrix, (b) rheological behaviour during the handling phase, (c) high stiffness of the reinforced cement with the potential to cause excessive load transfer to the cement–bone interface, and (d) biocompatibility.72

12.6

The effect of cement chemistry on fracture and fatigue

The distribution of the different fracture regions observed in fractographic analyses can vary depending on the cement formulation so that commercial cements frequently exhibit a variety of fracture behaviours. Although some of these differences can often be attributed to the radiopacifier or antibiotic additives, as discussed above, they can also stem from the differences in the proportions and/or presence of constituents that take part in the self-curing polymerisation reaction. Relative molecular weight is known to be a major factor in determining polymer mechanical properties. This is mainly due to the fact that polymers with higher molecular weights tend to have longer chains – these tend to become more entangled with each other, which leads to increased stiffness and fracture and fatigue resistance.4,57 Several factors can influence the final molecular weight distribution of the cement, for example: • • •

the molecular weight of the pre-polymerised powder; the conversion ratio of the reaction of the monomer to the polymer; the method of sterilisation of the powder.

Several studies have investigated the effect of molecular weight on fracture and fatigue strength, in particular the effect of powder sterilisation.73–76 All of these studies have demonstrated a tendency for reduced fatigue strength with reduction in molecular weight. Differences in cured molecular weight between the pre-polymerised powder beads and inter-bead matrix have been proposed as one of the reasons for a tendency for cracks to initiate and propagate more in the inter-bead matrix during early crack propagation.77 However, this effect may be counteracted by crack deflections at matrix–bead interfaces, crack arrest at or within beads and crack-tip shielding.30 Less research has been conducted on the influence of the initiator and activator components on fracture and fatigue properties. Hasenwinkel et al.78 found that altering the concentration of initiator (benzoyl peroxide) and activator (N,N-dimethyl-p-toluidine or DMT) in a novel two-solution bone cement caused differences in the fracture surface but did not find a significant difference for fracture toughness or fatigue life compared with commercial liquid–powder cements. Lewis et al.79 also found no significant difference in fatigue life or fracture toughness for different ratios of initiator and activator 5 years later. Alternatives to DMT have also been

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investigated in an effort to reduce cytotoxicity and have been shown to retain and even surpass fatigue and fracture properties of DMT-containing formulations.80 Cross-linking agents have also been used to improve mechanical properties.72,81,82 Many of these have been shown to retain rheological properties during the dough phase while increasing tensile properties.81,82 However, at least one cross-linking agent (bis-GMA, an adduct of bis-phenol A and glycidyl(methacrylate)) has been found to embrittle the polymer matrix at higher concentrations83 indicating that care should be taken in optimising the concentration of any proposed cross-linking agent. Perhaps the best known case of bone cement chemistry influencing arthroplasty revision risk has been that of Boneloc bone cement, which was withdrawn in 1995 only 4 years after worldwide release in 1991.84 Boneloc was designed to have low exotherm, low release of monomer and aromatic amines, sufficient mechanical properties, and to meet or exceed ISO and ASTM standards. Further aims were to provide a delivery system that was convenient and enabled low porosity mixing and delivery in the operating environment, while minimising personnel exposure to monomer vapours.85 Boneloc differed from conventional bone cements in both the liquid monomer and powdered PMMA constituents. Half of the methylmethacrylate (MMA) in the monomer was substituted with long-chain, high molecular weight, less volatile and less soluble methacrylates (ndecylmethacrylate, isobornyl-methacrylate), as well as alteration of the accelerator system to a mix of dihydroxypropyl-p-toluidine and DMT. The powder contained butylmethacrylate–MMA copolymers. These alterations lowered the glass-transition temperature, and were intended to enable complete mixing in the customised mixing and delivery system.85 By 1995 the Norwegian Arthroplasty Register had documented unacceptable revision rates in both femoral and acetabular sides of hip replacements86 and the Norwegian Orthopaedic Association recommended its withdrawal later that year.87 One of the first mechanical studies reported was by Kindt-Larsen et al.,85 who concluded that Boneloc had mechanical properties within the range of other conventional bone cements. However, closer examination indicates trends of reduced tensile properties, i.e. mean fracture toughness of Boneloc was between 57 and 59% of the vacuummixed values of the other conventional cements, while mean tensile strength was between 68 and 80% of the other cements tested. Later that year in Sweden, Thanner et al.88 demonstrated reduced mechanical properties, in particular tensile strength, for Boneloc specimens compared with Palacos. They also measured increased femoral prosthesis migrations for cases fixated using Boneloc using radiostereographic analysis. The cement was subsequently disbanded in Sweden, withdrawn by the distributor in Finland, and in April 1995 the manufacturer withdrew the cement from worldwide distribution.84

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Since then, numerous clinical follow-up studies have highlighted poor outcomes of arthroplasties fixated with Boneloc.89–101 However, relatively little has been published on the mechanical properties since the early studies of Kindt-Larsen et al.85 and Thanner et al.88 In 2000, Harper and Bonfield59 performed a study of ten commercial cements and paid particular attention to tensile properties, both static and fatigue. They were able to rank the cements they tested according to a Weibull median fatigue life in the same order as found clinically in the Swedish Hip Register at that time, with Boneloc being the lowest ranked cement. Fractographic analysis showed very different fracture behaviour for Boneloc specimens compared with the other cements. Boneloc fracture surfaces showed evidence of more ductile failure with regions separated into layers and there was little evidence of PMMA particle pull-out. There was some evidence, in the form of circular holes, of zirconia particle pull-out. Furthermore, they noted that Kindt-Larsen et al. found that Boneloc performed better than Simplex P when tested in strain-controlled fatigue but performed worse when tested in stress-controlled fatigue. Given the poor clinical performance of Boneloc, they concluded that stress-controlled fatigue should be recommended as a standard for bone cement testing. Most of these studies did not propose a chemical or microstructural cause for Boneloc’s inferior properties but, given that its elastic modulus was consistently lower than other cements (e.g. 1.62 GPa for Boneloc versus 2.41 GPa for Simplex),85 a possibility is that one of the contributing factors may have been a lower molecular weight than conventional cements; unfortunately none of the aforementioned studies examined the final molecular weight of Boneloc. This would also explain the more ductile behaviour and higher strain-to-failure observed for Boneloc. Another possibility is that the comonomer additions may have been incompatible with the copolymer mix, leading to an inhomogeneous mix. This may be an explanation for the laminar fracture surfaces seen by Harper and Bonfield.59 A link between the butylmethacrylate component and increased creep was also proposed by Thanner et al.88 However, because of the many differences in constituents and additives of various cements, the exact reasons for Boneloc’s inferior strength remain elusive without further testing. Perhaps the most salient material science-related lesson to be learned from the case of Boneloc is the importance of the tensile properties of a bone cement, in particular fatigue resistance. Furthermore, when testing in fatigue, it would appear that stress-controlled fatigue is to be recommended over strain-control.

12.7

Bone cement failure in joint replacements

Up to now, the fracture and fatigue properties of ‘pure’ bone cement specimens have mainly been considered. However, clinically the bone cement

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Debonded implantcement interface: Implant-cement debonding decreases load-bearing area of interface and increases cement stresses and damage accumulation. Relative motion of the implant and cement generates abrasive wear particles’ which can lead to further particulate reactions.

UHMWPE wear particles

Pores and cracks: Pores act as stress raisers and crack initiation sites. Cracks may also initiate from the interdigitated cement-bone interface or from localised debonded regions of the implant-cement interface (such bimaterial interface cracks are prone to branching into the weaker material). Large cracks with jagged faces are likely to experience further abrasion, stimulating particulate reactions

Soft-tissue interface with UHMWPE and PMMA wear particles: Stress-shielding-induced bone resorption and interfacial failures open a route for particulate debris. Weakening of the interface leads to increased relative micromotions with resulting interface damage and bone resorption with the formation of a soft-tissue interface.

Debonded implant-cement interface and branching crack

12.6 Schematic illustrating the interaction of bone cement failure with aseptic loosening of a cemented total hip replacement.

usually forms a composite structure with the other materials it comes in contact with, e.g. bone and metal in a cemented total hip arthroplasty (Fig. 12.6). Huiskes102 has proposed two interacting failure scenarios to account for aseptic loosening of cemented joint reconstructions: (a) the particulate reaction scenario and (b) the damage accumulation scenario. In the ‘particulate reaction scenario’, biological reactions to particulate wear debris cause deterioration of the bone until it no longer supports the implant. In the ‘damage accumulation scenario’, debonding of the prosthesis from the cement, along with microcracking in the cement, proceeds until there is no longer sufficient fixation for the prosthesis. Interactions between failure scenarios can occur from either increased loading of the cement as the living tissues become unable to support the load, or from increased wear particle production as the debonded prosthesis abrades the deteriorating cement layer (see Fig. 12.6 for a schematic illustration of these interactions). Some of the bone cement-related factors affecting these two scenarios will be discussed in more detail in the following sections.

Fracture toughness and fatigue characteristics of bone cements

12.8

279

Failure at interfaces

The interfaces between the cement and the surrounding materials and tissue can have important consequences for the evolution of damage accumulation in the cement. The two most common interfaces are (a) between cement and bone, both cancellous and cortical, and (b) between cement and implant, usually metal (e.g. cobalt chrome or titanium), ceramic (e.g. alumina), or polymer (e.g. UHMWPE).

12.8.1 Cement–bone interface Bone cement may need to adhere to either cancellous and/or cortical bone in many of its clinical applications. As it does not chemically bond with bone, it relies on mechanical interlock to affix to its host material. Obviously, the bond strength with cortical bone is relatively weak, as it can only interlock with surface texture. However, good interdigitation can be achieved with cancellous bone, in particular when the bone is cleaned, e.g. using pulse lavage. Maintenance of the integrity of this interface is crucial to survival of prosthesis fixation in heavily loaded applications such as hip replacements.14,103–105 Retrieval studies14,15 as well as in vitro tests of prosthesis–cement–femur constructs106–110 have demonstrated damage at this interface. The importance of the cement–bone interface has long been evident to clinicians and is frequently used to define aseptic loosening by radiographic analysis.105 However, given its importance, there has been relatively little biomechanical research into its failure mechanisms.111 Some studies have described regional variations in interface strength or morphology and have suggested that this was related to interdigitation.112,113 This was demonstrated by Mann et al.111 by relating quantification of interdigitation, using computed tomography, and measurements of interface failure during monotonic loading. Interface failure typically followed a strain-softening pattern. Fractography showed several characteristic failure patterns: (a) fracture of bone trabeculae at the extent of the cement region leaving bone trabeculae in the cement side; (b) a mixture of bone and cement fracture with cement spicules left in the bone side and bone trabeculae in the cement side; and (c) fracture of cement spicules at the extent of the bone region with cement spicules left in the bone side. A subsequent series of studies by the same group investigated both static and fatigue failure of cement–bone interfaces. They found that the interface was stronger in shear than in tension114 and proceeded to develop mixed-mode failure criteria for the interface for monotonic loading.115,116 Fatigue testing of interface specimens found somewhat different behaviour than for static testing.117,118 In particular, failure tended to dominate at the contact interface (i.e. debonding) between the bone and cement of the interdigitated

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region. Furthermore, deformation tended to progress in a creep-like manner, i.e. permanent deformation of the interface occurred at constant-amplitude cyclic loading. Some differences were also apparent between loading modes – tensile fatigue caused strain softening (i.e. a reduction of stiffness),117 while no such change was observed for shear fatigue.118 However, both loading modes exhibited a three-phase damage process of (a) rapid primary deformation followed by (b) steady-state secondary deformation and ending with (c) rapid deformation to failure over a few cycles. Although initially lagging behind other aspects of bone cement failure analysis, there has been considerable progress made in the study of cement– bone failure mechanisms in recent years. As the mechanical properties of the interface become more established, no doubt attention will turn to the consequences of mechanical damage on biological adaptation, which remains a relatively unexplored area of research.

12.8.2 Cement–metal interface The interface between bone cement and metal prostheses has long been of considerable interest, in particular for cemented hip arthroplasty. In 1976 Fornasier and Cameron119 found in an autopsy retrieval study that femoral prostheses frequently debond from the cement early in the lifetime of the implant, often resulting in a thin fibrous tissue film between the implant and cement. Subsequent mechanical testing of cement–metal interfaces showed that both static120 and fatigue121 strengths of this interface were substantially less than the bulk cement material. Fractographic analysis of ex vivo retrieved cement mantles demonstrated that prosthesis debonding occurs early and is followed by distributed slowly developing fractures in the cement initiated by stress concentrations at the implant–cement interface as well as from pores in the cement.14,15 Bone cement and metal do not chemically bond and so mechanical interlock between the cement and the metal’s surface roughness is required to achieve fixation between the two materials.23,122 Furthermore, there is also a tendency for porosity to form near this interface,35,36,123 which can further reduce the interface strength. Push-out tests have frequently been used to examine the shear strength of the interface in both static and fatigue conditions.120,121,124,125 However, these tests have been shown to be sensitive to boundary conditions so it can be difficult to compare results from different studies.126 Observation of stem–cement interface cracks that had propagated through the thickness of the cement mantle inspired fracture mechanics studies of the bi-material interface under both tensile and torsional loads.127–129 These studies have demonstrated that the material stiffness mismatch at the cement–metal interface leads to mixed-mode loading (modes I (tension), and II (shear)) of interfacial cracks even under application of

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uniaxial mode I loads. Furthermore, one of these studies predicted that the stiffness mismatch can promote crack branching away from the interface into the bulk cement under certain load conditions.127 This situation is particularly evident in torsional loading of the interface, e.g. due to prosthesis rotation.108 The inherent weakness of the interface has led to the development of several measures that aim to strengthen it and delay the onset of loosening. Two approaches have dominated: (a) increased surface roughness of the metal prosthesis and (b) pre-coating the prosthesis with a layer of bone cement. Although increasing surface roughness of the metal does improve static interface strength130,131 and decrease prosthesis subsidence in fatigue testing,132,133 clinical results have indicated poor performance of prostheses with matt surface finishes, in particular straight-taper designs.51,134 Verdonschot et al. demonstrated in a series of studies132,135–137 that this is most likely due to the fact that matt prostheses are still likely to debond and that, when this occurs, abrasive wear of the rough metal surface increases damage and particle generation in the cement. This can provoke biological responses if these particles reach the surrounding tissues. Similarly, precoating the implant has also been demonstrated to increase interface strength120,125,138 but clinical data suggest that some prosthesis designs do not perform well with a pre-coat.139,140 Some have speculated that strengthening the cement–metal interface in this way may increase loading at the cement– bone interface and lead to earlier failure there instead.141 Relatively poor clinical results have led some to question whether strengthening the cement–metal interface is necessary, given that it is likely to debond anyway and the strengthening measures are likely to negatively affect subsequent fatigue of the structure.134,142 Huiskes et al.143 later hypothesised that the question of whether a stem should be matt or polished should be considered in the context of its shape and design objective.

12.9

Residual stress and the initiation of damage

Residual stress is a common problem in manufacturing PMMA components144 and early evidence of its existence in bone cement was reported by Kusy.1 Shrinkage has been described as having three mechanisms of volume change during curing.34,145 Firstly, a volume change of approximately 20% occurs as the liquid monomer polymerises to a higher density solid.144 However, as bone cement is a two-phase mixture of MMA and PMMA beads this volume change is not as large as might be expected. It depends on the ratio of liquid monomer to polymer beads – shrinkage for typical liquid–powder ratios is approximately 7%.145 Secondly, there is bulk expansion due to the formation and expansion of gas bubbles. Finally, there is thermal expansion during the exothermic polymerisation reaction followed

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by thermal shrinkage during cooling to ambient temperature.145–147 The thermal expansion coefficient for a cement mass cooling after polymerisation has been measured to be in the range 7–9%°C−1.145 Whelan et al.146 measured internal strain during polymerisation with fibre-optic Bragg grating sensors and found residual strains of the order of 4000 microstrain, the major portion of which corresponded to shrinkage from the peak temperature. A computational model of this experiment was able to predict the residual strain and further predicted that internal tensile stresses could be generated in this free-shrinking specimen due to a mismatch in shrinkage between the interior and exterior regions of the cement mass (∼3 MPa).148 Some studies have also found evidence for residual stress generation in the inter-bead matrix of cured cement in free-shrinking cement specimens.6,30 Since this thermal contraction is frequently restricted by the composite nature of an implant–cement–bone reconstruction, even higher residual stress generation can be expected in many clinical applications. Although shrinkage appears to be the predominant mechanism of residual stress generation, the point at which the rapidly polymerising cement mass begins to support internal loading and stress is difficult to observe and measure directly and remains a frequent topic of investigation.146,149–154 Although the mechanism of stress generation has been the subject of some debate, the effects of residual stress on the damage accumulation scenario for bone cement are relatively clear. A number of in vitro models have found cracks in cement mantles prior to the application of any loads.106,107,149,155 Some of these studies have clearly demonstrated that residual stress can interact with porosity to initiate cracks in the cement mantle of intramedullary prostheses.149,156 Acoustic emission monitoring of bone cement polymerisation has also confirmed that cracking occurs during curing.151

12.10 Viscoelasticity, creep and creep–fatigue interaction Like other amorphous polymers, PMMA is viscoelastic – it experiences two relaxation process which are due to:57,144 (a) motion and rotation of the molecular backbone of the polymer chain (α-relaxation); (b) rotation of the COOCH3 side group (β-relaxation). The glass transition corresponds to the α-relaxation and occurs at approximately 105–110°C, while β-relaxation occurs at approximately 50°C; these relaxation processes are thus significantly retarded at the typical service (i.e. body) temperatures of bone cement. Nevertheless, both creep and stress relaxation have been demonstrated for bone cement at body temperature.157–160

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Interaction between creep and fatigue is a frequently studied topic in failure analysis of metals, in particular in high-temperature applications such as nuclear engineering.161 This is because fatigue in metals is characterised mainly by transcrystalline microcracks initiating from surface flaws, while creep occurs due to cavity formation at grain boundaries162 – thus, simultaneous creep and fatigue tends to accelerate damage accumulation in metals. In contrast, bone cement creep is due to molecular rearrangements, as described above. These rearrangements lead to relaxation of stress and can therefore be expected to decrease damage accumulation. There has been relatively little experimental investigation of creep during fatigue loading and any possible interaction between the two processes. Several studies have measured creep under cyclic loading conditions.163–167 However, such studies cannot answer the question of whether creep can decrease damage accumulation. In their investigations on the effect of implant surface finish, Verdonschot and Huiskes132 did show increased creep with less damage for polished stems compared with rough stems; however, this was primarily due to abrasive wear. Nevertheless, the greater capacity for creep around the polished stems may have been a contributing factor. Nguyen et al.168 considered creep zones in fatigue crack propagation using a theoretical model based on creep constants obtained from the literature. They proposed that creep acts to reduce the stress intensity ahead of a crack tip similar to a plastic zone in metals. During low-frequency loading this creep zone has more time to grow and hence reduce the stress intensity. In contrast, high-frequency loading does not allow enough time for the creep zone to relieve the stress intensity and the situation tends towards the linear elastic fracture mechanics assumption. Computational simulations have also predicted creep-induced deceleration of damage accumulation169,170 and at this time it seems that this is the most plausible hypothesis for an interaction between creep and fatigue in bone cement.

12.11 Fracture and fatigue properties Some of the factors affecting both static and fatigue failure of acrylic bone cements have been presented in the preceding sections. This final section attempts to provide some fracture toughness and fatigue data for two common cement brands, Simplex P and Palacos R. These were chosen primarily as they are both commonly used, have different viscosities when in the dough phase and have been widely studied. Inter-study comparisons are exceedingly difficult to make due to a wide array of differing test parameters (e.g. additives, preparation methods, storage conditions), loading modes (e.g. different stress levels, fully reversed versus tension-only axial loading or flexural loading, specimen crosssectional shape, etc.) and reported measures (e.g. mean and standard

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deviation of fatigue life at a given stress (Nf), Weibull mean fatigue life (NW), fatigue performance index (IW) obtained from three-parameter Weibull analysis, etc.). Standardisation of bone cement testing and reporting has received increasing attention in recent years and is likely to be a subject of debate for some time to come.61,171 In Tables 12.1 and 12.2, fracture toughness values were taken from tests using a variety of different specimen configurations and test parameters. Fully reversed fatigue of cylindrical specimens at 15 MPa was chosen as it Table 12.1 Selected fracture toughness and fatigue life values for simplex P bone cement. Nf, mean and standard deviation of fatigue life at a given stress; Nw, mean Weibull fatigue life. KIC, fracture toughness Property

Hand mixed

KIC (Mpa(m)0.5)

1.52 1.24 1.33–1.41 1.03 1.5

Centrifuged

Vacuum mixed

172 173 174 175 40

1.55 1.4 1.78

1.71 ± 0.05 Fatigue (±15 MPa)

Reference

176 177 79

24 218 11 841 15 147 ± 24 690

34 594 34 239 ± 5 889

2 101 ± 2 459

71 479 ± 22 626

12 838 ± 21 838 65 325 ± 92 037

(NW)178 (NW)179 (Nf)180 (Nf)181 (Nf)182

Table 12.2 Selected fracture toughness and fatigue life values for Palacos R bone cement for different mixing methods Nf, mean and standard deviation of fatigue life at a given stress; Nw, mean Weibull fatigue life; Iw, fatigue performance index obtained from three-parameter Weibull analysis Property

Hand mixed

Centrifuged

KIC (Mpa(m)0.5)

2.02

1.96

7 753 11 504 ± 6387 75 433 46 174 ± 10 222

Reference 40

1.81 1.92 ± 0.10 1.95 ± 0.17

1.75 ± 0.07 1.76 ± 0.11 1.78 1.85 ± 0.12 Fatigue (±15 MPa)

Vacuum mixed

183 184 74 5 185

8 800 8 687 ± 1 258 177 477 130 750 ± 36 973

(NW)65 (Nf)180 (IW)186 (Nf)187

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is one of the most widely available tests in the literature. Data from common studies are presented in common rows and the reported measure and reference are provided in the rightmost column. Although values are given for different mixing methods, these data should not be used to compare mixing methods across studies due to the many differences in test configurations and reported measures. Data from standard formulations were chosen – i.e. including the standard radiopacifiers for a given cement but without other additives such as antibiotics.

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87 linder, l. Boneloc – The Christiansen experience revisited. Acta Orthop Scand 66, 205–206(1995). 88 thanner, j., freij-larsson, c., kärrholm, j., malchau, h. and wesslén, b. Evaluation of Boneloc. Chemical and mechanical properties, and a randomized clinical study of 30 total hip arthroplasties. Acta Orthop Scand 66, 207–214(1995). 89 riegels-nielsen, p., sørensen, l., andersen, h.m. and lindequist, s. Boneloc cemented total hip prostheses. Loosening in 28/43 cases after 3–38 months. Acta Orthop Scand 66, 215–217(1995). 90 engesaeter, l.b., furnes, a., havelin, l.i., lie, s.a. and vollset, s.e. [The hip registry. Good economy for society]. Tidsskr Nor Laegeforen 116, 3025– 3027(1996). 91 furnes, a., lie, s.a., havelin, l.i., engesaeter, l.b. and vollset, s.e. The economic impact of failures in total hip replacement surgery: 28,997 cases from the Norwegian Arthroplasty Register, 1987–1993. Acta Orthop Scand 67, 115– 121(1996). 92 nilsen, a.r. and wiig, m. Total hip arthroplasty with Boneloc: loosening in 102/157 cases after 0.5–3 years. Acta Orthop Scand 67, 57–59(1996). 93 furnes, o., lie, s.a., havelin, l.i., vollset, s.e. and engesaeter, l.b. Exeter and charnley arthroplasties with Boneloc or high viscosity cement. Comparison of 1127 arthroplasties followed for 5 years in the Norwegian Arthroplasty Register. Acta Orthop Scand 68, 515–520(1997). 94 wedderkopp, n., andersen-ranberg, f., andersen, m.b. and termansen, n.b. Aseptic loosening of BonelocR cemented hip prostheses. Int Orthop 21, 87–90(1997). 95 nilsson, k.g. and dalen, t. Inferior performance of Boneloc® bone cement in total knee arthroplasty – a prospective randomized study comparing Boneloc® with Palacos® using radiostereometry (RSA) in 19 patients. Acta Orthop Scand 69, 479–483(1998). 96 gebuhr, p., stentzer, k., thomsen, f. and levi, n. Failure of total hip arthroplasty with Boneloc bone cement. Acta Orthop Belg 66, 472–476(2000). 97 havelin, l.i., engesaeter, l.b., espehaug, b., furnes, o., lie, s.a. and vollset, s.e. The Norwegian Arthroplasty Register: 11 years and 73 000 arthroplasties. Acta Orthop Scand 71, 337–353(2000). 98 walczak, j.p., d’arcy, j.c., ross, k.r., james, s.e., bonnici, a.v., koka, s.r. and morris, r.w. Low-friction arthroplasty with Boneloc bone-cement: outcome at 2 to 4 years. J Arthroplasty 15, 205–209(2000). 99 markel, d.c., hoard, d.b. and porretta, c.a. Cemented total hip arthroplasty with Boneloc bone cement. J South Orthop Assoc 10, 202–208(2001). 100 petruskevicius, j., nielsen, s. and overgaard, s. [Total knee arthroplasty with Boneloc cement. Clinical and radiological results]. Ugeskr Laeger 163, 6433– 6436(2001). 101 abdel-kader, k.f., allcock, s., walker, d.i. and chaudhry, s.b. Boneloc bonecement: experience in hip arthroplasty during a 3-year period. J Arthroplasty 16, 811–819(2001). 102 huiskes, r. Mechanical failure in total hip arthroplasty with cement. Curr Orthop 7, 239–247(1993). 103 holz, u., hemminger, w. and gasse, h. [Mechanical tests with polymethylmethacrylate in fresh specimens and specimens obtained from loosened

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prosthesis (author’s translation)]. Arch Orthop Trauma Surg 91, 121–127 (1978). krause, w.r., krug, w. and miller, j. Strength of the cement–bone interface. Clin Orthop Relat Res 163, 290–299(1982). harris, w.h. Is it advantageous to strengthen the cement–metal interface and use a collar for cemented femoral components of total hip replacements? Clin Orthop Relat Res 285, 67–72(1992). mccormack, b.a.o., walsh, c.d., wilson, s.p. and prendergast, p.j. A statistical analysis of microcrack accumulation in PMMA under fatigue loading: applications to orthopaedic implant fixation. Int J Fatigue 20, 581–593(1998). mccormack, b.a.o. and prendergast, p.j. Microdamage accumulation in the cement layer of hip replacements under flexural loading. J Biomechanics 32, 467– 475(1999). mccormack, b.a.o., prendergast, p.j. and o’dwyer, b. Fatigue of cemented hip replacements under torsional loads. Fatigue Fracture Engng Mater Struct 22, 33–40(1999). race, a., miller, m.a., ayers, d.c. and mann, k.a. Early cement damage around a femoral stem is concentrated at the cement/bone interface. J Biomechanics 36, 489–496(2003). mann, k.a., gupta, s., race, a., miller, m.a., cleary, r.j. and ayers, d.c. Cement microcracks in thin-mantle regions after in vitro fatigue loading. J Arthroplasty 19, 605–612(2004). mann, k.a., ayers, d.c., werner, f.w., nicoletta, r.j. and fortino, m.d. Tensile strength of the cement–bone interface depends on the amount of bone interdigitated with PMMA cement. J Biomechanics 30, 339–346(1997). bean, d.j., convery, f.r., woo, s.l. and lieber, r.l. Regional variation in shear strength of the bone–polymethylmethacrylate interface. J Arthroplasty 2, 293– 298(1987). maher, s.a. and mccormack, b.a.o. Quantification of interdigitation at bone cement/cancellous bone interfaces in cemented femoral reconstructions. Proc Instn Mech Engrs, Part H: J Engng Med 213, 347–354(1999). mann, k.a., werner, f.w. and ayers, d.c. Mechanical strength of the cement– bone interface is greater in shear than in tension. J Biomechanics 32, 1251–1254 (1999). mann, k.a., mocarski, r., damron, l.a., allen, m.j. and ayers, d.c. Mixed-mode failure response of the cement–bone interface. J Orthop Res 19, 1153–1161 (2001). mann, k.a. and damron, l.a. Predicting the failure response of cement–bone constructs using a non-linear fracture mechanics approach. J Biomech Eng 124, 462–470(2002). kim, d., miller, m.a. and mann, k.a. Creep dominates tensile fatigue damage of the cement–bone interface. J Orthop Res 22, 633–640(2004). kim, d., miller, m.a. and mann, k.a. A fatigue damage model for the cement– bone interface. J Biomechanics 37, 1505–1512(2004). fornasier, v.l. and cameron, h.u. The femoral stem/cement interface in total hip replacement. Clin Orthop Relat Res 116, 248–252(1976). ahmed,a.m., raab,s. and miller,j.e. Metal/cement interface strength in cemented stem fixation. J Orthop Res 2, 105–118(1984).

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121 raab, s., ahmed, a.m. and provan, j.w. The quasi-static and fatigue performance of the implant/bone–cement interface. J Biomed Mater Res 15, 159–182(1981). 122 davies, j.p. and harris, w.h. Tensile bonding strength of the cement–prosthesis interface. Orthopedics 17, 171–173(1994). 123 mann, k.a., damron, l.a., miller, m.a., race, a., clarke, m.t. and cleary, r.j. Stem-cement porosity may explain early loosening of cemented femoral hip components: experimental–computational in vitro study. J Orthop Res 25, 340–350(2007). 124 mann, k.a., bartel, d.l., wright, t.m. and ingraffea, a.r. Mechanical characteristics of the stem–cement interface. J Orthop Res 9, 798–808(1991). 125 davies, j.p. and harris, w.h. Strength of cement–metal interfaces in fatigue: comparison of smooth, porous and precoated specimens. Clin Mater 12, 121– 126(1993). 126 harrigan, t.p., kareh, j. and harris, w.h. The influence of support conditions in the loading fixture on failure mechanisms in the push-out test: a finite element study. J Orthop Res 8, 678–684(1990). 127 mccormack, b.a. and prendergast, p.j. An analysis of crack propagation paths at implant/bone–cement interfaces. J Biomech Eng 118, 579–585(1996). 128 mann, k.a., edidin, a.a., ordway, n.r. and manley, m.t. Fracture toughness of CoCr alloy-PMMA cement interface. J Biomed Mater Res 38, 211–219 (1997). 129 heuer, d.a. and mann, k.a. Fatigue fracture of the stem–cement interface with a clamped cantilever beam test. J Biomech Eng 122, 647–651(2000). 130 müller, r.t. and schürmann, n. Shear strength of the cement metal interface – an experimental study. Arch Orthop Trauma Surg 119, 133–138(1999). 131 wang, j., taylor, m., flivik, g. and lidgren, l. Factors affecting the static shear strength of the prosthetic stem–bone cement interface. J Mater Sci: Mater Med 14, 55–61(2003). 132 verdonschot, n. and huiskes, r. Surface roughness of debonded straighttapered stems in cemented THA reduces subsidence but not cement damage. Biomaterials 19, 1773–1779(1998). 133 verdonschot, n., tanck, e. and huiskes, r. Effects of prosthesis surface roughness on the failure process of cemented hip implants after stem–cement debonding. J Biomed Mater Res 42, 554–559(1998). 134 schmalzried, t.p., zahiri, c.a. and woolson, s.t. The significance of stem–cement loosening of grit-blasted femoral components. Orthopedics 23, 1157–1164 (2000). 135 verdonschot, n. and huiskes, r. Mechanical effects of stem cement interface characteristics in total hip replacement. Clin Orthop Relat Res 329, 326–336 (1996). 136 verdonschot, n. and huiskes, r. Cement debonding process of total hip arthroplasty stems. Clin Orthop Relat Res 336, 297–307(1997). 137 verdonschot, n. and huiskes, r. The effects of cement–stem debonding in THA on the long-term failure probability of cement. J Biomechanics 30, 795–802 (1997). 138 barb, w., park, j.b., kenner, g.h. and recum, a.f.v. Intramedullary fixation of artificial hip joints with bone cement-precoated implants. I. Interfacial strengths. J Biomed Mater Res 16, 447–458(1982).

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139 ong, a., wong, k.l., lai, m., garino, j.p. and steinberg, m.e. Early failure of precoated femoral components in primary total hip arthroplasty. J Bone Joint Surg Am 84-A, 786–792(2002). 140 cannestra, v.p., berger, r.a., quigley, l.r., jacobs, j.j., rosenberg, a.g. and galante, j.o. Hybrid total hip arthroplasty with a precoated offset stem. Four to nine-year results. J Bone Joint Surg Am 82, 1291–1299(2000). 141 gardiner, r.c. and hozack, w.h. Failure of the cement–bone interface. A consequence of strengthening the cement–prosthesis interface? J Bone Joint Surg 76-B, 49–52(1994). 142 ling, r.s.m. The use of a collar and precoating on cemented femoral stems is unnecessary and detrimental. Clin Orthopa Relat Res 285, 73–83 (1992). 143 huiskes, r., verdonschot, n. and nivbrant, b. Migration, stem shape, and surface finish in cemented total hip arthroplasty. Clin Orthop Relat Res 355, 103–112 (1998). 144 kine, b.b. and novak, r.w. Acrylic and methacrylic ester polymers, In Encyclopaedia of Polymer Science and Engineering, Eds H.F. Mark, N.M. Bikales, C.G. Overberger, G. Menges, & J.I. Kroschwitz, pp. 234–299 (John Wiley and Sons, New York, 1987). 145 ahmed, a.m., pak, w., burke, d.l. and miller, j. Transient and residual stresses and displacements in self-curing bone cement – Part I: Characterization of relevant volumetric behaviour of bone cement. J Biomech Engng, Trans ASME 104, 21–27(1982). 146 whelan, m.p., kenny, r.p., cavalli, c., lennon, a.b. and prendergast, p.j. In Proceedings of the 12th conference of the European Society of Biomechanics, Eds P.J. Prendergast, A.J. Carr, & T.C. Lee p. 252 (Royal Academy of Medicine in Ireland, Dublin, Ireland, 2000). 147 muller, s.d., green, s.m. and mccaskie, a.w. The dynamic volume changes of polymerising polymethyl methacrylate bone cement. Acta Orthop Scand 73, 684–687(2002). 148 lennon, a.b., prendergast, p.j., whelan, m.p., kenny, r.p. and cavalli, c. In Proceedings of the 12th conference of the European Society of Biomechanics, Eds P.J. Prendergast, T.C. Lee, & A.J. Carr, p. 253 (Royal Academy of Medicine in Ireland, Dublin, Ireland, 2000). 149 lennon, a.b. and prendergast, p.j. Residual stress due to curing can initiate damage in porous bone cement: experimental and theoretical evidence. J Biomechanics 35, 311–321(2002). 150 ahmed, a.m., pak, w., burke, d.l. and miller, j. Transient and residual stresses and displacements in self-curing bone cement – Part II: Thermoelastic analysis of the stem fixation system. J Biomech Engng, Trans ASME 104, 28–37 (1982). 151 roques, a., browne, m., taylor, a., new, a. and baker, d. Quantitative measurement of the stresses induced during polymerisation of bone cement. Biomaterials 25, 4415–4424(2004). 152 nuño, n. and avanzolini, g. Residual stresses at the stem–cement interface of an idealized cemented hip stem. J Biomechanics 35, 849–852(2002). 153 nuño, n. and amabili, m. Modelling debonded stem–cement interface for hip implants: effect of residual stresses. Clin Biomech 17, 41–48(2002).

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154 li, c., wang, y. and mason, j. The effects of curing history on residual stresses in bone cement during hip arthroplasty. J Biomed Mater Res B Appl Biomater 70, 30–36(2004). 155 lennon, a.b., mccormack, b.a.o. and prendergast, p.j. The relationship between cement fatigue damage and implant surface finish in proximal femoral prostheses. Med Engng & Physics 25, 833–841(2003). 156 orr, j.f., dunne, n.j. and quinn, j.c. Shrinkage stresses in bone cement. Biomaterials 24, 2933–2940(2003). 157 lee, a.j.c., ling, r.s.m. and vangala, s.s. The mechanical properties of bone cements. J Med Engng Technol 1(3),137–140(1977). 158 pal, s. and saha, s. Stress relaxation and creep behaviour of normal and carbon fibre reinforced acrylic bone cement. Biomaterials 3, 93–96(1982). 159 ebramzadeh, e., mina-araghi, m., clarke, i.c. and ashford, r. Loosening of well-cemented total-hip femoral prosthesis due to creep of the cement. In Corrosion and Degradation of Implant Materials, Eds A.C. Franker, & G.D. Griffin, pp. 373–399 (American Society for Testing and Materials, Philadelphia, 1983). 160 yetkinler, d.n. and litsky, a.s. Viscoelastic behaviour of acrylic bone cements. Biomaterials 19, 1551–1559(1998). 161 chaboche, j.l. Continuum damage mechanics: present state and future trends. Nuclear Engng Design 105, 19–33(1987). 162 chaboche, j.l. Thermodynamically founded CDM models for creep and other conditions, CISM Courses. In Creep and damage in Materials and Structures, Eds H.J. Altenbach & J.J. Skrzypek, pp. 209–283 (Springer-Verlag, New York, 1999). 163 verdonschot, n. and huiskes, r. Creep properties of three low temperaturecuring bone cements: a preclinical assessment. J Biomed Mater Res 53, 498–504(2000). 164 verdonschot, n. and huiskes, r. In Second International Symposium on Computer Methods in Biomechanics and Biomedical Engineering, Ed. J. Middleton, pp. 25–33 (Gordon and Breach Publisher, The Netherlands, 1995). 165 verdonschot, n. and huiskes, r. Creep behavior of hand-mixed Simplex P bone cement under cyclic tensile loading. J Appl Biomater 5, 235–243(1994). 166 jeffers, j.r.t., browne, m. and taylor, m. Damage accumulation, fatigue and creep behaviour of vacuum mixed bone cement. Biomaterials 26, 5532–5541 (2005). 167 murphy, b.p. Aspects of the fatigue behaviour of acrylic bone cement. PhD Thesis, University of Dublin, Trinity College (2001). 168 nguyen, n.c., maloney, w.j. and dauskardt, r.h. Reliability of PMMA bone cement fixation: fracture and fatigue crack-growth behaviour. J Mater Sci: Mater Med 8, 473–483(1997). 169 verdonschot, n. and huiskes, r. Acrylic cement creeps but does not allow much subsidence of femoral stems. J Bone Joint Surg 79-B, 665–669(1997). 170 stolk, j., verdonschot, n., murphy, b.p., prendergast, p.j. and huiskes, r. Finite element simulation of anisotropic damage accumulation and creep in acrylic bone cement. Engng Fracture Mechanics 71, 513–528(2003). 171 lewis, g. and nyman, j.s. Toward standardization of methods of determination of fracture properties of acrylic bone cement and statistical analysis of test results. J Biomed Mater Res 53, 748–768(2000).

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172 krause, w.r., krug, w.h. and miller, j.e. In Proceedings of the 26th Annual Meeting of the Orthopedic Research Society, Atlanta, Georgia, p. 253 (1980). 173 weber, s.c. and bargar, w.l. A comparison of the mechanical properties of Simplex, Zimmer, and Zimmer low viscosity bone cements. Biomater Med Devices Artif Organs 11, 3–12(1983). 174 krause, w. and hofmann, a. Antibiotic impregnated acrylic cements: a comparative study of the mechanical properties. J Bioactive Compt Polym 4, 345– 361(1989). 175 pilliar, r.m., vowles, r. and williams, d.f. Fracture toughness testing of biomaterials using a mini-short rod specimen design. J Biomed Mater Res 21, 145–154(1987). 176 lewis, g. Effect of methylene blue on the fracture toughness of acrylic bone cement. Biomaterials 15, 1024–1028(1994). 177 friis, e.a., stromberg, l.j., cooke, f.w. and mcqueen, d.a. In The 19th Annual Meeting of the Society for Biomaterials, Birmingham, Alabama, p. 301 (1993). 178 lewis, g., janna, s. and carroll, m. Effect of test frequency on the in vitro fatigue life of acrylic bone cement. Biomaterials 24, 1111–1117(2003). 179 davies, j.p., o’connor, d.o., greer, j.a. and harris, w.h. Comparison of the mechanical properties of Simplex P, Zimmer Regular, and LVC bone cements. J Biomed Mater Res 21, 719–730(1987). 180 davies, j.p., jasty, m., o’connor, d.o., burke, d.w., harrigan, t.p. and harris, w.h. The effect of centrifuging bone cement. J Bone Joint Surg Br 71, 39–42(1989). 181 davies, j.p., o’connor, d., burke, d.w. and harris, w.h. In Transactions of the 34th Annual Meeting of the Orthopedic Research Society, p. 221 (1988). 182 davies, j.p., singer, g. and harris, w.h. The effect of a thin coating of polymethylmethacrylate on the torsional fatigue strength of the cement–metal interface. J Appl Biomater 3, 45–49(1992). 183 lewis, g. Effect of lithotriptor treatment on the fracture toughness of acrylic bone cement. Biomaterials 13, 225–229(1992). 184 lewis, g. Apparent fracture toughness of acrylic bone cement: effect of test specimen configuration and sterilization method. Biomaterials 20, 69–78 (1999). 185 lewis, g. and mladsi, s. Correlation between impact strength and fracture toughness of PMMA-based bone cements. Biomaterials 21, 775–781(2000). 186 kim, h.y. and yasuda, h.k. Improvement of fatigue properties of poly(methyl methacrylate) bone cement by means of plasma surface treatment of fillers. J Biomed Mater Res 48, 135–142(1999). 187 lewis, g. and janna, s. The influence of the viscosity classification of an acrylic bone cement on its in vitro fatigue performance. Biomed Mater Eng 14, 33–42(2004).

13 Dynamic mechanical properties of bone cements S. N. N A Z H AT, McGill University, Canada; and J. V. CAU I C H R O D R Í G U E Z, Centro de Investigacíon Cientifica de Yucatan A.C., Mexico

Abstract: Polymeric bone cements are viscoelastic materials. Dynamic mechanical analysis (DMA) is a powerful technique for the characterisation of polymers. Viscoelastic parameters of storage and loss moduli, mechanical loss factors (tan d) as well as important thermal transition temperature and polymer relaxation processes are easily quantified as a function of temperature and frequency. This chapter explores the viscoelasticity of polymers, DMA as a characterisation tool and the use of DMA in the characterisation of polymeric bone cements. Key words: dynamic mechanical analysis, glass transition temperature, modulus, frequency, mechanical loss factor.

13.1

A brief introduction to viscoelasticity in polymers

The complex nature of polymeric materials results in their mechanical properties being both time and temperature dependent, i.e. viscoelastic. Simplistically, viscoelastic materials simultaneously display both viscous and elastic behaviours. While elastic materials store energy as they are being deformed and viscous materials dissipate this energy, polymers on the other hand combine these two characteristics. Elastic materials obey Hooke’s law where the stress is proportional to the strain and is independent of the loading time or rate. In contrast, liquids have a mechanical behaviour that can be represented at low strain rates by Newton’s law, where the stress is proportional to the strain rate and independent of the strain, i.e. time dependent. Polymers behave as solids at low temperatures and high strain rates, and as liquids at high temperatures and low strain rates. If a stress or strain is applied, the polymer’s response depends on the rate and period of loading.1 The study of viscoelasticity is the investigation of the interplay of the elasticity, flow and molecular motion. There are three examples of viscoelastic responses in a polymer: 1

Creep – a delayed strain response after a rapid application of stress. Here, a constant load is applied on a material. After an initial elastic

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strain response the material will continue to undergo deformation for a considerable length of time. 2 Stress relaxation – occurs when a material responds quickly to strain, and subsequent decay of stress is observed. Here a sample is deformed rapidly under an applied stress and the material is held under constant strain and the decaying stress is measured as a function of time. 3 Dynamic response – this is when a material responds to a steady sinusoidal stress. This produces a strain oscillating with the same frequency, but out of phase with the stress. The response of a material is measured over a wide range of temperatures and frequencies of the applied loading. Molecules perturbed in this way store a portion of the imparted energy elastically, and dissipate a portion in the form of heat. Information can be obtained about the molecular motions in the sample, and these can affect the modulus, damping characteristics and structural transitions. This chapter is mainly concerned with the third form of viscoelastic behaviour of polymeric bone cements. The first part of this chapter describes the general viscoelastic behaviour of polymers and their properties, while the second part explores recent literature on the application of dynamic mechanical analysis (DMA) in the characterisation of polymeric bone cements and their composites.

13.2

Regions of viscoelasticity

The three major factors that influence the properties of polymers are the molecular configuration, the chain conformation, and the molar mass of the polymer chains and its distribution.2 Polymer molecules or chains are generally very large with a variety of conformations that prevents them from being a totally ordered solid, while their high molar mass results in chain entanglements, thus preventing them from being fluids. Compared with other classes of materials, temperature has a critical effect on the behaviour of polymers. It can also be generalised that, at high temperatures, molecules possess energy for (a) translational rotation of parts of the chains, (b) rotational movement of parts of molecules and (c) vibration of individual molecular bonds. This is in contrast to the behaviour of the chains at low temperature where the molecules possess energy for the latter two movement types, and may also be impeded by bonds with neighbouring chains. The viscoelastic regions of polymers can be represented by observing the change in the modulus as a function of temperature, which can be split into five regions depending on their temperature and state. Figure 13.1 shows this relationship schematically for a linear amorphous polymer (represented by the continuous line).2 In region 1, the polymers possess the

Orthopaedic bone cements 11

10

1

9

10 log E (dyne/cm2)

2

8

9

7

8 3 7

4

5

6 5 4

6

log E (Pa)

298

5

4 3

Temperature

13.1 The modulus–temperature relationships of polymers describing the five regions of viscoelasticity. The continuous line shows the behaviour of a linear amorphous polymer: region 1, glassy phase; region 2, glass transition temperature (Tg); region 3, rubbery plateau region; region 4, rubbery flow region; region 5, viscous flow region. The dashed line shows the behaviour of a semi-crystalline polymer with a higher modulus value for the rubbery plateau region as the crystals behave as reinforcing agents until the melting temperature. The dotted line shows the behaviour of a crosslinked network polymer with no rubbery flow and melting regions. (Taken from Introduction to Physical Polymer Science, Fourth Edition, L.H. Sperling, Wiley InterScience, 2006. Reprinted with permission of John Wiley & Sons, Inc.)

highest modulus values and are described as glassy. In this state, the chains in the polymer have only enough energy for movements (b) and (c) described above. Region 2 is the glass transition region. The glass transition temperature, Tg, of a polymer is arguably its most important property. The glass transition region is characterised by a rapid significant reduction in the modulus over a short temperature range. Qualitatively, it can be defined as the temperature at which there is an onset of long-range molecular motion. For example, while only 1–4 chains are involved in motions below Tg, some 10–50 chain atoms attain sufficient thermal energy to move in a coordinated manner in the glass transition region and the material becomes leathery in nature in this region. Region 3 is described as the rubbery plateau region where long-range segmental motion occurs but the thermal energy is insufficient to overcome entanglement interactions that inhibit flow. Hence, in this region, the modulus will continue to decrease but at a much lower rate than in region 2, and the width of the plateau region depends on the molar mass of the polymer; the higher the molar mass, the wider the width. Region 4 is described as the rubbery flow region, wherein an increase in temperature will cause the material to begin to flow, and a further decrease in modulus is seen over a narrow temperature range. Here, the polymer displays either rubber plasticity or flow properties depending on the time scale of the experiment. In short-time-scale experiments (high

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frequency) the physical entanglements are unable to relax and the material constituents behave in a rubbery fashion. With longer times, the increased molecular motion imparted by the temperature permits the movement of assemblies of chains in a co-ordinated manner and hence the polymer flows. Finally, region 5 is the viscous flow region, where the increased energy allotted to the chains permits the entanglements to open up and flow as individual molecules. Here, the modulus and viscosity are related through the molecular relaxation time. Two other morphologic cases can also exist in polymers: semi-crystalline and crosslinked polymers. The dashed and dotted lines in Fig. 13.1 represent the modulus–temperature relationships of these morphologies, respectively. While both of these polymer morphologies experience regions 1 and 2, semi-crystalline polymers have a longer rubbery plateau region 3 and a higher value for the modulus compared with the linear amorphous polymer during the rubbery plateau region since the crystalline regions act as reinforcing agent within the rubber (a crystal-reinforced rubber). The crystalline phase maintains this characteristic until it reaches the melting temperature. Crosslinked polymers, on the other hand, do not melt, or display flow behaviour, i.e. they do not experience regions 4 and 5 at all temperatures below degradation temperature. Therefore, these crosslinked polymers cannot be melt processed.

13.3

Theory of dynamic mechanical analysis (DMA)

DMA is a powerful technique for the analysis of polymers and their composites. Over the last few decades it has been recognised as a tool that gives critical information on the polymer structure and composition. Usually, the modulus values represented in Fig. 13.1 – which can be calculated, for example, from the linear region of a stress–strain curve generated from quasi-static standard mechanical testing such as tensile, flexural or shear testing – are complex in nature. DMA resolves this complex modulus into its real and imaginary components, which are also defined as storage and loss moduli, respectively.3 DMA measures the response or deformations of a material to periodic or varying forces. Generally, the applied force and the resulting deformation vary sinusoidally with time, and from such tests it is possible to obtain simultaneously both an elastic modulus and a mechanical damping. This modulus may be shear, tensile or bulk, depending on the type of loading, while the mechanical damping or loss factor (tan d) gives the energy dissipated as heat during deformation. During a typical DMA run, a sinusoidal stress is applied on a material; the resultant sinusoidal strain will be out of phase. This phase difference, along with the amplitudes of the stress and strain waves, are used to determine a variety of fundamental material parameters.

300

Orthopaedic bone cements e°

s′

d

s°

s″

13.2 A vector diagram illustrating the role of the phase angle and the relationship between stress and strain during DMA. (Adapted from Introduction to Physical Polymer Science, Fourth Edition, L.H. Sperling, Wiley Inter-Science, 2006.)

The stress and strain are not in phase, and the strain lags behind the stress by the phase angle, d. This relationship can be explained through a vector diagram demonstrating the relationship between stress and strain during DMA as shown in Fig. 13.2. The applied amplitude stress (so) vector can be resolved into its storage and loss components. The phase angle (d) also defines the in-phase and out-of-phase component of the stress. Therefore, by using the vector relationship:

V I = V o cos G

[13.1]

V II = V o sin G

[13.2]

and

The dynamic modulus can be expressed as EI =

VI = E * cos G Ho

[13.3]

E II =

V II = E * sin G Ho

[13.4]

where E*, EI and EII are the complex, storage and loss moduli, respectively, and e o is the amplitude strain. In terms of the complex notation and combined with Fig. 13.3 E* = E I + iE II and

[13.5]

Dynamic mechanical properties of bone cements

E∗

301

E″ d E′

13.3 The vector triangle representing the relationship between complex (E*), storage (E′) and loss (E″) moduli measured through DMA.

E = E * = ( E I2 + E II2 )

1/2

[13.6]

Therefore E II = tan G EI

[13.7]

It can also be seen that, when the phase angle is small, the value of the storage modulus will be very close to that of the complex modulus described above.

13.4

Material properties measured through dynamic mechanical analysis

DMA measures the response of a material to a sinusoidal stress over a wide range of temperatures and frequencies and is especially sensitive to the chemical and physical nature of polymers and their composites, where the stress and strain are out of phase. The main parameters obtained from DMA are storage modulus (EI), which represents the elastic phase of a system and is equivalent to the energy stored through deformation, loss modulus (EII) which represents the viscous phase of a system and is equivalent to the energy dissipated through deformation, and tan d, the mechanical damping factor and is the ratio of EII/EI. DMA can also pinpoint thermal transitions, e.g. typical output of tan d versus temperature will display a peak at Tg. Above Tg, the peaks correspond to the crystalline regions and eventually the melting temperature (Tm) is exhibited, if the polymer is semicrystalline. Morphologically, some polymers combine non-crystallinity (amorphous) and crystallinity, order and disorder in the solid state, and Tg and Tm are associated with the amorphous and crystalline regions, respectively. However, the ability of a polymer to possess crystalline phases depends on a number of factors including chemical composition, backbone chain flexibility, bulkiness of side groups and stereoregularity. Since Tm is

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only associated with the crystalline phase in a polymer, and not all polymers are capable of exhibiting crystalline regions, Tm does not occur in all polymers. With a dynamic test, a sample can be investigated over a wide temperature range in a short time and the overall performance of the material can be predicted from the results. Most DMA measurements are made using a single frequency and constant strain amplitude with varying temperature. Further information can be obtained with measurements where the amplitude of deformations is varied or where multiple frequencies are used. The mechanical damping factor, tan d, is an indication of the energy absorbing potential of a material, i.e. the ratio of energy dissipated as heat to the maximum energy stored in a material during a deformation cycle.4 Damping is often the most sensitive indicator of all kinds of molecular motions that occur in a material. These motions are of significant practical importance in determining the mechanical behaviour of polymers. Thus the absolute value of the damping, its sensitivity to the strain magnitude, temperature and frequency at which damping peaks occur are of considerable interest. Many mechanical properties are intimately related to damping, including fatigue life, toughness and impact strength, breaking strain, wear and coefficient of friction.

13.5

Applications of dynamic mechanical analysis in the characterisation of polymeric biomaterials

Over the past two decades, DMA has proven to be a useful technique for the characterisation of polymeric and composite biomaterials since it not only gives a quantitative assessment of material properties such as modulus and damping, but also provides structural and morphological information.5 This is because the dynamic mechanical properties of materials are sensitive to all kinds of transitions, e.g. Tg and Tm – which are thermodynamically classed as second- and first-order transitions, respectively – relaxation processes, structural heterogeneity and morphology of multiphase systems such as crystalline polymers, polyblends and composites. As a technique, DMA is also sensitive to the characterisation of polymers of similar chemical compositions, as well as being able to detect the presence of moderate quantities of additives such as plasticisers or leachable materials. The determination of the ratio of elastic to viscous components within a polymer is an important factor in the understanding of how a material would perform in a given application environment. The potential failure modes such as fatigue, creep and ageing can be related to the viscoelastic behaviour of the polymer.

Dynamic mechanical properties of bone cements

13.6

303

Applications of dynamic mechanical analysis in the characterisation of bone cements

The characterisation of bone cements by DMA has been exploited in several ways. This technique, commonly employed for studying the thermal behaviour of polymeric materials, can provide information regarding the temperature- and time-dependent behaviour in addition to the measurement of dynamic mechanical properties of acrylic bone cements. In terms of their temperature-dependent behaviour, α- (commonly associated with the Tg) and β-transitions have been studied. Closely related to these transitions, the effect of residual monomer has been determined. In terms of their time-dependent behaviour the effect of frequency has been studied and from these measurements information regarding activation energy (ΔE), shift factors (aT), etc. has also been obtained. In this way, activation energies of various transitions (α or β) or relaxation processes observed in different methacrylates can be related to compositional differences in bone cements. On the other hand, the shift factor allows the construction of a composite curve (log E versus time) where the ratio of either relaxation or retardation times can be obtained at different temperatures. Finally, the mechanical response derived from the storage modulus has also been studied by DMA.

13.6.1 Temperature-dependent behaviour of bone cements: determination of the glass transition temperature The temperature at which a polymer undergoes transformation from a rubber to a glass is known as the glass transition temperature, Tg. Several techniques can be used to determine Tg, such as dilatometry, differential scanning calorimetry (DSC), DMA, etc. The advantages of DMA are that the Tg can be measured as the change in slope from a plot of the storage modulus against temperature, from the maxima of the loss modulus peak versus temperature, or from the maxima of tan d versus temperature. When tan d is used for the determination of the Tg, a higher temperature is obtained compared with the loss modulus peak but the transition is easily identified because of the large intensity of the peak. The determination of Tg of bone cements is not included either in the ISO 58336 or ASTM F4517 standards for acrylic bone cements, but it has been determined on various experimental formulations using ASTM D7908 with the argument that this parameter affects their mechanical properties. In general, it has been observed that high-modulus bone cements exhibit high Tg values, while low-modulus bone cements exhibit low Tg values.9

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Orthopaedic bone cements

The Tg of commercial bone cements has been reported to range between 70°C and 100°C depending on their composition when measured with a horizontal dilatomer after 24 h of dry storage.10 However, a lower Tg is observed if the bone cement is conditioned in a wet environment as it has been observed that water or the conditioning media has a plasticising effect, which in turn reduces the Tg by up to 20°C. The primary reason for this is that polymethyl methacrylate (PMMA), even though it is considered a hydrophobic material, absorbs a significant amount of water. The Tg of bone cements can also be modified by the addition of comonomers and crosslinking agents,11–17 although these inclusions also increase water absorption, reduce shrinkage and modify radiopacity. In this manner, it has been observed that the expected Tg of a bone cement based on methyl methacrylate (MMA) (Tg = 100°C) is modified when either diethyl amino ethyl methacrylate (DEAEMA) or methacrylic acid (MAA) is used as comonomer.11 The low Tg observed in DEAEMAcontaining bone cements (Tg = 83.9°C) rests on the fact that the newly formed poly(N,N-diethyl aminoethyl methacrylate) (PDEAEMA) also exhibits a low Tg. On the other hand, the high Tg observed for MAAcontaining bone cements (Tg = 120°C) can be explained due to hydrogen bonding of the acidic units. Furthermore, as these polymers tend to be more hydrophilic, their Tg tends to decrease when the bone cement is conditioned in simulated body fluid due to the higher fluid uptake, which can plasticise the matrix. In a similar way, Cervantes-Uc et al.12 studied the effect of the addition of aliphatic and aromatic comonomers for bone cement preparation. Although the Tg was not substantially modified through the incorporation of aromatic structures such as 4-methacryloyloxybenzoic acid (MBA) or 4-diethyl amino benzyl methacrylate (DEABM), these systems also offer low shrinkage as these monomers exhibit high molar volume. Crosslinking agents such as ethylene glycol dimethacrylate (EGDMA), etoxy triethylene glycol methacrylate (TEG) and poly(ethylene glycol dimethacrylate) (PEGDMA) have been used by Deb et al.13,14 in bone cement formulations. Although the Tg was not measured in these systems by DMA, this parameter is expected to increase in the TEGDMA-containing bone cements as an increase in mechanical properties was observed. Bone cements containing high amounts of ethoxy triethylene glycol methacrylate (TEG), a hydrophilic monomer, were prepared by Pascual et al.15 By using a MKIII dynamic mechanical thermal analyser (DMTA) equipment in a temperature range from 20°C to 200°C at a heating rate of 3°C/min and 1 Hz, a broadening of the Tg was reported, especially in cements with high concentrations of this comonomer. This behaviour was related to the presence of different phases, where phase separation was enhanced when the samples were conditioned in water.

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Artola et al.16 conducted DMA experiments to determine the effect of the addition of radiopaque monomers either in the dry state or after reaching the equilibrium water uptake in distilled water. They found that the Tg decreased from 129°C in the dry state to 116.0°C after conditioning in water when 5% of 2, (2′,3′,5′-triiodo benzoyl) ethyl methacrylate (TIBMA) or 3,5 diiodine salicylic methacrylate (DISMA) were used as radiopacifiers. In a similar study Artola et al.17 reported that the addition of 20% of 4-iodo phenol methacrylate (IPMA) did not shift the Tg significantly, even when the copolymer composition changed. Although the observed Tg was similar to a formulation containing 10 wt% of barium sulphate as radiopaque agent, this temperature was reduced from 129°C to 116°C after the samples were conditioned in a wet environment. Addition of various fillers to bone cements – such as hydroxyapatite (HA), α-tri calcium phosphate (TCP) or Bioglass® – also tends to modify their Tg, although this phenomenon has been related to the presence of unreacted monomer. This behaviour was reported in one of the early works by Vallo18 using bone cements filled with glass spheres. The studies were conducted in three-point bending from 0°C to 150°C at a heating rate of 10°C/min and 1 Hz using a Perkin-Elmer DMA7-e. The results obtained suggested that the higher the monomer content, the lower the Tg due to its plasticising effect. Furthermore, two peaks were detected on the tan d plot, one for the partially polymerised MMA matrix and the second attributed to the prepolymerised PMMA beads. Conversely, Canul-Chuil et al.19 studied the addition of either hydroxyapatite (HA) or α-tricalcium phosphate (TCP) to bone cements containing DEAEMA or MAA and did not observe a change in Tg as the residual monomer increased from 3.5% (unfilled) up to 4.5% (filled). However, their results did show the appearance of two transitions suggesting phase separation or poor mixing.The first transition in formulations prepared with MAA and either HA or α-TCP was attributed to the presence of the polymer beads composed of methyl methacrylate-co-ethyl methacrylate. The second transition did not always appear at the same temperature and was explained by considering that methacrylic acid may undergo dehydration to yield glutaric and succinic polyanhydrides. Two transitions were also detected in DEAEMA bone cements with HA (5 and 10 wt%). The first transition coincides with the Tg of PDEAEMA located at 34°C and the second transition was assigned to the prepolymerised copolymer beads. Yang et al.20 reported that Tg decreased with the addition of TCP and that two transitions appeared in their composites and in the unfilled cement. These authors used up to 70 wt% of the ceramic but it was not stated that the lowering of the Tg was due to the presence of unreacted monomer. In conclusion, Tg measurement can be used as a qualitative indication of the mechanical behaviour of the bone cement, i.e. the higher the Tg, the higher

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the moduli. Furthermore, it has been suggested that bone cements with a Tg between 40°C and 50°C could have a disastrous clinical performance due to their poor load transfer that does not enable bone to be remodelled. However, care must be taken as Tg measurements tend to vary with the technique used for their determination, the geometry employed and the size of the specimens.

13.6.2 Time-dependent behaviour of bone cements In general, the mechanical properties of bone cements are related to their main transition (α, Tg), but materials exhibiting secondary transitions (β) can also offer information regarding their mechanical behaviours under impact. Sometimes, these transitions cannot easily be detected in a bead– matrix composite, such as a bone cement, at a single frequency but experiments at different frequencies make them clearer. In this regard, Vallo et al.21 conducted dynamic mechanical experiments on bone cements at 1, 5 and 10 Hz and found that the presence of a β-transition (located at 10– 25°C) was more evident at low frequencies. The presence of this transition explained the unexpected decrease in flexural and compressive modulus of commercial bone cements when tests were conducted at low deformation rates. Elvira et al.22 prepared experimental bone cements by adding different amounts of 2-hydroxyethyl methacrylate and 5-hydroxy-2-methacrylamido benzoic acid to the MMA liquid phase and observed an increase in the Tg with increasing frequency in the range from 1 to 30 Hz. With these measurements, they were able to calculate the activation energies, associated with the α-relaxation, which were found to be in the range of 210– 287 KJ/mol. The dependence of Tg on frequency was also observed by Cauich-Rodríguez et al.9 in bone cements prepared with butyl methacrylate (BMA) or tetrahydrofurfuryl methacrylate (THFMA) in the frequency range of 0.1–30 Hz. Activation energies of 153 and 235 kJ/mol were obtained for THFMA- and BMA-based bone cements. It has been found that viscoelastic properties depend not only on frequency but also on water uptake and heat treatment.23 Although an increase in storage modulus with frequency was observed in various commercial bone cements (CMW 1GTM, Simplex PTM and BraxelTM), the increase in loss modulus was not the same. Furthermore, the damping factor also increased with frequency, but at body temperature it stayed the same. On the other hand, Boesel et al.24 studied the frequency dependence of the storage modulus of partially biodegradable bone cements when they were immersed in a saline solution at 37°C. As expected, the storage modulus (EI) and the loss modulus (EII) increased with frequency, but the loss factor tended to be independent of this parameter.

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307

13.6.3 Determination of the mechanical properties of bone cements The effects of the addition of a comonomer, fillers (fibre, powder, etc.) and the presence of residual monomer on the mechanical properties of bone cements can also be followed by DMA. In this regard, Pascual et al.15 found that, with increasing amounts of triethylene glycol dimethacrylate (TEG), the storage modulus decreased for an MMA-based formulation. Further reductions in the storage modulus were observed when the samples were conditioned in water. In a similar work, Nien and Chen25 made use of a copolymer of methyl methacrylate–acrylic acid–allyl methacrylate for preparing bone cements as this has the ability to absorb water and control shrinkage. Using DMA they found that storage modulus increased with frequency in the range from 1 to 10 Hz, and that it was higher for the system prepared with the minimum amount of this copolymer. Yang et al.26,27 found that the addition of MMA-g-ultra-high molecular weight polyethylene (UHMWPE) grafted fibre to Simplex PTM radiopaque bone cement increased storage modulus compared with the non-grafted fibres system. This was explained by the formation of a better interphase between the matrix and the fibre. A further advantage of the grafting procedure was that the amount of fibre incorporated into the composite also increased from 2% to 6%. The effect of the presence of residual monomer on the mechanical properties of bone cements has also been discussed by Algers et al.28 They found that residual monomer can be eliminated in PalacosTM after a second heating cycle and that the storage modulus can be increased from 2.3 to 3.0 GPa. DMA has also been used for the long-term mechanical behaviour prediction of PMMA-based bone cements.29 The authors suggested that an analysis based on the time–temperature superposition principle (TTSP) should permit the prediction of viscoelastic properties on a long-term scale, based on short-term measurements at higher temperatures. However, for PMMA this principle was never fully assessed, but extrapolation based on a time power law applied to isothermal experimental data apparently allows much better long-term predictions.

13.7

Future trends

Other dynamic methods such as rheometry, dynamic nanoindentation tests and modulated DSC have been used in addition to DMA to assess the behaviour of bone cements. A Rheometric Ares rheometer in dynamic osscilation mode at a frequency of 1 Hz, and using parallel plate configuration, has been used on bone cements prepared with bismuth salicylate as radiopaque agent.30 For this vertebroplasty bone cement they observed a

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Orthopaedic bone cements

larger G″ (shear loss modulus) after mixing the components and an increase in G′ (shear storage modulus) as the polymerisation progressed. Dynamic nanoindentation tests were performed by Lewis et al.31 using a TriboIndenter applying a small constant oscillatory load of 15 μN superimposed on a quasi-static load of 200 μN from 1 to 200 Hz. They found a similar dynamic mechanical behaviour when three different activators (N,N, dimethyl-4toluidine; N,N dimethyl amino-4-benzyl laureate; N,N dimethyl amino-4benzyl oleate) were used, i.e. an increase in storage modulus with frequency of 4–6 GPa was reported. Dynamic compressisve creep tests conducted on CemexTM IsoplastiTM, CemexTM System, Cemex RXTM and Simplex PTM were performed by Verdonschot and Huiskes.32 They observed a linear relationship between creep deformation and the number of loading cycles when these bone cements were tested immersed in saline solution at 38.5°C. In a similar study Liu et al.33,34 studied CMW 1TM, Palacos R40TM, SmartSetTM and Simplex PTM bone cements under an equivalent stress of 10.6 MPa at 1 Hz and 6 000 000 cycles using a modified Durham MKII hip joint simulator. Two stages of creep were identified with a higher creep rate during early cycling followed by a steady-state creep rate.

13.8

Conclusions

This chapter has highlighted the significance of DMA for the characterisation of polymeric and composite bone cements. The mechanical thermal properties of materials are clearly linked to their structure and can have a significant effect on their clinical performance. DMA can provide such critical information.

13.9

References

1 ferry, j.d. Viscoelastic Properties of Polymers, Third Edition, Wiley, New York, 1980. 2 sperling l.h. Introduction to Physical Polymer Science, Fourth Edition, Wiley Inter-Science, New York, 2006. 3 sepe, m.p. Dynamic Mechanical Analysis for Plastics Engineering, Plastics Design Library, William Andrew Inc., New York, 1998. 4 nazhat, s.n. Thermal analysis of biomaterials, in Principles and Applications of Thermal Analysis, Edited by P. Gabbott (Chapter 7, pp. 256–285), Blackwell Publishing, Oxford, 2007. 5 nielsen, l.e., landel, r.f. Mechanical Properties of Polymers and Composites, Second Edition, Marcel Dekker Inc., New York, 1994. 6 ISO 5833: 2002 Implants for Surgery – acrylic resin cements. International Organization for Standardization, Geneva. 7 ASTM F451-95 Standard specification for acrylic bone cement. Annual Book of ASTM Standards. ASTM, Philadelphia, Pennsylvania.

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8 ASTM D790: Standard Test Methods for Flexural Properties of Unreinforced and Reinforced Plastics and Electrical Materials. Annual Book of ASTM Standards. ASTM, Philadelphia, Pennsylvania. 9 cauich-rodríguez, j.v., vázquez-torres, h., martínez-richa, a. Properties of THFMA-PEMA and BMA-PEMA based bone cements characterized by thermal analysis FTIR and NMR. J. Appl. Biomater. Biomech. 1 (2003) 108–116. 10 kühn, k.d. Bone Cements: Up-to-Date Comparison of Physical and Chemical Properties of Commercial Materials, Springer, Berlin, 2000, pp. 141–148. 11 islas-blancas, m.e., cervantes-uc, j.m., vargas-coronado, r., cauich-rodríguez, j.v., vera-graziano, r., martínez-richa, a. Characterization of bone cements prepared with functionalized methacrylates and hydroxyapatite. J. Biomater. Sci. Polym. Edn 12(8) (2001) 893–910. 12 cervantes-uc, j.m., vázquez-torres, h., cauich-rodríguez, j.v., vázquez-lasa b., san román del barrio, j. Comparative study on the properties of acrylic bone cements prepared with either aliphatic or aromatic functionalized methacrylates. Biomaterials 26 (2005) 4063–4072. 13 deb, s., vázquez, b., bonfield, w. Effect of crosslinking agents on acrylic bone cements based on poly(methylmethacrylate). J. Biomed. Mater. Res. 37 (1997) 465–473. 14 deb, s., braden, m., bonfield, w. Effect of crosslinking agents on poly(ethylmethacrylate) bone cements. J. Mater. Sci. Mater. Med. 8 (1997) 829–833. 15 pascual, b., gurruchaga, m., ginebra, m.p., gil, f.j., planell, j.a., vazquez, b., san roman, j., goñi, i. Modified acrylic bone cement with high amounts of ethoxy triethylene glycol methacrylate. Biomaterials 20 (1999) 453–463. 16 artola, a., gurruchaga, m., vazquez, b., san román, j., goñi, i. Elimination of barium sulphate from acrylic bone cements. Use of two iodine-containing monomers. Biomaterials 24 (2003) 4071–4080. 17 artola, a., goñi, i., gil, f.j., ginebra, m.p., manero, j.m., gurruchaga, m. A radiopaque polymeric matrix for acrylic bone cements. Biomed. Mater. Res. Part B: Appl. Biomater. 64B(1) (2003) 44–55. 18 vallo, c.i. Influence of filler content on static properties of glass reinforce bone cements. J. Biomed. Mater. Res. Appl. Biomater. 53 (2000) 717–727. 19 canul-chuil, a., vargas-coronado, r., cauich-rodríguez, j.v., martínez-richa, a., fernandez, e., nazhat, s.n. Comparative study of bone cements prepared with either HA or α-TCP and functionalized methacrylates. J. Biomed. Mater. Res. Appl. Biomater. 64B(1) (2003) 27–37. 20 yang, j.m., li, h.m., yang, m.c., shih, c.h. Characterization of acrylic bone cement using dynamic mechanical analysis. J. Biomed. Mater. Res. Appl. Biomater. 48(1) (1999) 52–60. 21 vallo, c.i., cuadrado, t.r., frontini, p.m. Mechanical and fracture behaviour evaluation of commercial, acrylic bone cements. Polym. Int. 43 (1997) 260– 268. 22 elvira, c., vazquez, b., san roman, j., levenfeld, b., ginebra, p., gil, x., planell, j.a. Acrylic bone cements incorporating polymeric active components derived from salicylic acid: curing parameters and properties. J. Mater. Sci. Mater. Med. 9 (1998) 679–685.

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23 de santis, r., mollica, f., ambrosio, l., nicolais, l., ronca, d. Dynamic mechanical behaviour of PMMA bone cements in wet environment. J. Mater. Sci. Mater. Med. 14(7) (2003) 583–594. 24 boesel, l.f., mano, j.f., reiss, r.l. Optimization of the formulation and mechanical properties of starch based partially biodegradable bone cements. J. Mater. Sci. Mater. Med. 15(1) (2004) 73–83. 25 nien, y.h., chen, j. Studies of the mechanical and thermal properties of crosslinked poly(methyl methacrylate-acrylic acid-allyl methacrylate) modified bone cement. J. Appl. Polym. Sci. 100(5) (2006) 3727–3732. 26 yang, j.m., huang, p.y., yang, m.c. Effect of ultra high molecular weight polyethylene fiber on mechanical properties of acrylic bone cements. J. Polym. Res. 4(1) (1997) 41–46. 27 yang, j.m., huang, p.y., yang, m.c., lo, s.k. Effect of MMA-g-UHMWPE grafted fiber on mechanical properties of acrylic bone cements. J. Biomed. Mater. Res. Appl. Biomater. 38(4) (1997) 361–369. 28 algers, j., maurer, f.h.j., eldrup, m., wang, j.s. Free volume and mechanical properties of Palacos R bone cement. J. Mater. Sci. Mater. Med. 14(11) (2003) 955–960. 29 guedes, r.m., gomes, m., simoes, j.a. DMTA analysis for the long term mechanical behaviour prediction of PMMA based bone cements. J. Biomater. Sci. Polym. Edn 17(10) (2006) 1173–1189. 30 hernández, l., fernandez, m., collía, f., gurruchaga, m., goñi, i. Preparation of acrylic bone cements for vertebroplasty with bismuth salicylate as radiopaque agent. Biomaterials 27 (2006) 100–107. 31 lewis, g., xu, j., deb, s., vazquez lasa, b., san román, j. Influence of the activator in an acrylic bone cement on an array of cement properties. J. Biomed. Mater. Res. 81A (2007) 544–553. 32 verdonschot, n., huiskes, r. Creep properties of three low temperature curing bone cements: A preclinical assesment. J. Biomed. Mater. Res. Appl. Biomater. 53(5) (2000) 498–504. 33 liu, c., green, s.m., watkins, n.d., gregg, p.j., mccaskie, a.w. Creep behaviour comparison of CMW1 and Palacos R40 clinical bone cements. J. Mater. Sci. Mater. Med. 13(11) (2002) 1021–1028. 34 liu, c.z., green, s.m., watkins, n.d., baker, d., mccaskie, a.w. Dynamic creep and mechanical characteristics of SmartSet GHV bone cements. J. Mater. Sci. Mater. Med. 16(2) (2005) 153–160.

Part IV Enhancing the properties of bone cements

14 Antibiotic-loaded bone cements S. D E B and G. K O L L E R, King’s College London, UK

Abstract: Infection is often a complication following joint replacement surgery and the use of antibiotic-impregnated cements is emerging to be a potentially effective clinical procedure that may assist in reducing the incidence of deep infection. The main advantage of local antibiotic delivery is the ability to achieve high levels of antibiotic at the target site without increasing systemic toxicity whilst providing sustained release over prolonged periods of time. Although antibiotic-loaded bone cements have been in use in Europe for over 30 years, they were approved by the US Food and Drug Administration (FDA) in 2003 and are now being increasingly used as a local drug delivery system. The present chapter examines the literature and discusses the effect of antibiotics on the physico-mechanical properties of cements and their clinical efficacy. Key words: antibiotic-laden bone cement, acrylic bone cement, antibiotics, infection, poly(methylmethacrylate) (PMMA).

14.1

Introduction

Early failure of hip replacements as first described by Charnley, led to the concept of infections related to joint replacement surgery. It was proposed that bacterial contamination of the implant site led to infections and thus caused failure of the joint replacement. Although the current rate of primary infections is low, at rates of less than 2%, few infected joints can be treated without removal of the prosthesis. Complications due to infection of a cemented prosthesis can lead to pain and deterioration of functional use of the implant (Charnley 1972), eventually leading to failure. Untreated infections can lead to severe complications, such as osteomyelitis, and may necessitate limb amputation and in extreme cases may be fatal. An emerging trend in clinical medicine is the use of combination devices such as antimicrobial catheters, drug-eluting stents and functional prosthetic implants. The inclusion of antibiotics within a bone cement matrix in cemented joint replacement surgery is one such technique for reducing infection rates through local delivery of antibiotic formulations (Welch 1978, Trippel 1986). Following joint replacement surgery, antibiotics are frequently administered prophylactically by conventional means, such as intravenous or per-oral routes, to reduce the incidence of such infections. Antibiotics are administered in varying doses, particularly in 313

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immunocompromised patients lower doses are used, whereas higher doses of antibiotics are used when treating an existing infection; however, such an antimicrobial therapy is beset with major disadvantages. One such disadvantage is the potential of adverse drug reactions after systemic administration, such as anaphylaxis in response to antibiotics. Furthermore, high concentrations of antibiotics have also been shown to be nephrotoxic and ototoxic (Hall 1977, Lieberman 2002). Antibiotic efficacy is also reduced in the presence of biomaterial-adherent strains of bacteria, the resistance of which is related to surface-adhesion phenomena (Naylor et al. 1990). Joint replacement sites can acquire infections in many ways, from perioperative primary seeding or contamination, due to non-sterile implant materials, during or after the surgical procedure due to haematogenous spreading, where the primary infection originates at a distant site and is transferred to the implant site. Staphylococcus aureus or Staphylococcus epidermidis are the most commonly encountered bacterial species associated with prosthetic infections, and the traditional method of treating these has been the administration of a systemic dose of antibiotics orally or intravenously, post-surgery. However, poor absorption of drugs and lack of circulation at the cement–bone interface (in cemented total hip replacement (THR)) limit the effectiveness of such systemic antibiotic therapy. The concept of local delivery of drugs to a target site is mainly designed to obtain maximum efficacy with minimal systemic side effects. The primary benefit of local administration of antibiotics is the high efficacy of the drug at the target site without increasing systemic toxicity (Hanssen 2004, Hanssen and Spangehl 2004). Despite improvements in surgical technique and implant design in orthopaedic and trauma surgery, implant-related infections are still a challenging problem for surgeons. Buchholz and Engelbrecht (1970) first incorporated antibiotics into a bone cement almost four decades ago, with the intention of providing a local dose of antibiotic directly to the infection site. Subsequently, the principle of local delivery of drugs in trauma and orthopaedic surgery found widespread introduction in the late 1970s for the treatment of chronic osteomyelitis to achieve high concentrations of antibiotic in situ. One of the most widely used antibiotics in bone cements is gentamicin; this has been in clinical use for over 50 years and exhibits a concentrationdependent antibacterial activity (Lacy et al. 1998). Furthermore, gentamicin in the form of gentamicin sulphate has a number of properties that made it particularly suitable for inclusion in acrylic bone cement formulations, such as wide-spectrum antibacterial activity, solubility in aqueous media, low allergenicity and its thermal stability. The use of antibiotic-impregnated bone cements is considered to be an effective strategy in preventing deep infection following total joint arthroplasty, also underlining its efficacy as a prophylactic agent for deep infection

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rather than as a treatment thereof (Buchholz et al. 1981). Infection continues to be one of the major complications following joint replacement surgery, second only to the serious complication of pulmonary embolism. Currently used surgical solutions to such infections include debridement and irrigation (Brook and Pupparo 1998), excision arthroplasty incorporating one-stage and two-stage (Buchholz 1979) excision arthroplasty without an antibiotic-loaded cement and finally two-stage exchange arthroplasty with an antibiotic-loaded implant. However, all the procedures available for the management of infection include further surgical intervention. While the addition of antibiotics such as gentamicin or tobramycin has been reported to reduce the risk of infection (Buchholz et al. 1981, Hanssen et al. 1994, Scott et al. 1999) and antibiotic-impregnated cements have an impact on the management of infection in total joint replacement surgery, the current literature presents conflicting results (Wahlig and Dingledein 1980, Bourne 2004, Hendriks et al. 2004). The Norwegian Register included 22 170 total hip replacements in a study (Engesaetar et al. 2003) that reported that antibiotic-containing cements reduced the rate of infection from 0.7% to 0.4% (p = 0.001) when comparing groups that received 24 h intravenous antibiotics prophylactically or received both 24 h intravenous treatment and an antibiotic-laden bone cement. Using the cement as a local antibiotic delivery vehicle has, without doubt, evolved to an accepted method of prevention of intra-operative infections. However, concerns continue on the safe use of these cements, especially with regards to the development of antibiotic resistance due to the slow, sub-lethal exposure to antibiotics over a period of time. There are very few reports pertaining to this aspect and recently Langlais and co-workers (2006) reported that the susceptibility of Staphylococci to gentamicin has continued to decrease in some French institutions. A prospective randomized study on 340 knee prostheses (Chiu et al. 2002) used to evaluate the efficacy of cefuroxime-impregnated cements in the prevention of infection after primary knee replacement showed that the cement did not display a discernible effect in the prevention of deep infection but was hypothesized to lead to stronger local resistance to infection due to the elution of the antibiotic into the joint fluid. It was believed that, in combination with oral antibiotics, the local delivery of the antibiotic would be more effective in prevention of infection.

14.2

Demands on acrylic bone cement systems

Generally, a pre-requisite for a viable drug delivery system is that the substrate must have the ability to incorporate the drug, retain it at the site and deliver effectively with time. In the case of an orthopaedic bone cement, it is imperative that it does not adversely affect the curing kinetics or the mechanical properties of the cement, such that there are no long-term

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effects on the longevity of the cement. It was earlier thought that the administration of antibiotics by means of a depot was beneficial prophylactically; however, the advantages of local delivery of the antibiotics were soon exploited for the treatment of infected prostheses. Effective local delivery of antibiotics through elution has the advantage of enhanced efficacy and also decreasing the systemic toxicity. However, there is still some concern regarding antibiotic resistance over time. Acrylic bone cements used in orthopaedics are dispensed in the form of powder and liquid components and the incorporation of an antibiotic is generally achieved via mixing of the antibiotic into the powder phase. The main methods of incorporating antibiotics in the cement composition are either industrial blending of the polymer powder with the antibiotic or hand mixing by the surgeon. The inclusion of antibiotics in bone cements can influence the setting kinetics, and the physical, chemical, as well as mechanical properties of the cement. The cement plays a crucial role in the longterm success of cemented joint replacement; therefore any adverse effect on the properties of the cement resulting from the inclusion of an antibiotic in a cement, needs assessment. Another important consideration in the selection of antibiotics for inclusion in orthopaedic acrylic bone cements is the ability of the antibiotic to withstand temperatures of up to ∼100°C, due to the temperature rise that occurs during the course of the polymerization reaction that occurs in situ during the setting of the bone cement. The polymerization exotherm varies for different cements; however, it does not usually exceed 100°C.

14.3

Antibiotic-loaded bone cements

Most commercial cements containing antibiotics are a modified version of the original plain cements that have been loaded with an antibiotic. Gentamicin is an aminoglycoside antibiotic and can be used to treat a variety of bacterial infections. Gentamicin, like all aminoglycosides, when administered orally is not systemically active due to its poor absorption in the small intestine and is found to be eliminated, largely in the urine. It is usually given intramuscularly, intravenously or topically because the drug binds avidly to certain tissues. Gentamicin is a heat-stable antibiotic that remains active even after autoclaving, which makes it particularly useful in the preparation of bone cements due to the exothermic nature of the polymerization reaction. Tobramycin, another member of the aminoglycoside family of antibiotics and also only effective when administered intramuscularly or intravenously, has also been incorporated in bone cements. Vancomycin is a glycopeptide antibiotic used in the prophylaxis and treatment of infections caused by Gram-positive bacteria. It is generally used to treat infections when other antibiotics have failed, such as in the case of MRSA

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strains now frequently encountered in clinical isolates, although other alternatives are being used due to the emergence of vancomycin-resistant bacterial strains. Some of the other heat-stable antibiotics that have been used in poly(methylmethacrylate) (PMMA) bone cements include clindamycin, erythromycin, oxacillin, cefuroxime and colistin.

14.3.1 Basic composition of the cement The composition of commercial acrylic bone cement systems does not vary vastly from one formulation to another, except that there are various modifications such as the addition of copolymers of PMMA and some different comonomers in the liquid. As has been indicated earlier, the powder phase of bone cements predominantly contains PMMA powder; however, some cements use methylmethacrylate/ethylmethacrylate/styrene copolymers, butylmethacrylate/methylmethacrylate copolymers or methylmethacrylate/ styrene copolymers in varying percentages to enhance handling and properties. The antibiotic is usually incorporated within the powder phase as a pre-blended mixture and in some cements (see Fig. 14.1) the antibiotic is dispensed separately and mixed by the surgeon in the powder phase prior to application. The cement powders also contain radiopacifiers, which are present in plain, unloaded cements as well. The liquid phase remains the same as that found in plain cements, the main variation being the initiator:activator ratio, inhibitor content and pigments. The parameters that can influence the properties and thus performance of bone cements have been discussed elsewhere in detail (Chapter 9); however, the effect of the inclusion of antibiotics on the properties needs to be considered.

14.1 Vacuum mixing using an antibiotic-laden bone cement.

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14.3.2 The effect of antibiotics on curing kinetics of the cement It has been established that the inclusion of additives such as cross-linking agents, modified monomers (Pascual 1999) and radiopaque agents – namely barium sulphate or zirconia – within acrylic bone cements influences the setting kinetics of the cement (Deb and Vazquez 2001). The curing kinetics provide important information for appropriate clinical handling, thus addition of antibiotics and their influence on the working and setting time is important. The setting time and peak temperatures of different cements are sometimes difficult to compare as different protocols have been followed to determine these parameters. However, Kühn (2000) compared the setting time of various plain cements and that of their antibiotic-containing counterparts and reported that no significant changes occurred with the inclusion of the antibiotic. Lewis et al. (2005) reported that the method of blending the antibiotic in a cement powder had no significant effect on the curing kinetics and thus no differences in the thermal stabilities of the cements were found, which is expected, as the blending methods do not influence the composition of the bone cement per se.

14.4

The effect of antibiotics on the mechanical properties of the bone cement

The inclusion of antibiotics within acrylic bone cements may affect the mechanical properties. Thus, numerous studies have focused on this area and, in general, the literature suggests that properties such as compressive and tensile strength and Young’s modulus, are not significantly affected; however, this depends on the amount and type or types of antibiotics incorporated. Kühn (2000) compared the compressive strength of several antibiotic-containing cements and those of their plain counterparts to report a general lowering in compressive strength, but the values were well within acceptable limits as set by the standards. It has also been reported that the amount of antibiotic incorporated influences the compressive strength, which decreases with increasing amount of antibiotic (Lee et al. 1978). Similarly, the bending strength was also not significantly affected but, as expected, the admixing of antibiotic powder caused inhomogeneous mixtures to form, resulting in the lowering of strength and a 10% reduction of the bending strength has been reported. Generalisations regarding the effects of antibiotics on the mechanical properties of bone cement are difficult, not only due to the different cements and antibiotics used but also because of the small sample size of cement used in each study. Klekamp and co-workers (1999) reported the preparation of acrylic bone cements using injectable, lyophilized vancomycin (vancomycin-L) and powdered

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vancomycin (vancomycin-P), and tested the compressive strengths of the cements. The cements Palacos® and Simplex (Stryker, NJ, USA) exhibited no differences in compressive strengths with or without antibiotic, and vacuum mixing inevitably increased the compressive strengths. However, the addition of a second antibiotic (vancomycin) in the cement has been reported to decrease the bending strength to 56 MPa (Palacos R) and 53 MPa (CMW 1), which is only slightly above the ISO recommended minimum value of 50 MPa. This suggests that, although the use of more than one antibiotic offers better prophylaxis, this may affect the mechanical properties adversely. Gentamicin powder in small quantities manifests as discrete domains within a bone cement matrix, while higher concentrations lead to the interconnectivity of the gentamicin domains within the cement mantle and, as the mobility of fluids is greater in gentamicin due to the higher solubility compared with the hydrophobic cement, high concentrations ease the elution of the drug. However, with the progressive elution of the drug, voids may appear, which may cause the onset of fracture. One possible cause attributed to the lowering of mechanical properties is the net lowering of the density due to the lower mass of the drug incorporated and, as density is a measure of the mass per unit volume and hence is a function of porosity, this is therefore a possible mechanism for the inferior mechanical properties. It is known that variations in mixing methods of PMMA powder and monomer can produce cement mixtures with different porosities. The surgeon often mixes antibiotics manually or uses an industrially preblended powder when an antibiotic-laden bone cement needs to be used. However, there are limited data in the literature regarding the issue of the ‘mixing method’ used to blend antibiotics into the cement powder and its impact on the properties. Recently, Lewis and co-workers (2005) have related the method of blending the antibiotic to the PMMA powder to the properties of the cement. The study compared three types of cements with antibiotic incorporated by manual mixing, mixing with a commercially available mixer or using an industrial blending process. They evaluated the modulus, work to fracture in a four-point bend, plane-strain fracture toughness, fatigue life and diffusivity in phosphate-buffered saline. The results indicated that none of the methods had a significant influence on the properties evaluated; however, the authors recommended that manual blending of antibiotic using a mechanical mixer is likely to yield a more reproducible mix than simple manual mixing. A hip joint is subjected to forces as great as three times the bodyweight up to 1 million times each year. Thus the fatigue properties of the cement are important factors in the survival of bone cements in vivo; in fact, they will largely determine if a correctly used cement will fail or not. Numerous studies evaluating the fatigue properties of cements are documented in the

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literature; however, it is difficult to derive conclusive results due to the variation in the method of testing. Nevertheless, the common consensus is that porosity is one key factor. Klekamp and co-workers (1999) reported that the inclusion of vancomycin and tobramycin in commercial cements such as Palacos® had no detrimental effects on the compressive strength; however, it adversely affected the fatigue life. A comparison of vacuummixed Simplex cement, vacuum-mixed Simplex with 2 g of vancomycin-P in 40 g of powder and vacuum-mixed Simplex with 2 g of vancomycin-L in 40 g of powder, showed that the fatigue life decreased by about three times in the cement containing vancomycin-P; however, the fatigue life was about 10-fold less for cements with vancomycin-L. Thus, vancomycin-P had a substantially less detrimental effect on fatigue strength than vancomycin-L, although both are able to maintain their biological activity. Neut et al. (2003) examined the effects of vancomycin on the compressive strength and fatigue life, both alone and in combination with tobramycin. The cements containing the antibiotics did not reveal any significant differences in compressive strengths; however, the fatigue life was reported to decrease significantly in the presence of vancomycin (Hendriks et al. 2005). There is a considerable variation in testing methods, which makes it difficult to achieve a reasonable comparison of the fatigue properties of antibiotic-containing bone cements. Lewis (2002) and Lewis et al. (2005) attempted to compare data on fatigue properties of cements containing antibiotics and highlighted that the different mixing methods, specimen fabrication methods, specimen geometies, test conditions and the frequency at which the testing was carried out make it very difficult to draw reasonable conclusions.

14.4.1 Effect of vacuum mixing Porosity is widely recognized to inf luence the physical and mechanical properties of bone cements, and consequently the life of the arthroplasty. Macro- and micro-pores exist within the cement mantle as a result of: (a) incorporation of air during mixing; (b) loss of volatile monomer due to the steep rise in temperature; and (c) entrapment of air during transfer of mix. Porosity is mostly considered to have an adverse effect on the survival of cemented total hip arthroplasties and has been shown to affect mechanical properties. Pores in cement mantles have been clearly identified as affecting the fatigue properties of cements. Retrieved bone cement samples show that the pores acted as nucleation sites for micro-cracks that could accelerate the crack propagation within the cement mantle. Antibiotic elution occurs predominantly through bulk diffusion; however, the combination of surface roughness, porosity and wettability has also been correlated and greater surface roughness enhances the release of antibiotic in the initial stages. Sustained release of antibiotic occurs due to penetration of water or

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fluid into the cement matrix. A more porous matrix allows greater fluid absorption and therefore the amount of antibiotic leached is higher. The vast majority of in vitro elution tests have been carried out on static samples of antibiotic-loaded cements, whereas in vitro tests show that constant dynamic loading causes development of micro-cracks, which is expected to have a bearing on the release kinetics. These defects in the cement provide more surface area for the release of antibiotic. The concept of vacuum mixing was introduced to reduce the amount of air present in the set cement, in a successful attempt to improve the mechanical properties of the cement. However, the reduction in porosity achieved through vacuum mixing has a deleterious effect on the amount of antibiotic released. The level of antibiotic release is even lower for hand-mixed cements; this is significant in the USA where the majority of antibioticloaded cement used for total hip arthroplasties is hand mixed (Neut et al. 2003). The inclusion of antibiotics within bone cements has an effect on the mechanical properties; however, generalising the effects is not straightforward due to the different cements and antibiotics used.

14.5

Release of antibiotics from bone cements

Antibiotic elution from PMMA-based bone cements may occur predominantly through bulk diffusion; however, surface roughness, porosity and wettability have all been correlated to rate and amount of release. PMMA is a highly hydrophobic polymer and thus impervious to diffusion of drug molecules; however, water uptake studies have shown that diffusion does occur in PMMA cements (1–2% water uptake). Although the mechanism of elution of antibiotics from PMMA cements is not well established, elution does occur and it is highly likely that the combination of diffusion with release from surface and bulk defects such as cracks, imperfections and voids contributes to the elution process. The type of antibiotic incorporated within a bone cement also influences the drug release profile and studies have shown that elution from PMMA is unique to individual antimicrobial agents. Greater surface roughness has also been reported to release more antibiotic in the initial stages. Increasing antimicrobial concentration in PMMA cements increases peak concentrations and area under the curve. For effective functioning of the antibiotic, the amount eluted should be above the minimum inhibiting concentration and the minimum bactericide concentration of the respective pathogens. The amount of antibiotic released from any bone cement is, in general, much lower than the amount incorporated in the matrix. The release pattern of the antibiotic in bone cements is expected to follow the laws of diffusion and depends on factors such as surface area, porosity, wettability and fluid uptake. The polymer powder bead size, shape and distribution are factors that affect several properties

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of acrylic bone cements and these factors may also contribute to the elution rates and amounts of antibiotic from the cement mantle. The wettability as determined by contact angles has been shown to be similar for most bone cement formulations and ranges between 70° and 80°, thus this parameter is not expected to alter the release of antibiotics from different cements. However, surface roughness and porosity are parameters that vary widely from cement to cement and are markedly influenced by the mixing and dispensing technique. For example, the release of gentamicin from different cements – namely CMW 1, CMW 3, CMW Endurance, Palacos and Palamed – was reported to range from 4.0% to 5.3% of the total amount of antibiotic in the first week (van de Belt et al. 2000), which was attributed to the varying degrees of porosity and roughness because the wettability was similar in the different cements. The examination of the microstructure of acrylic bone cement samples fabricated by different techniques – such as hand mixing, vacuum mixing, and vacuum and centrifugation – show specimens with very irregular pores with numerous small craters that may contribute to the rate and total amount of drug released in vivo. The release of antibiotics from bone cements has been researched fairly extensively using both in vitro and in vivo methods. The results produced vary widely, mainly due to the non-uniformity in the studies, thus making it difficult to derive firm conclusions. However, it is important to bear in mind that even a very systematic evaluation of the elution of drugs may not equate to the clinical situation due to the wide variation in the cement mixing and dispensing techniques and insertion of the cement in the femoral cavity. Surgeons can use different pressurization techniques and the clinical variations are expected to lead to cements with varying degrees of porosity, imperfections in the cement mantle and different surface areas, all of which will contribute to the actual elution of the drug into the surrounding tissues. Gentamicin, being one of the most commonly used antibiotic in bone cements, has been the subject of many research papers. It is thus prudent to establish a good, sensitive and reproducible method for quantitative analysis of the eluates for gentamicin obtained from the various in vitro studies. A wide variety of methods to quantify aminoglycoside antibiotics have been reported, including: fluorescence polarization immunoassay, enzyme immunoassay, chromatography and spectrophotometry. Chromatographic methods have been used in most pharmacopoeias for reliable testing; however, these methods are time consuming. Spectrophotometric methods are rapid and reliable, but since gentamicin does not absorb ultraviolet or visible light well, indirect methods are required for assay. Derivatizing agents include ninhydrin as it can react with primary and secondary amines present in the gentamicin molecule that can yield chromophoric products (Frutos et al. 2000). Cabanillas et al. (2000) thoroughly validated

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a method of derivatizing gentamicin sulphate using o-phthaldialdehyde that was reported to be more sensitive than using ninhydrin derivatives, and in addition, the o-phthaldialdehyde–gentamicin complex was more stable than the derivative with ninhydrin. Aminoglycosides have been shown to elute well from bone cement mantles. A comparison of the elution of tobramycin and vancomycin release from vacuum-mixed Simplex indicated that tobramycin eluted 10 times more efficiently than vancomycin. A combination of vancomycin and tobramycin showed that the elution efficiency of tobramycin was reduced in the presence of vancomycin; however, tobramycin had no appreciable effect on the elution rates of vancomycin (Klekamp et al. 1999). Literature reports show contradictory results, and data for the release of gentamicin from identical commercial cements suggest that it may be eluted easily (Holm and Vejlsgaard 1976) and may not be eluted (Elson 1977; Wahlig and Dingledein 1980) from the cement matrix. The contradicting reports in the literature may be attributed to different testing methods or to limited accuracy of the release study, and thus have generated an interest in more standard studies to allow comparison. However, the variable amounts of antibiotic released from the similar cements also depend on the porosity of the cement, which is one factor that can vary quite easily from user to user. van de Belt et al. (2000) studied the influence of release of gentamicin on biofilm formation on different cements. The cements used in the study were CMW 1, CMW 3, CMW Endurance, CMW 2000, Palacos® and Palamed; they were tested for Staphylococcus aureus biofilm formation, and the influence of gentamicin release was quantified in terms of the number of colonyforming units on the gentamicin-loaded cements in comparison with the number of viable organisms on cements without gentamicin. The study clearly showed that, in the early stages, microbial colonization was similar on antibiotic-loaded and unloaded cements; however, for Palacos® and CMW 1 there was a distinct reduction in biofilm formation after 24 hours. There was no clear relation observed between the gentamicin release and biofilm formation as the other cements were more effective either at 48 or 72 hours, which does not relate it to the amount of antibiotic release. However, a degree of caution is required in interpreting the results, as it is evident that the effectiveness of the antibiotic was variable for the different cements. Torrado and co-workers (2001) proposed an in vitro drug diffusion model to elucidate the drug release mechanism of antibiotics from cement mantles. Identical methods of cement mixing were adopted to minimize variation in specimen preparation and two analytical methods – namely, high-performance liquid chromatography (HPLC) and fluorescence polarization immunoassays – were used to quantify the amounts of gentamicin

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eluted from CMW 1 gentamicin cement (DePuy International). A diffusion cell with a buffered solution (pH 7.4) was used to collect the eluants at 37°C. The data obtained showed that large amounts of drug were eluted in the first few hours and that this gradually decreased with time and reached a plateau; however, high levels of dispersion within the sample size were reported. This observation was similar to other trends reported in the literature (Klekamp et al. 1999, Penner et al. 1999). Cerretani et al. (2002) described in vitro elution characteristics of vancomycin combined with imipenem–cilastatin in acrylic bone cements. Three commercial acrylic bone cements namely CMW 1 (DePuy, UK), Palacos R (Schering Plough, Germany) and Simplex P (Howmedica, UK), were used to prepare discs containing 2 g of vancomycin and 2 g each of vancomycin and imipenem–cilastatin per 40 g of cement. Discs were used for the elution tests and it was observed that the amount of vancomycin released in saline increased by 30.58, 50.52 and 50.15% in the presence of imipenem–cilastatin for CMW 1, Palacos R and Simplex P, respectively. CMW 1 showed higher release when treated with vancomycin alone; however, the amounts of vancomycin released were greater from Palacos R and Simplex P in the presence of the additive. In an earlier study (Klekamp et al. 1999), the elution rates of vancomycin and tobramycin from different cements were examined and, subsequently, the effect of combination of antibiotics on elution rates. Vancomycin was reported to be less efficient in terms of elution in comparison with tobramycin from cements such as Simplex P and Palacos R; however, it was claimed that the amount of antibiotic eluted and the rate were independent of the mixing method, which is in contradiction with other studies. The difference in the elution rates between tobramycin and vancomycin was attributed to the higher molecular weight of vancomycin. The presence of vancomycin was reported to decrease the elution of tobramycin; however, tobramycin did not affect the elution rate of vancomycin. With the emergence of organisms resistant to vancomycin and or other glycosides, other drugs such as cefazolin, ciprofloxacin, gatifloxacin, levofloxacin, linezolid and rifampin have now been reported to be suitable for incorporation within bone cements. The release profiles recorded were bimodal with an initial burst followed by a slow sustained release of the drug and, more importantly, the release profiles were dependent on the drug that was used. However, the effects on mechanical properties and polymerization kinetics of the cement are not clearly stated (Anquita-Alonso et al. 2006). The vast majority of in vitro elution tests have been carried out on static samples of antibiotic-loaded cements, whereas in vitro tests show that constant dynamic loading causes development of micro-cracks, which is expected to have a bearing on the release kinetics. These micro-cracks/

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defects in the cement provide more surface area for the release of antibiotic, thus elutions under static conditions may not be truly relective of the in vivo situation. Antibiotic release from bone cements typically follows a biphasic pattern with high initial bursts with a long tail. In order to understand the release pattern, and the factors that influence it, it is important to consider any biomechanical factors that the cement is likely to be subjected to. One of the important considerations in the characterization of bone cements is cyclic loading, thus its effect on the release of antibiotics from cements needs evaluation. As walking results in cyclic loading, it has been suggested that the release of antibiotics be studied under conditions of cyclic loading. Hendriks and co-workers (2003) designed two models of the frontal aspect of a femoral stem, which were cemented with CMW 1 Radiopaque G, Palacos R-G and Palamed G, and cyclically loaded at 5 Hz during the immersion period in water. Their results suggested that, after 10.8 × 106 cycles, initial release of gentamicin from Palamed G was increased significantly for loaded over unloaded specimens, but not from CMW 1 Radiopaque G and Palacos R-G. The results also highlight the fact that normal physiological loading patterns may influence the release of antibiotics; thus this is yet another variable that needs to be taken into consideration. It has also been reported (Hendriks et al. 2003) that freshly prepared cements, when subjected to ultrasound, show an increase in the release of gentamicin, which supports the findings that external stresses have a role to play in the release of drugs. The general consensus among orthopaedic surgeons, attained from 30 years of antibiotic-loaded bone cements use in Europe, is that antibioticloaded cements have certain benefits, but there are some concerns over the long-term presence of the antibiotic. So far, no toxic concentrations of antibiotics have been found in the blood or urine of patients with antibioticloaded bone cements. Although numerous studies have been conducted on the elution mechanism and kinetics of the release of drugs from bone cements, there are still conflicting theories that do not provide a clear concept. It is most likely that such discrepancies will continue to appear in the literature even with further refinements. The facts that commercial bone cements differ in composition, molecular weight of the pre-polymer PMMA, bead size and size distribution, have different additives in terms of radiopacifiers and cross-linking agents, and have different copolymers and small amounts of different comonomers within a cement pack, are enough to cause microstructural differences in the hardened cement. In addition, mixing techniques such as vacuum mixing and the use of closed proprietary systems are expected to influence the presence of defects and voids within the cement. Although some studies claim that vacuum mixing has no effect on the amount and rate of drug release, it is also claimed that the release

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of the drug is predominantly through voids and defects rather than through diffusion. As vacuum mixing decreases the net amount of pores in a cement matrix, it is difficult to conceive that antibiotic elution remains unaffected in a denser matrix. It is the opinion of the author that no, one mechanism occurs in the elution of drugs from cement matrices, as it is a combination of diffusion and loss of molecules through surface defects and bulk irregularities in the cement matrix. The amount of drug present in the cement, the composition of the acrylic cement – i.e. the presence of hydrophobic additives – the blending method of the drug within the uncured cement, the surface area of the cement mantle and porosity are all factors that will influence the elution of drugs. In general, it has been reported that elution is greatest in the early phase and that only a proportion of the impregnated drug is released; this may suggest that any complications due to the prolonged presence of antibiotics in the systemic circulation will be minimal; however, long-term recall studies will reveal a truer picture.

14.5.1 Antifungal derivatives in acrylic bone cements Periprosthetic eukaryotic microbial infections are relatively rare in relation to joint prostheses; some cases have been reported, however, with Candida albicans is the most common causative organism quoted, although C. parapsilosis, C. tropicalis and C. glabrata are also cited (Gaston and Ogaden 2004, Lazzarini et al. 2004). Currently, the treatment of candidal infection involves a two-stage irrigation and suction drainage with or without an antifungal spacer, which is usually carried over a considerable length of time. Intravenous administration of amphotericin B is also given to the patient for a short period of time, followed by long-term antifungal drugs either orally or intravenously. Suggestions have been made on the impregnation of bone cement spacers containing antifungal drugs (Selmon et al. 1998, Marra et al. 2001). Surgery also plays an important role in the treatment of Candida prosthetic joint infections, although treatment is similar to that for native joint infections (Phelan et al. 2000). Prosthesis removal and antifungal treatment are also deemed to be successful procedures in the elimination of fungal infections. Lazzarini et al. (2004) also reported a case of recurring Candida spp. joint infection following prosthetic joint surgery. Although candidal prosthetic infections are not common, they display a devastating effect with risk of immunosuppression and neutropenia (Darouiche 1989). Thus, the incorporation of antifungal agents within bone cement matrices has been advocated to alleviate problems arising from prosthetic fungal infections. The fact that the inclusion of antibiotics in bone cement mixes cause losses in the mechanical properties means that it is important to examine the properties of bone cements containing antifungal drugs. Goss et al. (2007) recently reported the effect of antifungal agents on the mechan-

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14.2 Structure of amphotericin B.

ical properties and elution from two bone cements – namely, Simplex T containing tobramycin and Palamed Gentamicin with gentamicin. Cement samples from Simplex T were prepared by hand mixing under atmospheric pressure with 25, 50, 100 and 200 mg of amphotericin B in 40 g of the bone cement powder. The incorporation of amphotericin B led to an increase in the compressive strength of the cements, contrary to the expected lowering in values; 100 mg of the antifungal agent resulted in a statistically significant rise in compressive strength of about 20% in comparison with the control cement. The authors attributed this observation to the structure of amphotericin B (see Fig. 4.2), which consists of a long unsaturated backbone with reactive double bonds; thus with an initiator–activator present in the mix, the polymerization as propagated by the acrylic monomer may react with the double bonds present in the amphotericin molecule. This is expected to then chemically bind the drug to the matrix, thus generating cross-links in the polymer to improve the mechanical properties. This theory was further confirmed by the lack of elution of the amphotericin B from the matrix due to its chemical linkage with the matrix polymer. However, tobramycin release was observed, albeit to a lesser extent in comparison with the control; this can be attributed to the more cross-linked network in cements containing amphotericin B. This study clearly showed that the compatibility of the drug with the in situ polymerizing mass is an important parameter and any chemical interaction of the drug with the polymerizing mass would lead to compromise of the biologics of the drug and also have an impact on the physico-mechanical properties. Thus, inclusion of drugs such as amphotericin B is not viable in PMMA-based bone cements and this is thus an ineffective method for treatment of fungal infections.

14.6

Other additives in bone cement

The addition of fluoride, growth hormones and bioactive glasses has been attempted in efforts to improve the interface and reduce the risk of prosthetic loosening. Addition of fluoride can achieve a biochemical effect

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wherein bonding with hydroxyapatite in the bone mineral can lead to the formation of fluorapatite, which is more resistant to osteoclastic resorption and, in addition, is also known to promote osteoblast differentiation and proliferation.

14.7

Conclusions

There is a considerable volume of literature on the inclusion of antibiotics in acrylic bone cements and their effect on the physico-mechanical properties, release kinetics and clinical performance. A review of the literature suggests that the most commonly used antibiotic in acrylic bone cement is gentamicin, with other antibiotics such as tobramycin and vancomycin being used either alone or in conjuction with a second antibiotic. There is a lack of detailed systematic studies that would allow a reasonable comparison of the properties of the antibiotic-laden cements due to the variation in the intrinsic factors (such as molecular weight of the PMMA powder, viscosity, amount of antibiotic, type of antibiotic) and extrinsic factors (such as mixing method, testing methods, sample environment); however, the general trend suggests that quasi-static mechanical properties – such as ultimate tensile strength, compressive strength and modulus – are not significantly altered in the presence of the antibiotic, but fatigue properties are generally lowered. A comparison of elution rates from different cements is also difficult, given the fact that there is a lack of systematic evaluation and the different studies use different cements that have been used in different environments with different antibiotics. The general consensus among orthopaedic surgeons, attained from 30 years of antibiotic-loaded bone cements use in Europe, is that antibioticloaded cements have certain benefits; however, some concerns remain over the long-term presence of the antibiotic. No doubt, providing higher concentrations of antibiotics locally is more effective, but the balance between the decrease in systemic toxicity and the potential increase in local toxicity needs to be maintained for safe use.

14.8

References

anquita-alonso p, rouse ms, piper ke, jacofsky dj, osmon dr and patel r, Comparative study of antimicrobial release kinetics from polymethylmethacrylate. Clin Orthop Relat Res 445; 239–244, 2006. bourne rb, Prophylactic use of antibiotic bone cement: an emerging standard – in the affirmative. J Arthroplasty 19; 69–72, 2004. brook dh and pupparo f, Successful salvage of a primary total knee arthroplasty infected with Candida parapsilosis. J Arthroplasty 13; 707, 1998. buchholz hw and engelbrecht h, Depot effects of various antibiotics mixed with Palacos resins. Chirurg 41; 511–515, 1970.

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buchholz hw, Results of exchange operations for infection. In: RA Elson and ADS Caldwell, Editors, Revision Arthroplasty. Proceedings of a Symposium held at Sheffield University, Medical Education Service, Oxford; pp. 103–110, 1979. buchholz hw, elson ra, englebrecht e, lodenkamper h, rottger j and siegel a, Management of deep infection of total hip replacement. J Bone Joint Surg Br 63; 342, 1981. cabanillas pf, pena ed, barrales rienda jm and frutos g, Validation and in vitro characterisation of antibiotic loaded bone cement release. Int J Pharmaceutics 209; 15–26, 2000. cerretani d, giorgi g, fornara p, bocchil, neri l, ceffa r, ghisellini f and ritter ma, The in vitro elution characteristics of vancomycin combined with imipenemcilastatin in acrylic bone-cements: A pharmacokinetic study. J Arthroplasty 17; 619–626, 2002. charnley j, Postoperative infection after total hip replacement with special reference to air contamination in the operating room. Clin Orthop 87; 167–187, 1972. chiu fy, chen cm, lin cfj and lo wh, Cefuroxime – impregnated cement in primary total knee athroplasty: A prospective, randomized study of three hundred and forty knees. J Bone Jt Surg 84; 759–762, 2002. darouiche ro, hamill rj, musher dm, young ej and harris rl, Periprosthetic candidal infections following arthroplasty. Rev Infect Dis 11; 89, 1989. deb s and vazquez b, The effect of cross-linking agents on acrylic bone cements containing radiopacifiers. Biomaterials 22; 2177–2181, 2001. elson ra jephcott ae, mcgeghie db and verettas d, Antibiotic loaded acrylic cement. J Bone Jt Surg 59; 200–205, 1977. engesaetar lb, lie sa, espehaug b, furnes o, vollset se and havelin li, Antibiotic prophylaxis in total hip arthroplasty: effects of antibiotic prophylaxis systemically and in bone cement on the revision of 22,170 primary hip replacements followed 0–14 years in the Norwegian Arthroplasty Register. Acta Orthop Scand 74; 644– 651, 2003. frutos p, torrado s, perez-lorenzo me and frutos g, A validated quantitative colorimetric assay for gentamycin. J Pharm Biomed Anal 21; 1149–1159, 2000. gaston g and ogden j, Candida glabrata periprostheic infection: a case report and literature review. J Arthroplasty 19; 7, 2004. goss b, lutton c, weinrauch p, jabur m, gillett g and crawford r, Elution and mechanical properties of antifungal bone cement. J Arthroplasty 22(6); 902–908, 2007. hall fj, Anaphylaxis after gentamycin. Lancet, 2; 455, 1977. hanssen ad, rand ja and osmon dr, Treatment of the infected total knee arthroplasty with insertion of another prosthesis: the effect of antibiotic impregnated bone cement. Clin Orthop 309; 44, 1994. hanssen ad, Prophylactic use of antibiotic bone cement: an emerging standard – in opposition. J Arthroplasty 19; 73–77, 2004. hanssen ad and spangehl md, Practical applications of antibiotic loaded bone cement for treatment of infected joint replacements. Clin Orthop 427; 79–85, 2004. hendriks jge, ensing gt, van horn jr, lubbers j, van der mei hc and busscher hj, Increased release of gentamicin from acrylic bone cements under influence of low-frequency ultrasound. J Controlled Release 92; 369–374, 2003.

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hendriks jge, van horn jr, van der mei hc and busscher hj, Backgrounds of antibiotic-loaded bone cement and prosthesis-related infection. Biomaterials 25; 545–556, 2004. hendriks jge, neut d, hazenberg jg, verkerke gj, van horn jr, van der mei hc and busscher hj, The influence of cyclic loading on gentamicin release from acrylic bone cements. J Biomechanics 38; 953–957, 2005. holm nj and vejlsgaard r, The in vitro elution of gentamycin sulphate from methylmethacrylate bone cement. A Comparative study. Acta Orthop Scand 47; 144–48, 1976. klekamp j, dawson jm, haas dw, deboer d and christie m, The use of vancomycin and tobramycin in acrylic bone cement: Biomechanical effects and elution kinetics for use in joint arthroplasty. J Arthroplasty 14; 339–346, 1999. kühn kd, Bone Cements: Up-to-date Comparison of Physical and Chemical Properties of Commercial Materials, Springer-Verlag, New York, 2000. lacy mk, nicolau dp, nightingale ch and quintiiliani r, The pharmacodynamics of aminoglycosides. Clin Infect Dis 27; 23–27, 1998. langlais f, belot n, ropars m, thomazeau h, lambotte jc and cathelineau g, Antibiotic cements in articular prostheses: current orthopaedic concepts. Int J Antimicrob Agents 28; 84–89, 2006. lazzarini l, manfrin v and de lalla f, Candidal prosthetic hip infection in a patient with previous candidal septic arthritis. J Arthroplasty 19; 248–252, 2004. lee ajc, ling rs and vangal ss, Some clinically relevant variables affecting the mechanical behaviour of bone cement. Arch Orthop Traumat Surg 92; 1–18, 1978. lewis g, Effect of loading rate on the apparent fracture toughness of acrylic bone cement. Biomed Mater Eng 12(2); 149–155, 2002. lewis g, janna s and bhattaram a, Influence of the method of blending an antibiotic powder with an acrylic bone cement powder on physical, mechanical, and thermal properties of the cured cement. Biomaterials 26; 4317–4325, 2005. lieberman p, Anaphylactic reactions during surgical and medical procedures. J Allergy Clin Immunol 110; S64–S69, 2002. marra f, robbins gm, masri ba, duncan c, wasan km, kwong eh, jewesson pj, Amphotericin B loaded bone cement to treat osteomyelitis caused by Candida albicans. Can J Surg 44; 5, 2001. masri ba, duncan cp and beauchamp cp, Long-term elution of antibiotics from bonecement – An in vivo study using the prosthesis of antibiotic-loaded acrylic cement (PROSTALAC) system. J Arthroplasty 13; 331–338, 1998. naylor pt, myrvik qn and gristina a, Antibiotic resistance of biomaterial-adherent coagulase-negative and coagulase-positive staphylococci. Clin Orthop 261; 126– 133, 1990. neut d, van de belt h, van horn jr, van der mei hc and busscher hj, The effect of mixing on gentamicin release from polymethacrylate bone cements. Acta Orthop Scand 74; 670–676, 2003. pascual b, gurruchaga m, ginebra mp, gil fj, planell ja, vázquez b, san román j and goñi i, Modified acrylic bone cement with high amounts of ethoxytriethyleneglycol methacrylate. B Biomaterials 20; 453–463, 1999. penner mj, duncan cp and bassam am, The in vitro elution characteristics of antibiotic-loaded CMW and Palacos-R bone cements. J Arthroplasty 14; 209–214, 1999.

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phelan dm, osmon dr, keating mr and hanssen ad, Delayed reimplantation arthroplasty for candidal prosthetic joint infection: a report of 4 cases and review of the literature. Clin Infect Dis 34; 930, 2000. scott cp, higham pa and dumbleton jh, Effectiveness of bone cement containing tobramycin. An in vitro susceptibility study of 99 organisms found in infected joint arthroplasty. J Bone J Surg 81-B; 440–443, 1999. selmon gp, slater rn, shepperd ja and wright ep, Successful 1-stage exchange total knee arthroplasty for fungal infection. J Arthroplasty 13; 1, 1998. torrado s, frutos p and frutos g, Gentamycin bone cements: characterisation and release (in vitro and in vivo assays). Int J Pharmaceutics 217, 57–69, 2001. trippel sb, Current concepts review antibiotic-impregnated cement in total joint arthroplasty. J Bone J Surg 68; 1297–1302, 1986. van de belt h, neut d, uges dra, schenk w, van horn jr, van der mei hc and busscher hj, Surface roughness, porosity and wettability of gentamicin-loaded bone cements and their antibiotic release. Biomaterials 21; 1981–1987, 2000. wahlig h and dingledein e, Antibiotics and bone cements. Experimental and clinical long term observations. Acta Orthop Scand 51; 49–56, 1980. welch ab, Antibiotics in acrylic bone cement. In vitro studies. J Biomed Mater Res 12; 679–700, 1978.

15 Modifications of bone cements J. S A N R O M Á N, B. V Á Z Q U E Z L A S A, M. R. AG U I L A R and L. F. B O E S E L, Institute of Polymer Science and Technology (CSIC) and CIBER-BBN, Spain

Abstract: The chapter is dedicated to the analysis and description of the state of the art of bone cement formulations, mainly acrylic selfhardening systems, and the reasons for the development of modifications. The flexibility or stiffness of cured formulations, and the ways in which to control and change these properties, are discussed according to the addition of appropriate compounds that participate in the polymerisation process. Considering that these systems can be applied as bioactive materials or as a controlled delivery system, different approaches with regard to the function and control of the release of bioactive compounds and drugs are considered. Biohybrid composites for bone and cartilage regeneration are of importance because of the implications in the regeneration of tissues and, in connection with this point, future challenges and advances in designs for tissue engineering applications are considered in detail. Key words: acrylic bone cements, self-curing formulations, hydrophilic/ hydrophobic bone and cartilage cements, drug delivery systems, injectable systems for vertebroplasties, scaffolds for bone and cartilage regeneration.

15.1

Introduction

The application of self-curing acrylic bone cements was a pioneering application carried out by Sir John Charnley in the late 1950s in orthopaedic surgery for the fixation and adaptation of joint prostheses, mainly hip prostheses.1,2 Later, the technique was extended to fill bone cavities and for the fixation of other devices applied in the human body. In view of the good results obtained with acrylic resins and acrylic monomers in the formulation of composites for dental applications, they have been applied with a few modifications in many surgical applications in old and young patients, or even children. The success of the process is due to the relative simplicity of the procedure and, in general, to the good biological and biomechani332

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cal behaviour of the system. This is associated with the physical characteristics of the poly(methylmethacrylate) (PMMA)/methylmethacrylate (MMA) formulation, as a consequence of the good solubility of the polymer in the corresponding monomer, which gives rise to the formation of a dough paste that can be manipulated easily during a short, but adequate, period of time. The viscosity of the paste and the time of manipulation (also called working time) can be controlled with certainty by manipulating the composition of the mixture of the polymer and monomer (solid to liquid ratio) as well as by changing the concentration of the initiator (usually benzoyl peroxide (BPO)) and the activator (usually N,N-dimethyl 4-toluidine (DMT)). Many changes to the initial formulation have been proposed over the years, but essentially the mechanism and the main components remain the same. A free radical addition polymerisation of the acrylic monomer in physiological conditions is initiated by a redox mechanism of the peroxide/amine, which occurs on mixing of the two phases of the cement formulation. Commercial products have been rigorously tested both from a chemical and biological point of view, with very successful applications in fields such as orthopaedic surgery, dental fields, tissue engineering and drug delivery systems. Probably one of the most interesting advantages of this kind of system is the versatility, since it is possible to design a wide family of formulations that can be well adapted to different applications; for example, as pure cements of initial formulations for joint prostheses, as fluid and injectable self-hardening formulations for filling of complex cavities, and as drug delivery systems at very different levels. However, despite these recognised characteristics and properties, some negative factors have to be taken into consideration. The most important of these are: (a) the relatively high temperature that is reached at a local level when the polymerising polymer/monomer mixture is applied; (b) the long-term loosening of the prosthesis as a consequence of the absence of secondary fixation, and the migration and inherent toxicity of the acrylic monomer; and (c) the continued presence of the residual initiator/activator redox system in the cement mantle. In particular, aromatic tertiary amines such as DMT are highly toxic and can provoke secondary effects that are of importance in a small number of patients.3 The present chapter is dedicated to the analysis and consideration of the most recent modifications of the current acrylic bone cement formulations which aim to improve the biomechanical behaviour, the biological stability, the bioactivity and the pharmacological activity. The modification can be associated with the addition of active compounds and drugs, either physically or chemically, to one or both components of the current formulations in order to obtain self-hardening acrylic cements for biomedical applications.

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15.2

Modulation of the hydrophilic/hydrophobic character of bone cements and the consequences on the properties and behaviour of such formulations

The traditional acrylic bone cements4 are not true adhesives, but they act as a material to accommodate the prosthesis in the natural and normal position of coupling within the bone cavity. The formulations based on PMMA/MMA are very hydrophobic, with a capacity of hydration of no more than 2 or 3 wt%, but they offer excellent biostability against physiological fluids. One interesting characteristic is the lack of adhesive properties, which, in terms of the coherence of the components of the joint (dynamic and fixed), offers a relative biomimetic approach. The absence of a strong adherence of components gives a certain degree of freedom and capacity for adaptation of the prosthetic components to the natural counterparts, with the most natural solution to the adaptation of the prosthesis to the profile of the edges of the bone cavity. Moreover, as the application of the cement occurs during the hardening process, there is enough time for the best adaptation of the prosthetic components within the cavity, reaching the most natural orientation and position of the components. This fact facilitates the operative surgical process in general with very good success in both the short and long term. The incorporation of additional ingredients in the solid or liquid phases offers excellent opportunities to attain specific properties that can be modulated according to the exact application and function of the prosthesis (see Fig. 15.1). For example, it is recognised that, from a biomechanical point of view, the main function of the bone cement is to transfer stress from the prosthesis to the bone, to reach the optimal load distribution on the prosthesis. An excess of loading at a local point can be fatal in that it results in the consequent failure of the prosthesis.5 The viscoelastic mechanical properties of polymeric formulations of acrylic bone cements are strongly dependent on temperature and on the hydration of the polymeric mass in the physiological medium. These characteristics depend on the microstructure of the polymer chains, and the size and size distribution of microdomains in the hardened materials, which physically is related to the glass transition temperature, Tg, of the polymer. On cooling a specific polymer from the soft state (above the Tg) to a temperature well below the so-called Tg, Brownian motions of chain segments and local bulk side groups are frozen, with subsequent changes in properties from a dough to a brittle material, and other changes that influence the coefficient of thermal expansion and the mechanical or even electric stresses.6 Tg depends on the molecular structure of the monomeric components, the molecular weight of polymeric chains, the water content or even the microdomain distribution

Modifications of bone cements

O

HN

O

O

O

O

O O

COOH

OH

335

O

OH

HO

O

O

O HO

O O

N

15.1 Chemical structures of monomers incorporated in acrylic bone cement formulations to modulate their hydrophilic/hydrophobic character.

in the complex composite system constituted by the polymer particles added to the formulation and the polymer matrix formed by the polymerisation of the monomer (liquid) component. Tg has become an important reference characteristic of bone cements and has been considered for additional characterisation of bone cements.7 Cements with high Tg correspond to brittle materials that have an increased likelihood of breaking with corresponding loosening of the components of the joint prosthesis. The introduction of monomers that lower the Tg of the final cement is an approach that has been targeted by different research groups in order to offer toughness, and therefore better stability in the long term (see Table 15.1). It is necessary to take into consideration the fact that, for applications in the physiological medium, the behaviour of materials depends on the capacity for water absorption, which modifies the Tg noticeably with respect to the Tg of the same material in the dry state. Of course, this depends on the chemical structure of the components that is manifested by the hydrophilic character of the bulk material. For example, the Tg of Palacos R in the dry state is 86°C, but if the cured cement is immersed in water at 37°C for a week, the Tg decreases to 78°C and at 2 and 4 weeks the Tg is 66 and 67°C, respectively. Even if the same cement is immersed in a lipidic medium (10%), the Tg decreases to 76, 73 or 64°C after 1, 2 or 4 weeks, respectively.6

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Table 15.1 Values of glass transition temperature (Tg) of several formulations of modified acrylic bone cements Liquid phase MMA/HEMA MMA/HEMA/5-HMA MMA/HPMA

MMA/HPMA

MMA/MAA MMA/DEAEMA

Tg (°C) 116–117 116–118 99–106 (first peak), 120–126 (second peak) 63–91 (first peak) 120–126 (second peak) 80 65–73

Observations

Reference

DMTA, 1 Hz, dry samples DMTA, 1 Hz, dry samples DMTA, 1 Hz, dry samples

8

DMTA, 1 Hz, wet samples

9

DSC, dry samples DSC, dry samples

10

8 9

10

HEMA, 2-hydroxyethylmethacrylate; 5-HMA, 5-hydroxy-2-methacrylamido benzoic acid; HPMA, 1-hyaroxypropylmethacrylate; DEAEMA, N,Ndiethylaminoethyl methacrylate; DMTA, values obtained by dynamomechanical thermal analysis; DSC, values obtained by differential scanning calorimetry.

The modification of acrylic bone cements by incorporation of a hydrophilic monomer has been attempted by several authors with the aim of compensating for the volume shrinkage due to the increase in swelling, increasing the cement pressure on the prosthesis stem and on the bone.11 The introduction of up to 20% of 2-hydroxyethylmethacrylate (HEMA), modifies the mechanical behaviour and lowers its Tg in the wet state. A slight swelling occurs after sorption depending on the HEMA amount. In tension, the elongation at break increases and strength lowers with a rather evident decay of the elastic modulus. The presence of HEMA alters the kinetics of the reaction, showing a significant increase in the rate of polymerisation.12 Elvira et al.8 reported the introduction of HEMA in conjunction with 5hydroxy-2-methacrylamido benzoic acid (5-HMA), as components of the liquid phase. The incorporation of the acrylic derivative of salicylic acid, 5-HMA, contributed to the formation of intermolecular complexes with salts containing calcium ions as well as to the potential pharmacological effect (anti-inflammatory and analgesic) associated with the salicylic residue of the acrylic components. This modification led to lower exotherms and shorter setting time, giving rise to slightly crosslinked polymeric systems. Incorporation of these two monomers in the classical formulation provides some elastic character, increasing tensile strength and elastic modulus as a consequence of the hydrogen bonding interactions and crosslinking points. However, the tensile strength of wet specimens decreases, but strain to failure and elastic modulus increase (see Table 15.2). The surface characteristics of the new formulations studied by contact angle measurements

Table 15.2 Values of setting time (tsetting), maximum temperature (Tmax) and mechanical properties of acrylic bone cements modified with HEMA and 5-HMA monomers s (MPa) 5-HMA (wt%)

HEMA (wt%)

tsetting (min)

Tmax (°C)

Dry

— — — 2.25 5.00 10.0

9 20 40 9 20 40

5.75 5.25 3.75 6.10 5.15 5.00

92 96 99 89 84 77

50.1 52.1 60.2 51.0 54.8 62.3

Wet ± ± ± ± ± ±

2.0 3.5 4.2 2.5 3.1 3.4

e (%)

Et (GPa)

36.8 40.4 36.6 40.0 31.6 28.6

Dry ± ± ± ± ± ±

0.4 6.5 4.6 3.7 4.6 0.9

2.39 2.50 2.83 2.46 2.79 2.92

Wet ± ± ± ± ± ±

0.60 0.11 0.23 0.10 0.14 0.12

2.42 2.53 2.93 2.52 2.82 3.10

Dry ± ± ± ± ± ±

0.15 0.12 0.31 0.11 0.32 0.30

2.7 2.4 2.0 2.4 2.4 2.3

Wet ± ± ± ± ± ±

0.3 0.2 0.1 0.1 0.1 0.2

3.8 3.9 4.4 4.7 4.7 5.6

± ± ± ± ± ±

0.4 0.8 0.9 0.1 0.2 0.2

338

Orthopaedic bone cements

reveal an increase in the polar character of the surface, which indicates that this modification provides materials of improved biocompatibility. Other researchers studied the introduction of 1-hydroxypropyl methacrylate (HPMA) in the liquid phase of the cement.9 The incorporation of the hydrophilic monomer accelerates the polymerisation kinetics with respect to conventional acrylic bone cement, with an increase of the maximum temperature and a decrease of setting time. Polymerisation shrinkage decreases with the content of HPMA in the cement, and the optimum value is achieved at 50% HPMA, where the shrinkage practically compensates for the swelling. Mechanical properties in compression are acceptable for cements modified with up to 40% HPMA. However, in tension the cement containing 20% HPMA presents values of tensile strength comparable with those of the control, and higher concentrations of the hydrophilic monomer provide materials with a drastic decrease in tensile properties. Analysis of the fracture surface reveals a higher deformation of the matrix because of the introduction of the hydrophilic monomer, showing good adhesion between the beads and the matrix for the cements containing less than 30% HPMA. However, for higher contents of this monomer, the propagating crack goes around the bead along the bead– matrix interface indicating a deterioration of the adhesion between both phases. In order to improve the compatibility between solid and liquid phases, cements can be prepared using an equimolecular MMA/HPMA copolymer as the solid phase; however, a loss of the mechanical properties is obtained as the solid phase is enriched in HPMA, due to the resultant very soft material.13 Following this trend, Pascual et al.14 prepared new formulations of bone cements by substitution with high percentages (up to 60% v/v) of the more hydrophilic monomer, ethoxytriethyleneglycol methacrylate (TEG). The essential advantages of these formulations are the decrease in the polymerisation exotherm and the improvement of ductility. The peak polymerisation temperature decreases linearly with the content of TEG and, when the cement is prepared with 60% v/v TEG, the value of the maximum temperature is approximately 40°C lower than in the control. Setting time shows an exponential increase with the content of TEG, with an increase of 8 min for the cement prepared with 60% v/v. The residual monomer content of the modified cements is comparable with that of the PMMA formulation for TEG contents lower than 50% v/v; however, at higher levels, the residual monomer content increases considerably. Due to the greater volume of TEG, one of the advantages obtained with the introduction of this monomer in the bone cement formulation is the noticeable decrease in the polymerisation shrinkage, which is of great importance since it could improve the contact between prosthesis stem and bone cement, providing a better transfer of loads through the interface. The viscoelastic behaviour of TEG-

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containing cements was studied by means of dynamic mechanical thermal analysis. The results for wet samples, after reaching the equilibrium in physiological conditions, show that storage modulus decreases with the content of TEG and the loss tangent curves become broader, displaying two peaks at high proportions of the hydrophilic monomer; this is attributed to heterogeneities of the bone cement. Static mechanical properties evaluated in tension and compression show that the introduction of TEG changes the brittle behaviour characteristic of the PMMA cement, with an increase in strain to failure but a decrease in strengths and Young’s modulus. The modified cements present a greater ductility with an increase in tensile toughness when the cements are formulated with up to 50% v/v TEG. These results are confirmed by the scanning electron microscopy (SEM) fractographic analysis of the fracture surface of the modified cements showing a higher matrix deformation. The incorporation of functionalised monomers into the liquid phase of acrylic bone cements is documented in the literature.10 The addition of methacrylic acid (MAA) to the liquid phase produces a decrease in setting time and an increase of the maximum temperature reached, during the polymerisation, due to the higher propagation constant and the lower termination constant of MAA in comparison with those of MMA.15 These results are consistent with those reported earlier by Brauer et al.16 on the addition of carboxylic acids to peroxide–amine systems. Accordingly, the addition of MAA in an equimolar concentration to that of the accelerator to compositions containing other high molecular weight monomers, such as 30% dicyclopentenyloxyethyl methacrylate, shows that setting time lowers in 3 min and peak temperature rises in 8°C approximately. The presence of N,N-diethylaminoethyl methacrylate (DEAEMA), on the other hand, decreases the maximum temperature and slows down the curing process. MAA improves the mechanical properties with respect to PMMA cement; however, DEAEMA produces the opposite trend, but the minimum compressive strength (70 MPa) required for bone cements in the international standards (ISO 5833) is fulfilled in all formulations. Further reductions in mechanical properties are also expected because of the water-absorbing capability of functionalised methacrylates. In general terms, low modulus bone cements can be obtained when DEAEMA is present in the formulation, while high modulus bone cements are achieved by the addition of MAA. The effect of the introduction of hydroxyapatite (HA) or tricalcium phosphate (α-TCP), in different proportions to the functionalised formulations has also been investigated.17 The mechanical properties of the filled cements depend mainly on the composition and type of testing, but generally all systems fulfil the minimum compressive strength required for bone cement application, with significantly lower values for the alkaline comonomer-containing systems.

340

Orthopaedic bone cements

15.3

Control of the flexibility or stiffness of acrylic cement formulations

One of the relative limitations of acrylic bone cements is the brittle character of the hardened system after the free radical polymerisation in physiological conditions. This makes the system too stiff to guarantee adequate fatigue strength during the life of the prosthesis in the corresponding application. In order to solve this problem, two different alternatives are used, based on complementary approaches. The first one is based on the addition of a second monomer to the liquid component of the original formulation, with enough reactivity to be incorporated into the growing polymer chains, which contributes to the decrease of the Tg, as has been commented in the previous section. The second method is the addition of micro-/nanoparticles or fibres to the solid phase, which provides an increase in the toughness of the cured formulation with respect to the pure PMMA original system. Several schools, based on the rationale that these components would contribute to an increase in the toughness and fatigue life of the resulting system, have suggested the addition of metal or polymeric fibres. This approach has been considered favourably, not only because it can be applied to cements employed in the fixation of joint prostheses, but also because it contributes to the mechanical properties of self-curing formulations applied in the filling of bone or dental cavities, as well as in the biomechanical stabilisation of bone defects. Metal fibres in the form of short segments or whiskers, if added to the formulation in the appropriate proportion, provide good biomechanical reinforcement and, in addition, they can contribute to decreasing the maximum temperature reached during the polymerisation process.18–20 The addition of 2% v/v short fibres provides a clear increase of the tensile modulus, and contributes to the viscoelastic properties by decreasing the creep of traditional formulations.21,22 The toughness and the fatigue resistance of self-curing formulations based on polyacrylic systems have also been controlled or modulated by the addition of polymeric particles of traditional polymers, such as high-density poly(ethylene) (HDPE), or ultra-high molecular weight poly(ethylene) (UHMWPE). In fact, the addition of microparticles of this polymer family in the right proportion provides a noticeable increase of the toughness of the cured cements that is not dependent on the content of microparticles added, over a wide range of concentrations.20 However, comparative studies on the inclusion of different polyethylene samples show conflicting data in the literature, which can be expected. The inclusion of microparticles or microfibres of UHMWPE and those containing the same quantity of low-density poly(ethylene) (LDPE) is not directly comparable due to differences in the type of the poly(ethylene) particles. According to

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the viscoelastic behaviour of both poly(ethylene)s, it is clear that the LDPE provides conditions for a better dissipation of the mechanical energy than UHMWPE, but if the time of application and periodical charges are applied (simulating a real situation), the result is not the same, and the properties of the LDPE-charged cements change with time faster than those of UHMWPE-charged cements. This can probably be attributed to the poor adhesion of microfibres of LDPE to the acrylic cements, leading to the formation of segregated microdomains during the curing process, resulting in lower toughness and fatigue resistance in comparison with cements containing other poly(ethylene) fibres or microparticles.23 One of the main problems associated with the addition of fillers in the form of fibres or microparticles, is the lack of adhesion to the continuous phase of the cement. This means that a migration of the fibres or microparticles can occur, with the corresponding release of microparticles that can activate the inflammatory processes of the facing joint and, in time, lead to the aseptic loosening of the prosthesis.24–26. According to the classical formulations and the morphology of polyacrylic systems, it can be considered that the PMMA-based cured bone cement, in fact, constitutes a composite of PMMA beads distributed more or less homogeneously in a continuous PMMA matrix. Although the chemical structure of both components is essentially the same, the real system is a homogeneous composite of PMMA particles, in a continuous PMMA matrix. There is a good adhesion of the two components because the smallest PMMA particles are soluble or partially soluble in the monomer MMA, but the larger particles can only swell through diffusion of the monomer in the polymer beads during the mixing phase. In order to improve the adhesion of both components with the other ingredients introduced into the original formulation, the functionalisation of fillers with different kinds of reactants has been proposed.25,27,28 The incorporation of components to improve the adhesion between phases and the adhesion of the cement to the surrounding tissue, has been studied exhaustively over the last 20 years. An elegant approach has been formulated on the basis of the addition of reactive monomers such as 4-methacryloyloxyethyltrimellitate anhydride (4-META), to increase the strength of the cement fixation. The cement consisted of: 4-META and MMA as monomers; tri-n-butylborane (TBB) as an initiator; and PMMA powder as the solid phase.29 The tensile strengths of 4-META bone cement bonded to smoothened stainless steel and titanium alloy stems were 28 and 17 MPa, respectively, whereas those for CMW 3 bone cement were 10 and 7.4 MPa.30 4-META cements have also been reinforced with HA, showing good mechanical performance. The mechanical strength of the reinforced cement was maintained with increasing HA content, which was not the case for the cement prepared in the absence of the adhesive monomer,31 and a

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cohesive failure was observed when fracture surfaces are examined using SEM. The in vivo behaviour of the cement, implanted in dogs, shows evidence of direct attachment and interdigitation of the cement into the bone with no fibrous tissue formation, which can have beneficial effects in stabilisising the prosthesis in the long term.32 Toughening of acrylic cements via the inclusion of beads exhibiting elastomeric properties has also been reported in the literature. Rubber particles with a more or less irregular shape and with adequate size and size distribution, if homogeneously distributed through the entire continuous phase, provide a mechanism of releasing stresses and therefore a corresponding toughening of the cement. This phenomenon is more effective if a good interface between the cement and the rubber particles is attained by the modification of the surface of the particles.33,34 Planell and co-workers35,36 have extensively analysed the influence of the rubber-like particles on the properties of toughness and fatigue response of acrylic cement formulations. The addition of acrylonitrile–butadiene–styrene (ABS) particles changes the fracture behaviour of the cements drastically, which results in a clear ductile surface, as has been shown conclusively by SEM. Biomechanical analysis of the fracture process indicates that the energy of fracture reaches a maximum when 10% ABS particles are added to the liquid component of the original formulation. In a similar manner, a clear influence on the fatigue crack propagation of normalised probes is observed and the addition of 10% ABS particles results in a decrease in the crack propagation rate of one or even two orders of magnitude, despite the fact that this parameter is also dependent on the medium of storage and the preparation conditions.37 We have studied the addition of poly(ε-caprolactone) (PCL) to PMMA beads by suspension polymerisation of MMA in the presence of PCL.37,38 After the free radical polymerisation, the beads obtained show microdomains segregation of PCL in the PMMA beads, due to the incompatibility of PCL and PMMA. This allows the introduction of the PCL as a microdomain in the whole cement formulation and the result is reduced porosity of the cured cement, and easier mixing of the liquid and solid phases, usually for relatively low contents of PCL (lower than 5%). However, owing to the relative limitation of the crystallisation of microdomain PCL phases, and the Tg of this polymer (Tg = −60°C), the higher loading results in a plasticising effect with a decrease in the adhesion of the beads to the polymerised continuous matrix. Thus, a limited amount of PCL can be incorporated and 15% PCL results in optimum mechanical properties. This composition is also optimum for the release of vancomycin, an antibiotic that is added to the initial mixture of liquid/solid components.38 The addition of antibiotics to this relatively complex system is very attractive because the release of the antibiotic can be produced not only as a result of a pure

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diffusion mechanism, but it also profits from the biodegradable and bioresorbable character of PCL, or even of other poly(hydroxyacids) such as poly(glycolic acid) or poly(lactic acid), which in addition offer an initial contribution to the biomechanical properties of the cements. However, the incompatibility of these biodegradable components with the PMMA matrix limits the use of these polymers. Charges of 10–15% of the polyester are the optimum for the preparation of formulations that could be applied as controlled delivery systems, and that would achieve biomechanical coherence with the bone, particularly with the cortical bone component.

15.4

Modification of formulations with bioactive and functionalised components with pharmacological activity

Self-curing acrylic formulations allow the incorporation of new monomers with specific bioactivity or pharmacological action, as well as the addition of drugs such as anti-inflammatory agents, anti-thrombogenic agents, immune suppressors, antibiotics, growth factors, etc. In fact, it is possible to modify the liquid phase, the solid phase or both, using activators of low toxicity, incorporation of crosslinking agents, monomers bearing pharmacologically active compounds, or polymer beads containing bioactive compounds. One of the drawbacks of acrylic cements is the release of toxic components and free radicals several days after the end of polymerisation.39 Free radicals are known to have injurious effects in cells and tissues, and consequently they have been implicated in the pathogenesis of many diseases.40,41 Some studies support the theory that the cytotoxicity of the PMMA bone cement is associated with the release of free radicals and mediated by lipoperoxides and other free radicals, and these phenomena could contribute to the further loosening of the prosthesis. The increase in lipoperoxide production in the presence of the medium exposed to PMMA reflects an oxidative process showing the toxicity of the bone cement, which in turn may be related to the inhibition of bone growth and stimulation of osteolysis.42 In order to overcome the damage produced by the release of free radicals during and after the implantation of the polymerising cement, new formulations with potential antioxidant character have been formulated by incorporation of a methacrylic derivative of vitamin E (MVE) (10–25 wt%) into the liquid phase of the cement.43 Vitamin E is a natural biological antioxidant, which prevents peroxides from accumulating and protects cells from the damaging effects of free radicals. Vitamin E also ensures the stability and integrity of biological membranes.44 The positive effects of the vitamin E-derived monomer are reflected in different

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aspects. The presence of MVE increases both the setting and working times of the modified formulations, which has benefits in the implantation of the reacting mass in the bone cavity. Partial replacement of MMA by MVE provides peak temperature values in the range 62 to 36°C, i.e. values rather lower than 56°C, which corresponds to the onset of coagulation of albumin,45 and values of residual monomer content are comparable with those of PMMA cements. Mechanical properties in tension and compression are also comparable with those of PMMA cements, with a significant increase of Young’s modulus in compression for contents of 15–20 wt% MVE in the cement. The cytocompatibility of cement containing c. 15 wt% MVE increases with respect to that of PMMA. This formulation does not release toxic extracts into the medium and provides a significant increase in cell viability compared with the negative control TMX. Cell viability, proliferation and differentiation indicate that the presence of the vitamin E-containing monomer can improve cytocompatibility with a significant increase in the DNA content at day 3 for the cement containing 15 wt% MVE.43 As stated in a previous section, the incorporation of a methacrylic monomer derived from salicylic acid, 5-HMA,10 to the classical bone cement based on the pure PMMA/MMA system provides an anti-inflammatory and analgesic local activity, and, in addition, the possibility of intermolecular complexes with cationic species such as calcium or even metallic ions.46 Apart from the biological activity found by the addition of 5-HMA to the liquid phase, a decrease of the peak temperature is observed, and with formulations prepared with 10% 5-HMA, values of tensile strength of 62 MPa and elastic modulus of 2.92 GPa are obtained, although the strain to failure decreases by 10% with respect to the classical PMMA/MMA systems. The addition of acrylic monomers with acidic or basic character modifies the parameters of the curing process (peak temperature, working time, viscosity of the polymer/monomer mixture) as well as the properties of the cured cements. An example is the effect of the addition of MAA or DEAEMA as comonomers with MMA. Although the residual monomer content when 15–20% of these monomers are added to the liquid phase of MMA is apparently not affected, with values ranging from 1.5 to 3%, the molecular weight of the cured system, the setting time, the maximum temperature and the Tg of the polymerised cement, are dependent on the composition. In this sense, a faster curing process, higher molecular weight and Tg of the cured samples are clearly observed with the addition of 10–15% MAA to the formulation. In addition, a slight modification of the hydrophobic character of the systems is also observed from the decrease of the contact angle with water. This results in improved behaviour in systems charged with bioactive fillers such as HA or bioactive glasses,

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giving mechanical properties above the standard minimum strength (70 MPa).47 Eugenol (4-allyl-2-methoxyphenol) is a major constituent (70–90%) of clove oil from Eugenia caryophyllata and occurs widely as a component of essential oils. This compound possesses analgesic and anti-inflammatory properties with the ability to relieve pain in irritated or diseased tooth pulp; however, it is not a true local anaesthetic.48 Eugenol exhibits antimicrobial and anti-aggregating activity that has been demonstrated in vitro, along with its influence on platelet function ex vivo.49 It also has a marked antipyretic activity50 when given intravenously and centrally, and may reduce fever primarily through a central action. Anti-anaphylactic properties of eugenol, by preventing mast cell degranulation, have also been reported in the literature.51 In addition, eugenol can prevent lipidic peroxidation in the initial stages due to the presence of the phenolic group, which can scavenge free radicals.52 Recently, we have studied the modification of eugenol by the linking of a methacrylic residue – in order to change the inhibitory character of the original eugenol – to a polymerisable molecule, but keeping the biological properties of the active free eugenol. This approach allows the eugenol derivative to participate in polymerisation reactions rather than to inhibit them, being more efficacious in the field of dental materials. The new derivatives could be incorporated in permanent restorative material, in injectable fluid cement formulations and in bone cements, delivering the bactericide effects of eugenol to the macromolecular chains.53 In this study, the synthetic methods employed explored two different methacrylic derivatives, where the acrylic and eugenol moieties are either directly bonded (e.g. eugenyl methacrylate (EgMA)) or separated through an oxyethylene group (e.g. ethoxyeugenyl methacrylate (EEgMA)). At low conversions, the polymerisation or copolymerisation of EgMA and EEgMA with ethyl methacrylate (EMA), provides soluble polymers consisting of hydrocarbon macromolecules with pendant eugenol moieties. At high conversions, only crosslinked polymers are obtained, this is attributed to the participation of the allylic double bonds of the eugenol in the polymerisation reaction. Analysis of thermal properties reveals a Tg of 95°C for PEgMA and of 20°C for PEEgMA, and an increase in the thermal stability for the eugenolderivative polymers and copolymers with respect to PEMA. Water sorption of the copolymers decreases with the eugenol-derivative content. Both EgMA and EEgMA monomers show antibacterial activity against Streptococcus mutans, producing inhibition halos of 7 and 21 mm, respectively (see Fig. 15.2).53 Investigations of the biocompatibility of the polymerised systems through cell culture studies reveal that the copolymers do not leach any toxic eluants and show good cellular proliferation with respect to PEMA. This study thus indicates that the EgMA derivatives are potentially good candidates for dental and orthopedic cements.

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25

Halo inhibition zone (mm)

20 15 10 5 0

DMSO

Eugenol

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15.2 Results of agar disc diffusion tests of eugenol and eugenolderivative monomers agains Streptococcus mutans CECT 479 after 24 h of incubation at 37°C. DMSO, dimethyl sutphoxide.

15.5

Improvement and modulation of the radiopaque character

Radiopacity is a desirable property in acrylic bone cements, to allow postoperative assessment of the implant using X-radiography. Most acrylic bone cements are rendered radiopaque by the addition of heavy metal salts of barium or zirconium as a contrast medium.54 However, the addition of radiopacifying compounds such as barium sulphate (BaSO4) or zirconium dioxide (ZrO2) affects different properties of the cement. In general terms, it has been recognised that the lack of adhesion between the inorganic particles and the PMMA matrix is the main reason for the detrimental effects of the radiopaque agent on mechanical properties. It has been demonstrated that BaSO4 reduces tensile strength considerably55 and this reduction seems to be lower for ZrO2.54 Reductions in the flexural strength with the addition of radiopacifying agents are also reported,56 although the influence of BaSO4 on the fracture toughness of the cement is rather controversial.57, 58 From a biological point of view, the release of radiopacifier particles into the surrounding tissues can have detrimental effects,59 ZrO2 being more harmful than BaSO4 due to its abrasive properties, and the addition of radiopaque agents to PMMA may contribute to the bone resorption and aseptic loosening.60 BaSO4 has also been shown to intensify the release of inflammatory mediators in response to PMMA particles.61 Taking all these factors into account, the investigation on this subject has been directed towards using radiopaque agents that are more compatible with the organic matrix. One approach has been the development of radiopaque monomers62,63 bearing covalently bound halogen atoms, i.e. bromine

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or iodine, which can be copolymerised with acrylic monomers producing a homogeneous radiopaque cement. Davy et al.64 prepared polymer beads based on copolymers of methacrylic monomers containing the group triiodobenzoate, and used them as the solid phase to formulate cold-curing systems with good mechanical properties. Other cements prepared from radiopaque polymer beads are based on an equimolar copolymer of MMA and 2-[4-iodobenzoyl]-oxo-ethyl-methacrylate (4-IEMA).65 These cements have intrinsic mechanical behaviour in compression tests superior to that of BaSO4-containing cement and the fatigue life of vacuum-mixed iodinecontaining cement is significantly better than that of the conventional cement. Results of in vitro biological evaluation of this novel radiopaque cement do not differ from those of translucent PMMA or BaSO4-containing cement and the in vivo experiments show no inflammation or foreign body reaction near the iodine-containing bone cement.66 Further studies on the mechanical, thermal and physical properties of the 4-IEMA-containing cement led to the conclusion that this formulation was a viable alternative for use in cemented arthroplasties in place of the current formulation.67 We have employed an iodine-containing methacrylate in the liquid phase of the acrylic bone cement. Thus, cements prepared with 5–7.5 wt% of 2,5-diiodo8-quinolyl methacrylate (IHQM)68 with respect to the liquid phase present sufficient radiopacity, have a lowered exotherm and provide a statistically significant increase in the tensile strength, fracture toughness and ductility with respect to the BaSO4-containing cement.69 However, the fatigue crack propagation resistance of these cements remains similar to that of the radiolucent one,70 indicating that the reinforcing effect that IHQM produces in static mechanical properties is not as effective as to hinder the stable crack propagation. The biocompatibility of the iodine-containing cements is good, showing neither chronic inflammatory response nor macrophages in the area of the implanted rods of cements in rat.68 Following this approach, Artola et al.71 have formulated radiopaque acrylic cements with different amounts (5–20% v/v) of 4-iodophenyl methacrylate (IPMA)71 in the liquid phase and the addition of 15% v/v IPMA is enough to attain a radiopacity similar to that of a 10 wt% BaSO4-containing cement. These iodine-containing cements present enhanced compressive and tensile strengths, elastic modulus and strain to failure with respect to conventional radiopaque bone cements. Other iodine-containing monomers that have been employed in the preparation of radiopaque acrylic bone cements are 2-[2′, 3′, 5′-triiodobenzoyl] ethyl methacrylate (TIBMA) and 3,5diiodosalicylic metacrylate (DISMA).72 The mechanical evaluation of the resulting cements showed an increase in the compressive strength and elastic modulus in comparison with BaSO4-containing cements. The cement prepared with 5 wt% DISMA showed the highest value for compressive strength in dry specimens although for wet samples the opposite behaviour

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was observed, probably due to the higher water uptake of this cement bearing carboxylic groups. Another approach to obtaining more homogeneous radiopaque bone cements has been the incorporation of triphenyl bismuth (TPB) as the radiopacifying agent. This compound is known to render PMMA polymers radiopaque73,74 and it is noticeably insensitive to moisture, which avoids the compound being leached into the aqueous environment. TPB has been investigated as a potential radiopaque agent for acrylic bone cements75 by following two routes. One of these is to blend TPB powder with the solid phase of the cement CMW 1 and the other consists of dissolving different amounts of TPB in MMA prior to cement formation. The latter route provides cements with unaltered exotherms but improved mechanical properties, which is attributed to the solution of TPB into the polymer matrix during polymerisation of the corresponding monomer, in which the compound is up to 70% soluble, and, therefore, leads to the formation of a homogeneous and continuous matrix with a much lower porosity. The addition of TPB via the dissolution method provides a statistically significant increase in the strain to failure in comparison with commercial acrylic cements containing BaSO4, thus reducing the brittleness of the cement. In addition the detrimental effects on the mechanical properties following conditioning in water are also much less pronounced in the homogeneous TPB cements. As stated above, one of the reasons for the detrimental effect of the radiopaque agent on mechanical and biological properties is the lack of adhesion between the particles and the polymeric matrix. One solution for improving interface adhesion consists of establishing covalent chemical bonding between both materials by using a silane coupling agent susceptible to reacting with the oxide surface of the inorganic particles and, further on, copolymerising with organic monomers. This approach has been reported by Behiri et al.76 in the reinforcement of the poly(ethylmethacrylate)-based cements with 3-(trimethoxysilyl)propyl methacrylate (γ-MPS)-treated HA particles, giving cements with enhanced tensile modulus and yield strength compared with untreated fillers. Abboud et al.77 proposed the use of alumina particles previously treated with γ-MPS to prepare radiopaque acrylic bone cements, where the grafted γ-MPS molecules would act as radiopacifying and reinforcing agents. The effectiveness of the surface modification of ceramic oxide particles such as Al2O3, TiO2 and ZrO2 through grafting of γ-MPS was previously reported.78 The mechanical characterisation of the so-prepared cements reveals an improvement in compression with values of modulus of the order of 3400 MPa; however, the formulations require high processability. New acrylic bone cements have been formulated with γ-MPS-modified alumina particles embedded in poly(methyl methacrylateco-ethyl acrylate) beads with about 7 mol.% of ethyl acrylate repeating

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units.79 The compressive strength of the cement cured with the hybrid beads decreases with alumina content whereas compressive modulus remains roughly constant. These results are in contradiction to those reported for cements based on a mixture of γ-MPS-treated alumina and unfilled beads, and they are interpreted in terms of alumina arrangement in the cement. Nevertheless, covalent bonding established between the acrylic matrix and the alumina fillers is expected to reduce the production of abrasive ceramic debris in tissue around prosthetic joints with the subsequent improvement of the biological properties.

15.6

New biohybrid composites for bone and cartilage regeneration

Bioactive material systems are those that provide a good bonding with the tissues in contact with the material, promoting the direct growth of tissue on the surface with a corresponding increase in the interfacial strength. In this regard, the addition of bioactive fillers based on calcium phosphates, HA, TCP, bioactive glasses and other calcium salts such as wollastonite or wollastonite/TCP has been extensively studied. In addition to the bioactivity, the incorporation of these types of fillers to the solid phase may contribute effectively to the increase in the biomechanical properties of the cured cements.80 On the other hand, bioactive compounds based on activators of cell proliferation and regeneration of tissue (bone, cartilage, vascular, etc.) seem to be very interesting for the stimulation of regenerative processes and the integration of the cement in the tissue. This is particularly important for the treatment of vertebroplasties by the injection of fluid cement formulations. Human growth hormone (HGH) stimulates bone and cartilage regeneration if applied in appropriate doses over an adequate period of time.81 Acrylic cements based on the classical PMMA/MMA formulation, charged with a relatively low concentration of GH, released only 1% of the charged drug in the first 24 hours; however, the bioactivity was retained, which was observed from the high concentration of the HGH in the surrounding tissues in comparison with normal physiological levels, and this was detectable after 36 days of implantation of the bone cement formulation.82 Self-curing formulations with a relatively hydrophilic character based on bioactive glasses as the inorganic component and bis-GMA (glycidyl methacrylate of bis-phenol A) and TEGDMA (triethyleneglycol dimethacrylate) as the polymerisable components have been extensively studied by Kokubo et al.83–85 These formulations present very good adhesion to the bone tissue, with improved strength, and the dissolution of the bioactive glass provides an excellent biointegration, with a clear growth of neo-formed bone tissue in the domains occupied by the bioactive glass particles. There is a positive

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balance of properties and biointegration, and as the cement mixture is fluid during the working time, these formulations can be applied by injection in the filling of bone cavities as well as being used for the fixation of prostheses.86 Wollastonite (calcium silicate, natural or synthetic) has been applied as a bioactive filler in bone tissues. This compound, when applied alone or in combination with calcium phosphate salts, noticeably activates bone regeneration and provides a very active surface for formulations with high bioceramic content.82,86,87,88,89 We have recently studied the properties and the in vitro behaviour of bioactive composites of wollastonite and a bioresorbable acrylic copolymer of ethyl methacrylate/vinyl pyrrolidone (EMA/ VP).90 The hydrophilic character of the composites depends on the composition of the copolymer and the content of the bioactive filler, wollastonite. Composites were initially prepared by bulk polymerisation of the mixture of the components at 50°C using azobisisobutyronitrile as free radical initiator. Chemical characterisation, compressive strength, flexural strength, degradation, bioactivity and biocompatibility were evaluated in specimens with a 40/60 EMA/VP ratio and a ceramic content in the range 0–60%.91 A good integration between phases was achieved. Greater compressive strength than the plain copolymer specimens was obtained only when the ceramic load was up to 60% of the total weight, whereas the maximum flexural strength was achieved for the composition with a 45% ceramic load. In vitro biocompatibility studies showed the absence of cytotoxicity for all formulations. The cells were able to adhere similarly on the TMX control and on the formulation containing 60% wollastonite, forming a monolayer and showing a normal morphology.

15.7

Future directions in the design and development of cements with specific properties

Although self-curing formulations based on the polymerisation of acrylic and vinyl monomers have been applied for more than 50 years, they still offer very interesting opportunities in the development of new strategies for regeneration of tissues. The original objective of bringing in a system for the accommodation of joint prostheses has changed to the concept of designing and applying bioactive formulations that can be injected without the risk of a surgical operation. The use of monomers with bioactive moieties is a biomimetic approach that allows the targeting of specific activators, growth factors, antioxidants and antibiotics, just at the site of the implant, with a local, specific action according to the needs of the patient. It is possible to prepare bioactive composites using a biomimetic approach and these composites can then be injected with all the ingredients to mimic

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as closely as possible the characteristics of integration and regeneration. The addition of drugs to suppress the immune response, antibiotics to control infection, anti-inflammatory agents to modulate inflammation processes and growth factors to stimulate tissue regeneration, are some of the objectives that are being considered as future directions for research in this area. However, it is necessary to consider that the addition of these specific ingredients cannot be done without establishing the adequate concentration of the corresponding drug. It is well known, for example, that the application of growth hormone can have negative effects if its concentration is not adequate, and this factor is variable and can vary from patient to patient. The same occurs with the application of bone morphogenetic proteins BMP-2 and BMP7: although the activity in regenerative bone processes is clear, if the concentration is very low, there is no effect and, in contrast, if the concentration is very high, secondary effects occur and bone of poor quality is formed in many cases. The application of acrylic formulations in tissue engineering is of particular importance if regenerative processes and treatments are applied for in situ regeneration. These systems allow the design and targeted application of the necessary components for the stimulation of a regenerative process. The systems can be designed with biodegradable or resorbable, hydrophobic or hydrophilic components, and charged with bioactive agents, in a nano-scale dimension, and the future for regeneration of tissues and targeted local therapies could be brilliant with the close cooperation of specialists in different fields, from the biomaterials world to those with expertise in biological and surgical procedures. The combination of elements from the inorganic field, and biodegradable or biostable but biocompatible polymers, offers unlimited options and possibilities, and eventually will result in the best biomimetic way to reproduce the natural regenerative processes as closely as possible.

15.8

References

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poly(ε-caprolactone) beads. J Biomed Mater Res Part B: Appl Biomater 2004; 70B: 340–347. moreau mf, chappard d, lesourd m, montheard jp, basle mf. Free radicals and side products released during methyl methacrylate polymerisation are cytotoxic for osteoblastic cells. J Biomed Mater Res 1998; 40: 124–131. kehrer jp. Free radicals as mediators of tissue injury and disease. Crit Rev Toxicol 1993; 23: 21–48. poli g, parola m. Oxidative damage and fibrogenesis. Free Rad Res Commun 1977; 22: 287–305. martins vale f, castro m, monteiro j, couto fs, pinto r, toscano rico jmg. Acrylic bone cement induces the production of free radicals by cultured human fibroblasts. Biomaterials 1997; 18: 1133–1135. méndez ja, aguilar mr, abraham ga, vázquez b, dalby m, disilvio l, san román j. New acrylic bone cements conjugated to vitamin E: Curing parameters, properties and biocompatibility. J Biomed Mater Res 2002: 62: 299–307. lucy ja, dingle jt. Fat-soluble vitamins in biological membranes. Nature 1964; 204: 156–160. planell ja, vila mm, gil fj, driessens fcm. Acrylic bone cements. In Wise DL, Trantolo DJ, Altobelli DE, Yaszemski MJ, Gresser JD and Schwartz ER, Editors. Encyclopaedic Handbook of Biomaterials and Bioengineering. Part B: Applications, Vol 2. 1995, Marcel Dekker, New York: pp. 879–921. rivas bl, moreno-villoslada i, elvira c, san román j. Analysis of the interactions of biologically active poly(methacrylic aminosalicylic acid) supports with Ca2+ and Zn2+ by ultrafiltration. J Memb Sci 2001; 192: 187–191. espigares i, elvira c, mano j, vázquez b, san román j, reis rl. New partially degradable and bioactive acrylic bone cements based on starch blends and ceramic fillers. Biomaterials 2002; 23: 1883–1875. sticht fd, smith pm. Eugenol: some pharmacologic observations. J Dent Res 1971; 50: 1531–1535. laekeman gm, van hoof l, haemers a, vanden berghe da, herman ag, vlietinck a. Eugenol a valuable compound for in vitro experimental research and worthwhile for further in vivo investigation. J Phytother Res 1990; 4: 90–96. feng j, lipton jm. Eugenol: antipyretic activity in rabbits. Neuropharmacology 1987; 26: 1775–1778. kim hm, lee eh, kim cy, chung jg, kim sh, lim jp, shim ty. Antianaphylactic properties of eugenol. Pharmacol Res 1997; 36: 475–480. okada n, satoh k, atsumi t, tajima m, ishihara m, sugita y, yokoe i, sakagami h, fujisawa s. Radical modulating activity and cytotoxic activity of synthesized. Eugenol-related compounds. Anticancer Res 2000; 20: 2955–2960. rojo l, vázquez b, parra j, lopez bravo a, deb s, san román j. From natural products to polymeric derivatives of eugenol. A new approach for preparation of dental composites and orthopaedic bone cements. Biomacromolecules 2006; 7: 2751–2761. rudigier j, kirschner p, richter ie, schweikert ch. Influence of different X-ray contrast materials on structure and strength of bone cements. In Hastings GW and Williams DF, Editors. Mechanical Properties of Biomaterials. 1980, Wiley & Sons, New York: p. 289.

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55 vázquez b, deb s, bonfield w. Optimisation of benzoyl peroxide concentration in an experimental bone cement based on poly(methyl methacrylate). J Mater Sci Mater Med 1997; 8: 455–460. 56 holm nj. The modulus of elasticity and flexural strength of some acrylic bone cements. Acta Orthop Scand 1977; 48: 436–440. 57 sih gc, berman at. Fracture toughness concept applied to methyl methacrylate. J Biomed Mater Res 1980; 14: 311–316. 58 molino ln, topoleski ldt. Effect of BaSO4 on the fatigue crack propagation rate of PMMA bone cement. J Biomed Mater Res 1996; 31: 131–137. 59 caravia l, dowson d, fisher j, jobbins b. The influence of bone and bone cement debris on counterface roughness in sliding wear tests of ultra-high molecular weight polyethylene on stainless steel. Proc Instn Mech Engrs, Part H, J Engng Medicine 1990; 204: 65–70. 60 sabokbar a, fujikawa y, murray dw, athanasou na. Radiopaque agents in bone cement increase bone resorption. J Bone Joint Surg 1997; 79B: 129– 134. 61 lazarus md, cuckler jm, schumacher hr, ducheyne p, baker dg. Comparison to the inflammatory response to particulate polymethylmethacrylate debris with and without barium sulphate. J Orthop Res 1994; 12: 532–541. 62 jayakrishnan a, chithambara thanoo b. Synthesis and polymerisation of some iodine-containing monomers for biomedical applications. J Appl Polym Sci 1992; 44: 743–748. 63 kruft mab, benzina a, bar f, van der veen fh, bastiaansen cwm, blezer r, lindhout t, koole lh. Studies on two new radiopaque polymeric biomaterials. J Biomed Mater Res 1994;28:1259–1266. 64 davy kwm, anseau mr, odlyha m, foster cm. X-ray opaque methacrylate polymers for biomedical applications. Polym Int 1997; 43: 143–154. 65 van hooy-corstjens csj, govaert le, spoelstra ab, bulstra sk, wetzels mr, koole lh. Mechanical behaviour of a new acrylic radiopaque iodine-containing bone cement. Biomaterials 2004; 25: 2657–2667. 66 van hooy-corstjens csj, bulstra sk, knetsch lw, geusens p, kuijer r, koole lh. Biocompatibility of a new radiopaque iodine-containing acrylic bone. J Biomed Mater Res Part B: Appl Biomater 2007; 80B: 339–344. 67 lewis g, van hooy-corstjens csj, bhattaram a, koolle lh. Influence of the radiopacifier in an acrylic bone cement on its mechanical, thermal and physical properties: Barium sulphate-containing cement versus iodine-containing cement. J Biomed Mater Res Part B: Appl Biomater 2004; 73B: 77–87. 68 vázquez b, ginebra mp, gil fj, planell ja, lopez-bravo a, san román j. Radiopaque acrylic cements prepared with a new acrylic derivative of iodoquinoline. Biomaterials 1999; 20: 2047–2053. 69 ginebra mp, aparicio c, albuixech e, fernandez-barragan e, gil fj, planell ja, morejon l, vázquez b, san román j. Improvement of the mechanical properties of acrylic bone cements by substitution of the radio-opaque agent. J Mater Sci Mater Med 1999; 10: 733–737. 70 ginebra mp, albuixech l, fernández-barragán e, aparicio c, gil fj, san román j, vázquez b, planell ja. Mechanical performance of acrylic one cements containing different radiopacifying agents. Biomaterials 2002; 23: 1873–1882.

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71 artola a, goñi i, gil j, ginebra p, manero jm, gurruchaga m. A radiopaque polymeric matrix for acrylic bone cements. J Biomed Mater Res Part B: Appl Biomater 2003; 64B: 44–55. 72 artola a, gurruchaga m, vázquez b, san román j, goñi i. Elimination of barium sulphate from acrylic bone cements. Use of two iodine-containing monomers. Biomaterials 2003; 24: 4071–4080. 73 delaviz y, zhang z-x, cabasso i, smid j. Homogeneous radiopaque polymers with organobismuth compounds. J Appl Polym Sci 1990; 40: 835–843. 74 rawls hr, granier rj, smid j, cabasso i. Thermomechanical investigation of poly(methylmethacrylate) containing an organobismuth radiopacifying additive. J Biomed Mater Res 1996; 31: 339–343. 75 deb s, abdulghani s, behiri jc. Radiopacity in bone cements using an organobismuth compound. Biomaterials 2002; 23: 3387–3393. 76 behiri jc, braden m, khorosani d, wiwattanadate d, bonfield w. Advanced bone cement for long term orthopaedic implantations. Bioceramics 1991; 4: 301–307. 77 abboud m, vol s, duguet e, fontanille m. PMMA-based composite materials with reactive ceramic fillers. Part III: Radiopacifying particle-reinforced bone cements. J Mater Sci Mater Med 2000; 11: 295–300. 78 abboud m, turner m, duguet e, fontanille m. PMMA-based composite materials with reactive ceramic fillers. Part 1. Chemical modification and characterisation of ceramic particles. J Mater Chem 1997; 7: 1527–1532. 79 abboud m, casaubieilh l, morvan f, fontanille m, duguet e. PMMA-based composite materials with reactive ceramic fillers: IV. Radiopacifying particles embedded in PMMA beads for acrylic bone cements. J Biomed Mater Res 2000; 53: 728–736. 80 vallo ci. Residual monomer content in bone cements based on PMMA. J Biomed Mater Res 2000; 53: 717–727. 81 goodwin cj, braden m, downes s, marshall j. Investigation into the release of bioactive recombinant human growth hormone from normal and low-viscosity poly(methylmethacrylate) bone cements. J Biomed Mater Res 1997; 34: 47–55. 82 tamura j, kawanabe k, kobayashi m, nakamura t, kokubo t, yoshihara s, shibuya t. Bioactive bone cement: The effect of amounts of glass–ceramic powder. J Biomed Mater Res 1996; 30: 85–94. 83 kawanabe k, tamura j, yamamuro t, nakamura, kokubo t, yoshihara s. A new bioactive bone cement consisting of BIS-GMA resin and bioactive glass powder. J Appl Biomater 1993; 4: 135–141. 84 tamura j, kawanabe k, yamamuro t, nakamura t, kokubo t, yoshihara s, shibuya t. Bioactive bone cement: The effect of amounts of glass powder and histologic changes with time. J Biomed Mater Res 1995; 29: 551–559. 85 senaha y, nakamura t, tamura j, kawanabe k, lida h, yamamuro t. Intercalary replacement of canine femora using a new bioactive bone cement. J Bone Joint Surg 1996; 78B: 26–31. 86 dufrane d, delloye c, mckay ij, de aza pn, de aza s, schneider yj, anseau m. Indirect cytotoxicity evaluation of pseudowollastonite. J Mater Sci Mater Med 2003; 14: 33–38. 87 tamura j, kitsugi t, lida h, fujita h, nakamura t, kokubo t, yoshihara s. Bone bonding ability of bioactive bone cements. Clin Orthop 1997; 343: 183–191.

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88 de aza pn, guitian f, de aza s. Bioeutectic: a new ceramic material for human bone replacement. Biomaterials 1997; 18: 1285–1291. 89 garcía carrodeguas r, de aza a, de aza pn, baudín c, jimenez j, lópez-bravo a, pena p, de aza s. Assessment of natural and synthetic wollastonite as source for bioceramics preparation. J Biomed Mater Res 2007; 83A: 484–495. 90 rodríguez-lorenzo lm, garcía-carrodeguas r, rodríguez ma, de aza s, parra j, san román j. Development of wollastonite-poly(ethylmethacrylate-covinylpyrrolidone) based materials for multifunctional devices. J Biomed Mater Res A 2007; 81A: 603–610. 91 rodríguez-lorenzo lm, garcía-carrodeguas r, rodríguez ma, de aza s, jiménez j, lópez-bravo a, fernández m, san román j. Synthesis, characterization, bioactivity and biocompatibility of nanostructured materials based on the wollastonite-poly(ethylmethacrylate-co-vinylpyrrolidone) system. J Biomed Mater Res 2008; Doi: 10.1002/JBM.a.31867

16 Design of bioactive bone cement based on organic–inorganic hybrids T. M I YA Z A K I, Kyushu Institute of Technology, Japan; and C. O H T S U K I, Nagoya University, Japan

Abstract: Problems with poly(methylmethacrylate) (PMMA) bone cement arise owing to loosening between the bone and the cement after a long implantation period, due to a lack of bone-bonding ability, i.e. bioactivity. Bone-like apatite formation in the body environment is needed for an artificial material to show bioactivity. The silanol (Si–OH) group and calcium ions are effective components for inducing apatite nucleation. Organic–inorganic hybrids prepared by organic modification of these components have attracted much attention as novel bonerepairing materials having both bioactivity and flexibility like natural bone. We attempted to prepare bioactive PMMA cement by chemical modification based on the organic–inorganic hybrids. Specifically, we added alkoxysilane and water-soluble calcium salts to the cements. The cement formed an apatite layer in simulated body fluid, which reproduces the body environment well, when the type of calcium salt was appropriately selected. Key words: poly(methylmethacrylate) (PMMA) bone cement, bioactivity, apatite, organic–inorganic hybrid, simulated body fluid.

16.1

The need for bioactive bone cements

Joints are essential tissues for smooth locomotion. However, they can be severely damaged by accidents and diseases, including rheumatism and osteoarthritis. Artificial joints, such as hip and knee joints, have been commonly used in order to restore the function of the damaged joints. Two types of artificial joints are currently adopted. One is the cemented type of artificial joint, wherein a self-setting poly(methylmethacrylate) (PMMA)based bone cement is used to fix the joint within the surrounding bone material.1 The other type of joint in popular use is the cementless artificial joint system.1,2 This type of joint is fixed by the anchoring of macroporous surface structures of the implant to the host bone. Some cementless artificial joints are modified with bone-bonding bioactive layers on their surfaces, by means such as hydroxyapatite thin film coatings and surface chemical treatments. The choice of which system is used, the case selection criteria, is primarily based on the age and the condition of the disease of the recipient patient. 358

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It is widely accepted that the clinical success of such joint replacement is related to the achievement and maintenance of a strong bond, directly or indirectly, with the patient’s bone. Significant problems of PMMA bone cement include loosening at the interface between the bone and the cement caused by lack of tissue bonding, strongly underlining the requirement for bioactive bone cements to achieve long-term stability of the PMMA cements within the body.

16.2

How do materials exhibit bioactivity?

How can we obtain bioactive PMMA bone cements? In order to design them, it is important to understand the the bone-bonding mechanism of artificial materials within the body environment. Artificial materials implanted into bone defects are generally encapsulated with a fibrous tissue, mainly composed of collagen; this is a normal reaction to protect the living body from foreign substances. Owing to this phenomena, the implanted material is isolated from the surrounding bone and does not bond to the living bone.This type of biological response is classified as the implant being a ‘bioinert’ type. In contrast, Hench et al. developed a novel glass in the early 1970s, commonly referred to as Bioglass® or bioactive glass, containing Na2Oˆ CaOˆSiO2ˆP2O5,3–5 with a composition of (mass%): Na2O, 24.5; CaO, 24.5; SiO2, 45; and P2O5, 6. Bioactive glass or Bioglass® has attracted a significant amount of clinical and research interest, as it possesses the attractive characteristic of direct bone-bonding in the body. This type of bone-bonding ability is called ‘bioactivity’. The glass of this composition shows the highest rate of bone-bonding among the melt-derived Na2OˆCaOˆSiO2ˆP2O5 glass systems. Bioglass® shows so high a biological affinity that it bonds not only to hard tissues but also to soft tissues. The high biological affinity of Bioglass® expands its clinical usefulness to a wide range of applications, including artificial middle ear implants and fillers for alveolar ridge reconstruction. Unfortunately, the mechanical strength of Bioglass® is lower than cortical bone, restricting clinical applications to non-load-bearing applications. This has resulted in the development of bioactive glass–ceramics in order to improve the mechanical properties of bioactive glasses. Kokubo et al.6,7 synthesized a bioactive apatite–wollastonite glass–ceramic (A–W), that has oxyfluorapatite (Ca10(PO4)6(O,F2)) and β-wollastonite crystals within an MgOˆCaOˆSiO2 glassy matrix by heat treatment of glass with composition of (mass%): MgO, 4.6; CaO, 44.7; SiO2, 34.0; P2O5, 16.2; and CaF2, 0.5. This glass–ceramic has β-wollastonite (CaO·SiO2) that consists of silicate chains of an SiO4 tetrahedral structure to increase its mechanical strength, as well as oxyfluorapatite. Specifically, glass–ceramic A–W is a kind of ceramic–ceramic hybrid that contains 38 mass% oxyfluorapatite

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and 34 mass% β-wollastonite grains in the size range of approximately 50–100 nm. Glass–ceramic A–W shows a bending strength of approximately 200 MPa in air, which is higher than that of human cortical bone. Glass–ceramic A–W specimens, when implanted in rabbit tibia, have been shown to bond tightly to the living bone and, on application of a perpendicular load, the fracture is seen to occur within the glass–ceramic rather than at the bone–implant interface. Glass–ceramic A–W has such a level of bioactivity and suitable mechanical properties that it has been clinically used as artificial iliac crests, vertebrae and bone fillers from 1991 to 2000, with a total of more than 60 000 placed implants. Hydroxyapatite (Ca10(PO4)6(OH)2) is known as the main inorganic component of natural bone, constituting 70 mass% of the bone matrix. Synthetic hydroxyapatite prepared by sintering processes also exhibits bioactivity,8 and was commercialized in the late 1980s as bone substitutes in the form of dense and porous body implants and granules. According to previous reports examining the interface between the bioactive ceramics and the bone in vivo, bonding to living bone takes place through an apatite layer, which is formed at the material interface by a sequential chemical reaction with body fluid, as shown in Fig. 16.1.9,10 It is

Bone tissue

Body fluid

Carbonate-containing hydroxyapatite with lattice defects

Bioactive ceramics

16.1 Bone-bonding mechanism of bioactive ceramics.

Design of bioactive bone cement based on organic–inorganic hybrids

361

important to note that some tricalcium phosphate (TCP) ceramics exhibit bioactivity without formation of such an apatite layer.11 The apatite layer formed is composed of carbonate hydroxyapatite of similar structure and composition to that of natural bone. Therefore, it is expected that osteoblasts would preferentially proliferate on the apatite layer and differentiate to form an extracellular matrix composed of biological apatite and collagen, resulting in direct contact of the bone with the surface of the materials, without interstitial fibrous tissue forming.12 When this bioactive phenomenon occurs, a tight chemical bond is formed between the surface apatite and the bone apatite in order to decrease the interface energy between them. The bone-bonding ability therefore depends on the rate of apatite formation when the materials are exposed to the body environment. The apatite formation on the bioactive ceramics and glass–ceramics can be well reproduced even in a simulated body fluid (SBF), with an inorganic ionic composition similar to that of human blood plasma, as shown in Fig. 16.2,13–16 proving that the apatite formation is caused by chemical reaction between their surface and surrounding fluid, as SBF does not contain any cells and proteins. Materials that form this apatite layer on their surfaces in SBF have been shown to possess the potential to bond to living bone. Furthermore, SBF can be used not only for evaluation of the bioactivity of artificial materials in vitro, but also for apatite coating on various materials under biomimetic conditions. Ion concentrations of SBF are shown in Fig. 16.2. The pH of SBF is adjusted to pH 7.25 at 36.5°C and the preparation protocol for SBF is as described in the literature.16 When the apatite-

Specimen

Ion

Concentration (mol m−3) Human blood plasma SBF

Na+

142.0

142.0

K+

5.0

5.0

Mg2+

1.5

1.5

Ca2+

2.5

2.5

147.8

103.0

HCO3−

4.2

27.0

HPO42−

1.0

1.0

SO42−

0.5

0.5

CI−

Buffered with HCI and 50 mol m−3 tris(hydroxymethyl) aminomethane

16.2 Ion concentrations of SBF in comparison with those of human blood plasma.

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forming ability of the specimen is not so high, the pH of the SBF is sometimes adjusted to pH 7.40. This fluid is a metastable solution containing calcium and phosphate ions and is already supersaturated with respect to the apatite. Although SBF is able to mimic chemical reactions of materials with body fluid in vitro well, there is still a gap in composition between SBF and human extracellular fluid, i.e. SBF contains larger amounts of Cl− ions and smaller amounts of HCO3− ions. Therefore, the apatite that is formed in SBF also contains larger amounts of Cl− ions and smaller amounts of HCO3− ions than natural bone. A novel SBF with inorganic ion concentrations exactly equal to those of human extracellular fluid has been proposed by revising the preparation protocol, including changing the type and amount of chemical reagents.17 We can see that bioactive materials can be designed by providing a material with apatite-forming ability. The capability for apatite formation is dependent on the composition and structure of the materials. For instance, the rate of apatite formation increases in the order: sintered hydroxyapatite < glass–ceramic A–W < Bioglass®.18–20 Sintered hydroxyapatite also forms a bone-like apatite phase when it is exposed to the body environment, through an ion-exchange reaction with the surrounding body fluid;21,22 however, this takes place at a low rate. In contrast, glass–ceramic A–W contains a glassy phase in the CaOˆSiO2 system, as well as crystalline oxyfluorapatite and β-wollastonite. This glassy phase contributes to rapid apatite deposition. What composition can act as basic component for bioactive ceramics? Ohtsuki et al.23 fundamentally investigated the compositional dependence of apatite formation on the surfaces of glasses in the CaOˆSiO2ˆP2O5 system after soaking in SBF. It is noted that surface apatite formation is restricted to the compositional regions of the CaOˆSiO2 system with SiO2 ranging from 30 to 70 mol%, but not the CaOˆP2O5 system. Both the calcium ions released from glasses in the CaOˆSiO2 system and the phosphate ions released from those in the CaOˆP2O5 system increased almost equally the degree of the supersaturation of the surrounding fluid with respect to apatite. In spite of that, the glasses from the CaOˆSiO2 system form the surface apatite layer, but not those from the CaOˆP2O5 system. This indicates that the surfaces of the glasses in the CaOˆSiO2 system exclusively provide favourable sites for apatite nucleation by formation of Si–OH and release of Ca2+. The glasses form a silica hydrogel layer before forming the apatite layer and during this process an appreciable amount of silicate ions are dissolved. This means that highly hydrated silica, implying silanol (Si–OH) groups, are abundant on the surfaces of the glasses. It is speculated that these silanol groups effectively induce heterogeneous nucleation of the apatite. This is confirmed by the observation that a pure silica gel, prepared by a sol–gel

Design of bioactive bone cement based on organic–inorganic hybrids Body fluid HCO3−

2+ HPO42− Ca

K+ SO42−

CI− Mg2+

HO

Si

O

O

O

O

Apatite OH

OH O

OH−

Na+

Ca2+ 2+ PO 3− − 4 OH Ca OH OH Ca2+ Si

363

Si

O

O

2+ O Ca

Si

O

Ca2+ increases degree of supersaturation with respect to apatite Si–OH group provides apatite–nucleation sites

O

CaO–SiO2 glass

16.3 Mechanism of apatite formation on CaO–SiO2 glass in the body environment.

method, forms bone-like apatite on its surface in SBF at pH 7.40,15,24 suggesting that a certain kind of silanol group is responsible for the apatite deposition on bioactive materials. The released Ca2+ ions increase the degree of the supersaturation of the surrounding fluid with respect to the apatite, which is already supersaturated even before the exposure of the glass– ceramics. Once apatite nuclei are formed on the surface of the materials, they can grow spontaneously by consuming calcium and phosphate ions from the surrounding body fluid. These mechanisms are shown schematically in Fig. 16.3. Addition of small amounts of Al2O3 and TiO2 to the glass from the CaO–SiO2 system significantly suppresses the apatite formation in SBF.25,26 This is because release of Ca2+ from the glasses into the surrounding fluid is reduced, and because the apatite-forming ability of hydrated silica containing Al2O3 and TiO2 is thought to be lower than that of pure hydrated silica formed on CaO–SiO2 glasses. What kind of chemical structures besides silicate compounds are effective in inducing bioactivity? In order to clarify this point, apatite deposition on various materials, including oxide gels prepared by the sol–gel process, selfassembled monolayers (SAMs) and chemically treated metal substrates have been investigated for their interaction with SBF. The study revealed that not only the Si–OH group but also various functional groups such as Ti–OH,24,27,28 Zr–OH,29,30 Ta–OH,31,32 Nb–OH,33 COOH,34–36 PO4H234 and SO3H37,38 are effective in triggering heterogeneous nucleation of the apatite. Uchida et al.27 reported that titania (TiO2), having an anatase structure, exhibits a higher ability of apatite formation in SBF than materials having rutile and amorphous structures. This means that not only the type of functional group, but also the crystalline structure, governs the apatite-forming ability of materials.

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The initial stage of apatite formation on CaO–SiO2-based bioactive glasses and sintered hydroxyapatite has been precisely investigated using techniques such as transmission electron microscopy (TEM), energydispersive X-ray (EDX) spectroscopy and zeta potential measurements.21,22,39 The results have revealed that the apatite formation progresses through the following processes. First, Ca2+ ions in SBF are selectively adsorbed onto negatively charged surfaces. Then phosphate ions are adsorbed on the calcium-rich surfaces to form an amorphous calcium phosphate. Finally, the resulting amorphous calcium phosphate is converted into low-crystalline apatite, since hydroxyapatite is the most insoluble in water among calcium phosphates in neutral conditions. Thus, we can summarize that selective adsorption of Ca2+ occurs in SBF onto negatively charged surfaces of materials and is a key issue in inducing apatite formation.

16.3

Design of bioactive bone cements using bioactive ceramics

On the basis of the results of bone-bonding mechanisms of bioactive ceramics, bioactive PMMA cements have been extensively investigated through replacing a part of the PMMA powder used by the bioactive ceramic fillers.40 Moroni and co-workers41 investigated the effect of hydroxyapatite powder addition to PMMA bone cements on the properties of the cements. Creep resistance was shown to be improved and the maximum temperature due to radical polymerization of MMA monomer was reduced by the addition of 5% (w/w) hydroxyapatite.41 Shinzato et al.42 compared the bonebonding properties of PMMA bone cements with various bioactive ceramics added, including hydroxyapatite, MgO–CaO–SiO2–P2O5–CaF2 glass and glass–ceramics A–W. The cements were prepared by mixing the bioactive ceramics with PMMA cement at 70 mass% to the total of the cement. They evaluated bone tissue affinity by means of the ratio of the length of bone tissue in direct contact with the cement surface to the total length of the cement, providing an affinity index. All the cements showed direct bonebonding after the implantation in rabbit tibia. Shinzato et al. also reported that the cements with MgO–CaO–SiO2–P2O5–CaF2 glass added showed a higher affinity index than those added with hydroxyapatite and glass– ceramic A–W. It has recently been found that anatase-type TiO2 shows high bioactivity; and, on the basis of this finding, Goto et al.43 prepared PMMA cements supplemented with nano-sized anatase powder. They reported that the addition of 50% or more by mass of the anatase powder to the cement was effective in exhibiting bioactivity. Furthermore, bioactivity of the cement was improved when the anatase powders were pretreated with methacyloxypropyltrimethoxysilane (MPS; H2C:CCH3COO(CH2)3Si(OCH3)3).

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This type of cement does not release calcium ions, which cause degradation of the cements. Therefore, the mechanical properties of the cements are expected to not deteriorate over a long period in the body.

16.4

Bioactive organic–inorganic hybrids

On the basis of the apatite formation mechanism on bioactive glasses and glass–ceramics, hybrid materials with bone-bonding ability can be developed by organic modification of Si–OH and Ca2+ at nanometric level. This type of material design would alleviate the problems currently encountered with bioactive ceramics. Generally, the fracture toughness of ceramics is lower and the Young’s modulus is higher than those of natural bone. For this reason, bioactive ceramics are difficult to be used under load-bearing conditions. Natural bone itself is a kind of organic–inorganic hybrid, where apatite nanocrystals and collagen fibers are three-dimensionally fabricated. Therefore, bioactive organic–inorganic hybrids are expected to have mechanical properties more analogous to those of natural bone. The molecular design of bioactive organic–inorganic hybrids is schematically illustrated in Fig. 16.4. In order to develop such organic–inorganic hybrids, limited heat treatment at lower temperature is permitted during the synthesis, since organic polymers are liable to be easily decomposed by the treatment. Sol–gel processing, which enables synthesis of various ceramics and oxide gels at lower temperature, is one of the attractive methods for preparing organic– inorganic hybrids. It has already been reported by Hu and Mackenzie44 that organically modified silicates can be synthesized from tetraethoxysilane (TEOS) with

Body fluid P(V)

Ca (II)

O H

Ca (II)

Apatite

Apatite

H

H

O

O

O Si

H

O−

−

Ca2+

O

Si

O

Si

Organic polymer

16.4 Design of bioactive organic–inorganic hybrids. Ca(II), divalent calcium ion; P(V), pentavalent phosphorus.

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incorporation of poly(dimethylsiloxane) (PDMS). The synthesized organic– inorganic hybrid, called an ORMOSIL, is expected to have properties of thermal stability from the silica block as well as flexibility from the polymer chain. Osaka et al. paid attention to the fact that ORMOSILs contain substantial amounts of Si–OH groups and reported that incorporation of calcium nitrate in ORMOSILs during the sol–gel processing gave the hybrids apatite-forming ability in SBF.45 Si–OH groups and Ca2+ ions on ORMOSILs trigger nucleation of the apatite in SBF. This implied that the OROMSILs incorporated with calcium salt show bone-bonding properties when implanted in the body. Chen et al.46 modified the protocol of the synthesis and obtained TEOS–PDMS–Ca(NO3)2 hybrids with higher mechanical strength. Synthesis of bioactive organic–inorganic hybrids has also been reported using vinyltrimethoxysilane (VS; H2C:CHSi(OCH3)3) and MPS as starting materials.47,48 Vinyl and methacryloyl groups in the starting organic substances were first radically polymerized to form organic polymers. Then the polymers were incorporated with Ca2+ ions by mixing with calcium salts and subjected to hydrolysis with water to form Si–OH groups. The synthesized organic–inorganic hybrids formed an apatite layer on their surfaces in SBF. We attempted the synthesis of bioactive organic–inorganic hybrids by incorporation of MPS (CH2:C(CH3)COO(CH2)3Si(OCH3)3) and calcium chloride into 2-hydroxyethylmethacrylate (HEMA; CH2: C(CH3)COO(CH2)2OH).49 HEMA has high hydrophilicity and high biological affinity, and its polymer is used for medical applications such as contact lenses and coating agents on artificial blood vessels.50 MPS and HEMA were dissolved in ethanol at a molar ratio of MPS :HEMA of 1 :9, at a total concentration of 1 mol/dm3. A solution of 100 cm3 was heated at 75°C for 3 h with 0.001 mol benzoyl peroxide (BPO) as the initiator for polymerization of HEMA and MPS. Then the polymer solution obtained was mixed with 20 cm3 of ethanol solution containing 0.01 mol CaCl2. Some of the solutions were mixed with 1 cm3 of either 1 mol/dm3 HCl or NH3 aqueous solution as a catalyst for hydrolysis. The resultant solutions were cast in polypropylene containers and dried at room temperature. After gelation, the gels were further dried under ambient conditions until the weight loss of the sample became less than 2% in 24 h. Hybrids without and with addition of HCl and NH3 were denoted as ‘NO’, ‘HC’ and ‘NH’, respectively. The tensile mechanical properties of the hybrids were evaluated using a universal testing machine under ambient conditions according to Japanese Industrial Standard (JIS) K7113. Nine specimens were subjected to tensile testing for NO, and seven specimens for HC and NH. The bioactivity of the obtained hybrids was evaluated by examining the apatite formation on their surfaces in SBF. Figure 16.5 shows representative stress–strain curves for the MPS–HEMA hybrids. Their tensile strength increased in the order: HC (0.15 ± 0.02 MPa)

367

Design of bioactive bone cement based on organic–inorganic hybrids

NH

Stress (MPa)

2

1 NO HC 0 0

20

40

60

80

Strain (%)

16.5 Representative stress–strain curves of the MPS–HEMA hybrids prepared with different catalysts. NO

HC

10 μm

NH

10 μm

10 μm

16.6 SEM photographs of the MPS–HEMA hybrids prepared with different catalysts after soaking in SBF for 7 days.

< NO (0.68 ± 0.04 MPa) < NH (2.10 ± 0.18 MPa). The Young’s modulus of the examined materials also increased in the same order: HC (0.24 ± 0.02 MPa) < NO (2.65 ± 0.19 MPa) < NH (41.1 ± 1.8 MPa). The Young’s moduli of the hybrids were shown to be similar to those of human cancellous bone (50–500 MPa) and articular cartilage (1–10 MPa). Such hybrids are therefore expected to provide novel materials for reconstruction of cancellous bone and articular cartilage. Scanning electron microscopy (SEM) micrographs of hybrids after soaking in SBF for 7 days are shown in Fig. 16.6. After the soaking, spherical particles were formed on the surfaces of NO and NH hybrids, but not the HC hybrid. The thin-film X-ray diffraction (TF-XRD) patterns in Fig. 16.7 gave peaks assigned to poorly crystalline apatite at about 26° and 32° for hybrids NO and NH after the soaking, but not for hybrid HC. Our results indicate that the hybrids NO and NH are expected to form apatite in the body and bond to living bone. In contrast, hybrid HC does not show bioactivity, in spite of the existence of silanol groups and calcium ion release. Release of HCl in hybrid HC would decrease the degree of supersaturation of the surrounding fluid with respect to apatite, and conse-

368

Orthopaedic bone cements

Intensity (arbitrary units)

Apatite

NH

HC

NO

20

30

40

2q (degrees)

16.7 TF-XRD patterns of the surfaces of the MPS–HEMA hybrids prepared with different catalysts after soaking in SBF for 7 days.

quently suppress apatite formation. There also remains a possibility that silanol groups did not have the appropriate structure with the addition of HCl, because the structural effect of silica gel on the apatite formation ability is also affected by the fabrication process of the silica gel.51,52 After the polymerization of MPS and HEMA solutions, a copolymer consisting of MPS and HEMA was produced. The alkoxysilane groups in the copolymer were hydrolysed by water from the atmosphere to form silanol groups during the ageing and drying process.53,54 The silanol groups then condensed to form siloxane bonds (˜Si–O–Si˜), providing crosslinkage among the MPS–HEMA copolymer. When HCl was added to the copolymer solution, it was expected that linear siloxane chains would be mainly formed due to the polycondensation of silanol groups.54 In contrast, when NH3 was added to the copolymer solution, it was expected that threedimensional siloxane networks would be predominantly formed. Therefore, hybrid NH is the hardest, as a result of the high concentration of siloxane networks, in comparison with hybrids NO and HC. The lower Young’s modulus of hybrid HC might be attributed to the lower concentration of the siloxane network structure in the hybrids. Various natural and synthetic polymers are available for synthesis of the hybrids. Recent researches have reported the preparation of bioactive organic–inorganic hybrids, including chitin and polycaprolactone, by modification with Si–OH and Ca2+.55,56 On the basis of these findings, Kamitakahara et al.57 synthesized bioactive anatase-poly(tetramethylene oxide) (PTMO) hybrids. Anatase crystals about 10 nm in size precipitated after hot water treatment of the organic–inorganic hybrids obtained from PTMO and titanium tetraisopropoxide (Ti(OCH(CH3)2)4). The synthesized

369

Design of bioactive bone cement based on organic–inorganic hybrids

hybrids form apatite in SBF within 3 days. It is noted that this type of hybrid forms apatite, even if it does not contain Ca2+. This means that calcium-free bioactive organic–inorganic hybrids can be designed by appropriately controlling the chemical state of the metal hydroxides. According to similar molecular design criteria, bioactive tantalum oxide (Ta2O5)–PTMO hybrids can be obtained.58 Such calcium-free bioactive hybrids are expected to maintain their strength in the body environment, since the release of Ca2+ from the hybrids is reported to decrease their strength in the body environment.

16.5

Design of bioactive bone cements based on organic–inorganic hybrids

Organic modification of Si–OH and Ca2+ was applied to the development of bioactive materials that can be injected into bony defects. In our laboratory, we attempted the preparation of bioactive PMMA bone cements by incorporation of calcium silicate gels resulting from MPS and various calcium salts.59–61 A PMMA powder with a molecular weight of approximately 100 000 and an average particle size of about 14 μm was used for the studies. The PMMA powder was mixed with one calcium salt selected from CaCl2, calcium acetate (Ca(CH3COO)2), calcium hydroxide (Ca(OH)2), calcium lactate (Ca(CH3CHOHCOO)2) and calcium benzoate (Ca(C6H5COO)2) at 20 mass% of the powder. BPO was then added to the powder as a polymerization initiator. MMA liquid was mixed with MPS at 20 mass% of the liquid. N,N-dimethyl-p-toluidine (NDT) was then added to the liquid as a polymerization accelerator. The composition of the cements is shown in Table 16.1. The cement denoted as ‘Reference’ has compositions similar to those of the commercially available bone cement (CMW® 1, Depuy), containing neither MPS nor calcium salts. The powder was mixed with the liquid at a powder : liquid ratio of 1 g : 0.5 g at 23 ± 2°C. The paste was shaped into a rectangular specimen of 10 × 15 × 1 mm3 in size. At a half of the setting time Table 16.1 Compositions of the cements Powder (mass ratio)

Liquid (mass ratio)

Sample

PMMA

Calcium salt

BPO

MMA

MPS

NDT

Reference Modified cement

0.971 0.776

0.000 0.194

0.029 0.029

0.992 0.794

0.000 0.198

0.008 0.008

MPS, methacryloxypropyltrimethoxysilane; NDT, N,N-dimethyl-p-toluidine.

BPO,

benzoyl

peroxide;

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Orthopaedic bone cements

Reference

CaCI2

20 μm Ca(OH)2

Ca(CH3COO)2

Ca(CH3CHOHCOO)2

20 μm

20 μm

20 μm

20 μm

Ca(C6H5COO)2

20 μm

16.8 SEM photographs of the surfaces of the cements modified with MPS and various kinds of calcium salts after soaking in SBF for 7 days.

of the specimens, they were soaked in 35 cm3 of SBF at pH 7.25 for 7 days to examine apatite-forming ability. The compressive strengths of the cements without and with exposure to SBF were measured with a universal material testing machine. Figure 16.8 shows SEM photographs of the surfaces of the cements modified with MPS and various kinds of calcium salts after soaking in SBF for 7 days. Assemblies of fine particles were observed on the cements modified with CaCl2, Ca(CH3COO)2 and Ca(OH)2 after the soaking. The formed particles were identified as poorly crystalline apatite by TF-XRD as shown in Fig. 16.9. Figure 16.10 shows the compressive strengths of the cements modified with MPS and various calcium salts before and after soaking in SBF for 7 days.The compressive strength of the modified cements decreased after exposure to SBF, except for the cement modified with Ca(OH)2 . Among the cements examined in this study, those modified with Ca(CH3COO)2, Ca(OH)2 or Ca(CH3CHOHCOO)2 showed a compressive strength near the lower limit required by ISO 5833. Modification of PMMA cement by incorporation of MPS and appropriate kinds of calcium salts makes the cement capable of apatite formation in the body environment. Incorporated alkoxysilane and calcium salts would form calcium silicate gels during setting and/or after exposure to SBF through the sol–gel reaction. The formed calcium silicate gels were

Design of bioactive bone cement based on organic–inorganic hybrids

371

Intensity (arbitrary units)

Apatite

Ca(C6H5COO)2 Ca(CH3CHOHCOO)2 Ca(OH)2 Ca(CH3COO)2 CaCl2 Reference 25

30 35 2q (degrees)

16.9 TF-XRD patterns of the surfaces of the cements modified with MPS and various kinds of calcium salts after soaking in SBF for 7 days.

ISO 5833 lower limit Reference 07 dd CaCI2 0 d 7d

Ca(CH3COO)2 0 d 7d

Ca(OH)2 07 dd Ca(CH3CHOHCOO)2 07 dd Ca(C6H5COO)2 07 dd 0

20 40 60 80 Compressive strength (MPa)

100

16.10 Compressive strengths of the cements modified with MPS and various calcium salts before and after soaking in SBF for 7 days.

expected to induce apatite nucleation in SBF. The solubility of the calcium salts in water decreases in the order: CaCl2 > Ca(CH3COO)2 > Ca(CH3CHOHCOO)2 > Ca(C6H5COO)2 > Ca(OH)2. The cements modified with highly water-soluble calcium salts have the tendency to form apatite in SBF. It is noted that the cement modified with Ca(OH)2 formed apatite in SBF, in spite of the fact that the solubility of Ca(OH)2 is the lowest among the calcium salts used in this study. The pH of the surrounding solution remarkably increased after soaking of the cement modified with Ca(OH)2 in SBF. The increase in pH was expected to accelerate the apatite nucleation, since OH− is a component of the apatite. These findings indicate that the increase in pH, as well as the release of Ca2+, governs the ability of the apatite formation on the modified cements. All the cements apart from the

372

Orthopaedic bone cements

one modified with Ca(OH)2 showed a decrease in compressive strength after soaking in SBF. This is attributed to the release of Ca2+ ions from the cement into the solution. When Ca2+ ions are rapidly released, pores are formed inside the cement, leading to a decrease in compressive strength. Therefore, an appropriate rate of release of Ca2+ is desired to retain a high mechanical strength of the cement. Bioactive PMMA bone cements were obtained through material designs based on organic–inorganic hybrids. In this type of cement, the amount of additives can be decreased up to 20 mass% without significantly decreasing the mechanical strength of the cements within body environment. This suggests that bioactivity can be improved while maintaining the excellent handling performance of the cements.

16.6

Conclusions

Bioactive organic–inorganic hybrids with high flexibility can be designed on the basis of the bone-bonding mechanism of bioactive ceramics. PMMA bone cement can be provided with bioactivity through the addition of components of the bioactive hybrids. These cements are expected to possess high bone affinity as well as suitable handling performances.

16.7

References

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11 ohtsuki c, kokubo t, neo m, kotani s, yamamuro t, nakamura t, bando y, ‘Bone-bonding mechanism of sintered β-3CaO·P2O5’, Phosphorus Res Bull, 1991 1 191–196. 12 loty c, sautier jm, boulekbache h, kokubo t, kim hm, forest n, ‘In vitro bone formation on a bonelike apatite layer prepared by a biomimetic process on a bioactive glass–ceramic’, J Biomed Mater Res 2000, 49(4) 423–434. 13 kokubo t, kushitani h, sakka s, kitsugi t, yamamuro t, ‘Solutions able to reproduce in vivo surface-structure changes in bioactive glass–ceramic A–W’, J Biomed Mater Res, 1990 24(6) 721–734. 14 ohtsuki c, aoki y, kokubo t, bando y, neo m, nakamura t, ‘Transmission electron microscopic observation of glass–ceramic A–W and apatite layer formed on its surface in a simulated body fluid’, J Ceram Soc Japan, 1995 103(5) 449–454. 15 cho sb, nakanishi k, kokubo t, soga n, ohtsuki c, nakamura t, kitsugi t, yamamuro t, ‘Dependence of apatite formation on silica gel on its structure: effect of heat treatment’, J Am Ceram Soc, 1995 78(7) 1769–1774. 16 kokubo t, takadama h, ‘How useful is SBF in predicting in vivo bone bioactivity?’, Biomaterials, 2006 27(15) 2907–2915. 17 oyane a, kim hm, furuya t, kokubo t, miyazaki t, nakamura t, ‘Preparation and assessment of revised simulated body fluid’, J Biomed Mater Res, 2003 65A(2) 188–195. 18 oonishi h, kushitani s, yasukawa e, iwaki h, hench ll, wilson j, tsuji e, sugihara t, ‘Particulate bioglass compared with hydroxyapatite as a bone graft substitute’, Clin Ortho Rel Res, 1997 334 316–325. 19 oonishi h, murata n, saito m, wakitani s, imoto k, kim n, matsuura m, ‘Comparison of bone growth behavior into spaces of hydroxyapatite and AW glass ceramic particles’, in LeGeros RZ, LeGeros JP (Eds), Bioceramics, Vol. 11, Singapore, World Scientific, pp. 411–414, 1998. 20 oonishi h, hench ll, wilson j, tsuji e, kin s, yamamoto t, mizokawa s, ‘Quantitative comparison of bone growth behavior in granules of Bioglass®, A–W glassceramic, and hydroxyapatite’, J Biomed Mater Res, 2000 51(1) 37–46. 21 kim hm, himeno t, kokubo t, nakamura t, ‘Process and kinetics of bonelike apatite formation on sintered hydroxyapatite in a simulated body fluid’, Biomaterials, 2005 26(21) 4366–4373. 22 kim hm, himeno t, kawashita m, kokubo t, nakamura t, ‘The mechanism of biomineralization of bone-like apatite on synthetic hydroxyapatite: an in vitro assessment’, J R Soc Interface, 2004 1(1) 17–22. 23 ohtsuki c, kokubo t, yamamuro t, ‘Mechanism of apatite formation on CaO– SiO2–P2O5 glasses in a simulated body fluid’, J Non-Cryst Solids 1992, 143(1) 84–92. 24 li p, ohtsuki c, kokubo t, nakanishi k, soga n, de groot k, ‘The role of hydrated silica, titania, and alumina in inducing apatite on implants’, J Biomed Mater Res, 1994 28(1) 7–15. 25 ohtsuki c, kokubo t, yamamuro t, ‘Compositional dependence of bioactivity of glasses in the system CaO–SiO2–Al2O3: Its in vitro evaluation’, J Mater Sci Mater Med, 1992 3(2) 119–125. 26 ohtsuki c, osaka a, kokubo t, ‘Effects of Al2O3 and TiO2 on bioactivity of CaO–SiO2 glasses’, in Andersson OH, Yli-Urpo A (Eds), Bioceramics, Vol. 7, Oxford, Butterworth-Heinemann Ltd, pp. 73–78, 1994.

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43 goto k, tamura j, shinzato s, fujibayashi s, hashimoto m, kawashita m, kokubo t, nakamura t, ‘Bioactive bone cements containing nano-sized titania particles for use as bone substitutes’, Biomaterials, 2005 26(33) 6496–6505. 44 hu y, mackenzie jd, ‘Rubber-like elasticity of organically modified silicates’, J Mater Sci, 1992 27(16) 4415–4420. 45 tsuru k, ohtsuki c, osaka a, iwamoto t, mackenzie jd, ‘Bioactivity of sol–gel derived organically modified silicates, Part I: In vitro examination’, J Mater Sci Mater Med, 1997 8(3) 157–161. 46 chen q, miyaji f, kokubo t, nakamura t, ‘Apatite formation on PDMS-modified CaO-SiO2-TiO2 hybrids prepared by sol–gel process’, Biomaterials, 1999 20(12) 1127–1132. 47 osaka a, ohtsuki c, tsuru k, ‘Preparation of bioactive polymers modified with silanol groups’, in Wilson J, Hench LL, Greenspan D (Eds), Bioceramics, Vol. 8, Oxford, Elsevier Science Ltd, pp. 441–445, 1995. 48 yabuta t, tsuru k, hayakawa s, ohtsuki c, osaka a, ‘Synthesis of bioactive organic–inorganic hybrid with γ-methacryloxypropyl trimethoxysilane’, J SolGel Sci Tech, 2000 19(1–3) 745–748. 49 miyazaki t, ohtsuki c, tanihara m, ‘Synthesis of bioactive organic–inorganic nanohybrid for bone repair through sol–gel processing’, J Nanosci Nanotech, 2003 3(6) 511–515. 50 ratner bd, hoffman as, schoen fj, lemons je, Biomaterials Science, 2nd edition, Amsterdam, Elsevier Academic Press, 2004. 51 cho sb, nakanishi k, kokubo t, soga n, ohtsuki c, nakamura t, ‘Apatite formation on silica gel in simulated body fluid: its dependence on structures of silica gels prepared in different media’, J Biomed Mater Res (Appl Biomater), 1996 33(3) 145–151. 52 cho sb, miyaji f, kokubo t, nakanishi k, soga n, nakamura t, ‘Apatite formation on silica gel in simulated body fluid: effects of structural modification with solvent-exchange’, J Mater Sci Mater Med, 1998 9(5) 279–284. 53 plueddemann ep, Silane Coupling Agents, 2nd edition, New York, Plenum, 1991. 54 brinker cj, scherer gw, Sol-Gel Science, San Diego, Academic Press, 1990. 55 miyazaki t, ohtsuki c, ashizuka m, ‘Synthesis of osteoconductive organic– inorganic nanohybrids through modification of chitin with alkoxysilane and calcium chloride,’ J Biomater Appl, 2007 22(1) 71–81. 56 rhee sh, ‘Bone-like apatite-forming ability and mechanical properties of poly(epsilon-caprolactone)/silica hybrid as a function of poly(epsiloncaprolactone) content’, Biomaterials, 2004 25(7–8) 1167–1175. 57 kamitakahara m, kawashita m, miyata n, kokubo t, nakamura t, ‘Apatite-forming ability and mechanical properties of CaO-free poly(tetramethylene oxide) (PTMO)-TiO2 hybrids treated with hot water’, Biomaterials, 2003 24(8) 1357–1363. 58 kamitakahara m, kawashita m, miyata n, kokubo t, nakamura t, ‘Preparation of bioactive flexible poly(tetramethylene oxide) (PTMO)–CaO–Ta2 O5 hybrids’, J Mater Sci Mater Med, 18(6) 1117–1124. 59 ohtsuki c, miyazaki t, kyomoto m, tanihara m, osaka a, ‘Development of bioactive PMMA-based cement by modification with alkoxysilane and calcium salt’, J Mater Sci Mater Med, 2001 12(10–12) 895–899.

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17 Clinical aspects of calcium phosphate bone cements S. L A R S S O N, Uppsala University Hospital, Sweden

Abstract: A number of injectable calcium phosphate cement (CPC) products have recently been developed for augmentation of crushed metaphyseal fractures. After injection, the cement hardens within minutes, after which a mechanical strength is achieved that corresponds to good cancellous bone. The chemical composition mimics bone and the material remodels through a cell-mediated process. Clinical studies have shown good results especially when used as bone graft substitute in tibial plateau fractures. In other anatomical areas the results have also been promising, although more studies are needed to define the proper indications. In the present chapter injectable CPC will be described more in detail with special reference to the use in different aspects of fracture treatment. Key words: calcium phosphate, cement, fracture treatment, surgery.

17.1

Introduction

The solution to some of the problems encountered when fixing fractures involving crushed cancellous bone, especially in osteoporotic patients, may lie in injectable calcium phosphate cement (CPC). Over the last decade a number of different products based on calcium phosphate have been introduced into the orthopaedic market (Fulmer et al. 1992, Driessens et al. 1994, Constantz et al. 1995, Kurashina et al. 1997). Several of these have also been studied in recent years as potential tools for use in fracture surgery. The use of CPC as a bone graft substitute for filling of voids in metaphyseal fractures involving compromised cancellous bone has been identified as a very promising indication. If CPC can be used instead of autologous bone for filling of metaphyseal defects, it also means that all the side effects related to the bone harvest site can be avoided. Stabilisation of compressive fragility fractures in the spine by percutaneous injection of cement, i.e. vertebroplasty, is now a well-recognised technique to reduce pain in osteoporotic patients. So far, the most commonly used materials for filling of crushed verterbrae are different compositions of poly(methylmethacrylate) (PMMA). Due to the side effects associated with these conventional bone cements, CPC compounds have been mentioned as a possible alternative. 377

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Orthopaedic bone cements

Another potential indication is to use CPC around screws to enhance the holding properties for screws when used for fixation of fractures in weak osteoporotic bone. An injectable cement that after curing improves the immediate fracture stability without impairing bone healing, and is strong enough to allow early and active rehabilitation, will most certainly add new possibilities when dealing with certain fractures, especially in weak bone (Larsson and Bauer 2002). In particular, if time with restricted weight bearing after surgery can be reduced, when compared with conventional treatment with a metal and bone transplant, it might be possible to undertake a more aggressive rehabilitation and thereby restore function more rapidly. In most fracture indications an ideal cement would be gradually resorbed with time and replaced by host bone. The rate of cement resorption must be balanced with the rate of new bone formation to avoid collapse at the fracture site, which might occur if resorption is too fast. The purpose of this chapter is to describe injectable CPCs, with special reference to their clinical use in different aspects of fracture treatment.

17.2

Material characteristics

17.2.1 General appearance Most orthopaedic surgeons are used to the handling properties and the performance that can be expected when using conventional PMMA in joint replacement procedures. It might therefore be tempting to transfer that experience into what can be anticipated when using CPC in fracture treatment. However, due to major differences between PMMA and CPC, it is very important to realise that CPC, in almost all aspects – i.e. chemical composition, setting time, handling properties, chemical reaction during setting, mechanical properties after setting, as well as the biological response over time after the procedure – is completely different when compared with PMMA. Almost all CPC products are delivered as kits that include one or two powders that consist of different calcium salts and a fluid that usually comprises a sodium phosphate solution. The chemical composition is described in more detail in Chapter 10 of this book. Briefly, it can therefore just be mentioned that, following curing, CPC resembles the mineral phase of bone tissue and, once implanted, will be subjected to a cell-mediated remodelling that seems similar to the natural bone remodelling. After mixing, either manually or with a mixing machine, a paste is formed in one or a few minutes, the timing varies slightly for different cement types. The paste that is formed has, for the vast majority of the commercially available CPCs, a consistency that resembles toothpaste. For a few minutes it is injectable

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through a needle with a size of about 8–12 gauge. For a few products, the composition has been designed in such a way that the material after mixing has a dough-like appearance, making it mouldable rather than injectable.

17.2.2 Mechanical properties After injection, the paste hardens in situ due to crystallisation within a few minutes, after which it attains an initial compressive strength of about 5–10 MPa. During the initial minutes of the curing process it is important to avoid movement of the material or subject it to vibration, as this might disturb the chemical process to such an extent that the material will not gain the anticipated mechanical strength. The speed of the crystallisation process that starts at the time of mixing is highly temperature sensitive, so it is important to consider this when working in, for instance, a very warm environment. Even a slightly higher than normal temperature in the operating room will shorten the time available for injection due to premature curing. The hardening process itself does not, in contrast to, for instance, PMMA cause any rise in temperature above physiological levels, or any change in the pH. These are important properties when dealing with fracture treatment as it means that the material will not cause any disturbance of the local healing capacity through, for instance, heat necrosis. During the first 12–24 hours after injection the crystallisation continues, leading to a final compressive strength of about 25–50 MPa. The strength varies between different cement types although, in general, the final strength will be in line with, or better than, the compressive strength in cancellous bone. In fact, following injection the material interdigitates with adjacent bone, forming a solid bone–cement structure that is more mechanically stable than when using either cancellous bone graft or preformed hydroxyapatite pellets or blocks. The mechanical strength in tension and shear is, however, much lower with reported values seldom exceeding 3–5 MPa. These mechanical properties, with good compressive strength and limited strength in tension and shear, are important facts to bear in mind when using CPC in the clinical setting. For many potential indications the consequence will be that the cement itself is strong enough to neutralise the compressive loadings that it will be subjected to; while some kind of metal, or other structural implant, should be used in conjunction with the CPC if the material will be subjected to a more substantial bending, tension or shear. As combined and often complex loadings are common in the skeletal system, this means that for most indications it is advisable to use a combined fixation technique. With such an approach the cement is used to counteract the compressive loading, while bending and shear are neutralised through a structural implant. If a combined fixation technique is not used, there is a

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risk that the CPC will fracture even at rather limited loads due to the brittle appearance when subjected to loadings other than pure compression.

17.2.3 Biological performance Several animal studies have shown that CPC is osteoconductive and undergoes gradual remodelling with time (Sarker et al. 2001, Oooms et al. 2003). In a canine study (Frankenburg et al. 1998), a bone defect created in the proximal tibia and filled with CPC showed prominent bone apposition as early as 2 weeks (Fig. 17.1). During the following months there was evidence of osteoclastic resorption of the cement, vascular penetration and bone formation in a pattern that suggested remodelling similar to that of normal bone, with no foreign body reaction, i.e. good biocompatibility. There have also been some reports of the histology of CPC in humans based on biopsies or assessment of retrieved specimens. Schildhauer et al. in 2000 described the use of CPC in augmenting internal fixation of calcaneal fractures. Biopsies were obtained from seven patients at the time of hardware removal, more than 1 year after injury. The biopsy specimens showed almost complete bone apposition to the CPC. Cement resorption was evident in the vicinity of osteoclasts and often was accompanied by new bone formation that appeared qualitatively similar to normal bone

Osteoblasts CaP

Bone

Osteoclasts

17.1 Photomicrograph demonstrating a focus of osteoclastic resorption and new bone formation associated with both bone and CPC in a biopsy obtained from an elderly patient more than 9 months after treatment with cement to reinforce an internal fixation of a proximal femoral fracture (Larsson et al. 1999).

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remodelling. Vascular ingrowth into the cement was evident, and there was no fibrous tissue. In a study on femoral neck fractures in human patients (Larsson et al. 1999), CPC was used for augmentation around cannulated screws used for internal fixation. Femoral heads were retrieved from several patients who had non-union and underwent total hip arthroplasty up to more than 2 years after the primary procedure. The femoral heads had a complex histological appearance, showing areas with extensive bone apposition to the cement, and in other areas, the bone marrow contained macrophages with fragments of cement that had been phagocytosed. Overall, the histological features showed processes of remodelling identical to those previously reported in animal studies and similar to those seen in normal bone. However, the study also showed that the cement may fragment if associated with unstable metal hardware.

17.2.4 Radiological appereance Owing to the high content of calcium salts, the CPC products will have inherently high radiographic attenuation, i.e. be naturally visible on radiographs without adding contrast medium to the product. A bone region implanted with CPC therefore appears quite radiopaque on conventional radiographic films. As the cement undergoes remodelling, as shown from biopsies and specimens mentioned above, one would expect to see a proportionate change in the radiopacity over time. However, owing to the low resolution and limitations of conventional radiographs, this change is usually not visible until a substantial part of the material has been replaced by bone. It has been shown that as much as 50% of the material can be taken away in vitro before the difference is detectable on radiographs. Conventional radiographs are therefore helpful when assessing positioning and containment of CPC following a procedure, but they are a blunt tool for assessment of the degree of remodelling. Still, even with the limitations when using conventional radiographs for assessment of remodeling, it seems reasonable to state that the remodelling process is very slow. One reason for the slow remodelling is the lack of macropores. This means that remodelling, or degradation, takes place as a surface phenomenon, layer by layer. If large volumes of CPC are being used, the degradation will probably not be complete even over several years. On the other hand, as the remodelling is cell mediated the resorption and the new bone formation are coupled, which means that there should be no risk for resorption being too fast, leaving a void before new bone will fill in. This is an important aspect as there are other bone graft substitutes with a high solubility similar to calcium sulphate products and which dissolve without a cell-mediated control, where formation of a new void has been described if the pace of the new bone formation is not fast enough.

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17.3

Surgical technique

Proper understanding of the mixing procedure and the injection technique, as well as knowledge of the mechanical properties that can be expected once the cement has cured, are vital in order to get reproducible results when using CPC. So far, the most common use of CPC is for filling of bone defects and subchondral voids in metaphyseal fractures. The cement is then often used as a substitute for bone transplant and in conjunction with conventional metal implants as previously mentioned. Following reduction of the fracture and provisional fixing with wires, the surgeon has two options as the cement can be injected either prior to or after the definitive placement of the hardware. Both options have advantages and drawbacks, and it might be a good idea to remember that both options are available and that for certain cases one option might be better than the other. It is well known that the material should not be manipulated during curing as this might negatively affect the mechanical strength. In order to avoid unintentional movement during curing, due to insufficient fixation, it is often technically easier and more reliable if the definitive hardware has been applied prior to cementing. Once the fracture is fixed with the definitive hardware, if necessary with additional provisional wire fixation still in place, the surgeon has got a stable situation. This will allow a thorough injection of the cement without any risk of secondary fracture displacement during injection or during curing of the cement. When using cement for filling of metaphyseal defects, it is very important to prepare the void carefully prior to cementing. In fact, this is one of the most important technical differences to consider when using CPC compared with when using bone transplant for filling of metaphyseal voids. When using conventional bone transplant for filling a void, the bone is usually impacted with instruments. This will also indirectly impact the cancellous bone that makes up the walls of the void. Thereby, a rather homogeneous structure is formed that is composed of well-impacted cancellous host bone and impacted bone transplant, whether autologous bone or allograft is being used. When CPC is injected into a void, it will not cause any impaction on the surrounding cancellous bone as the cement will only passively flow into the void and interdigitate with the bone. If the bone surrounding the void is crushed due to the primary injury, and is also of low density – as in elderly osteoporotic patients – the cancellous bone surrounding the void will become the weak part of the bone–cement mechanical construct. It is therefore very important that the surgeon, before cementing, impacts the cancellous bone that surrounds the void until it feels as if it has a homogeneous density and strength. In addition, there should not be any loose pieces of bone or debris in the void at the time of cementing, because such debris might negatively affect the goal of getting a monolithic block

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of CPC which in a solid way interdigitates with the surrounded wellimpacted cancellous bone. As previously mentioned, it is important to understand the mechanical properties of CPC, i.e. in general, strong in compression and weak in bending and shear. Even though these properties often make it necessary to retain some kind of hardware, such as screws or plate and screws, the overall properties of the CPC often make it possible to reduce the amount of conventional hardware. Such a combined technique with CPC and a reduced amount of hardware also paves the way for the development of new minimally invasive procedures.

17.4

Complications during surgery

17.4.1 Problems during injection A free flow through the cannula with no need for extensive injection pressure is a prerequisite for a successful cementing. If the cannula is too thin, or if there is some kind of obstruction in the needle that will reduce the working diameter of the needle, this will have major negative effects on the flow. If an obstruction is met with the application of more pressure on the syringe, there is a substantial risk for separation, or demixing, as the fluid comes out without the calcium phosphate particles. This is the same effect as seen when applying pressure on wet sand if the water can escape through a membrane, the so-called filter pressing effect. If there is a feeling of obstruction during injection, or if unexpected power is needed, this should be managed by changing to a new needle or by taking away whatever might be causing the obstruction. The most common problem seems to be a small bone piece obstructing the needle or the needle being pushed towards the wall of cancellous bone surrounding the void. Another problem during injection might be premature curing of the cement within the syringe or the cannula, which will prevent the final part of the cementing. This is often caused by unintentional warming up of the material prior to or during injection. Even a slight increase in the temperature will accelerate the curing process substantially. For instance, keeping the hand around the syringe during filling might be enough to increase the temperature for certain products to the extent where the curing properties will change resulting in a much more rapid setting than normal.

17.4.2 Incomplete void filling A major advantage with CPC compared with preformed blocks or pellets is that the cement can be used to completely fill any void, irrespective of

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size or shape. By a complete filling of the void, the mechanical strength of the bone–cement construct will be much higher, and the risk for secondary displacement will be far less, compared with a void that is only partially filled. As most CPC products are available as packages with different volumes, the surgeon has to estimate before cementing what amount, or volume, will be needed for complete filling. Metaphyseal voids are usually irregular and differ between cases in such a way that it is hard to get a good reference, or a standard, for different indications. This is especially obvious when using percutaneous techniques. The size of the void has then to be estimated, based on the tactile feeling from small instruments and the appearance on the fluoroscopy. In order to avoid incomplete filling, it is reasonable, as a general rule, to use a CPC package size that is anticipated to leave some excessive material once the void has been filled. By having some excessive material, this will allow a monolithic strong CPC block to be formed in the void. The excessive material can easily be washed away or simply left in the syringe. If the void is larger than expected and the first package of CPC does not provide complete filling, one can either wait until the first batch has cured or go on with the second injection before the first being cured. There has been some theoretical concern on formation of a weak interface between the two batches if the first one is allowed to cure before the second cementing is done. However, experimental testing has not revealed any major drawbacks attributable to a weak interface.

17.4.3 Leakage Most CPCs have a very low viscosity during injection. One advantage with this consistency is that it can be injected through needles with a relatively small diameter. Another potential advantage is that the cement will flow into the cancellous bone and interdigitate with the trabeculae creating strong bonds between the CPC and the bone. A drawback with the low viscosity is the risk of leakage outside the void through even very narrow fracture lines and cracks and thereby the cement will enter into the surrounding tissues. By definition, metaphyseal fractures are close to joints, which means that any intra-articular fracture line is a potential track for leakage of CPC into the joint during the injection. Leakage is also frequently seen and has been reported for various metaphyseal fractures. It has, however, been shown that the material disappears very quickly from the joint without causing any visible harm. The most likely mechanism is that the material, through shear loading in the joint, breaks into small pieces that later are phagocytosed by

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macrophages in the synovial membrane. In animal studies it has been shown that macrophages effectively take care of CPC that has intentionally been placed in joints, and it seems reasonable to believe that the same mechanism will prevail in humans. Leakage outside the bone into the surrounding muscles or into any other soft tissue is a common finding, even though cleaning and washing away of any visible protruded material outside the bone is part of the recommended technique. As the cement is often injected either percutaneously or by the use of minimally invasive techniques, it might be technically difficult to clear the soft tissue completely of all CPC that might have leaked outside the bone. However, based on the rapid disappearance of CPC from soft tissue, there is no need to use extensive methods to extract small or limited amounts from the soft tissue. It seems that the best way to avoid leakage, while still achieving adequate filling of the void, is to use a retrograde filling technique: starting with the tip of the needle at the far end of the void, the needle is slowly withdrawn, while the cement is injected by the application of an even pressure on the syringe.

17.4.4 Wet field properties Even when using a tourniquet during extremity surgery, the bone surfaces and the surrounding soft tissues are wet. If a tourniquet is not being used, all surfaces are wet and some degree of bleeding is always present. As CPC usually has low viscosity and hydrophilic properties, a wet field – and especially the presence of bleeding – might create technical problems during injection. The surgical technique should therefore always involve the use a tourniquet when possible and, in addition, different techniques and tricks should be included to get the surfaces as dry as possible at the time of cementing. For several of the more recently developed CPC products, the manufacturers claim that the wet-field properties have been greatly improved, with better cohesion compared with the first generation. There are also products available with a higher viscosity compared with the usually low viscosity in most CPCs. For such mouldable, rather than injectable, cements the material will more easily stay intact during curing in a wet environment. In addition, a higher viscosity and better cohesion will most probably reduce the risk of leakage. A potential drawback with mouldable CPC is that it might be difficult to achieve a filling of the void that is as complete as that achieved when using low viscosity CPCs. The higher viscosity will also make it necessary to use needles with a larger diameter or even an open application technique.

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17.5

Clinical applications

17.5.1 Metaphyseal fractures Internal fixation of metaphyseal fractures may be difficult because of the low strength in the cancellous bone that makes up the major part of the metaphysis. In addition, the fragment closest to the joint is often short, and the fracture frequently has an intra-articular component. Despite the technical difficulties introduced by these problems, internal fixation usually is recommended because good fracture stability is a prerequisite for early return to adequate joint function. Preservation of a good reduction of the articular surface until healing is also important in order to avoid malalignment, instability and subsequent degenerative arthritis. Good fracture stabilisation is especially important in patients with osteoporosis who may have general weakness that prevents them from reducing the load on their injured extremity during healing. Two factors frequently complicate stabilisation of metaphyseal fractures. The first is the crushing of subchondral cancellous bone by depressed articular fragments at the time of injury. Reducing the fracture during surgery creates a metaphyseal defect that must be filled to prevent the joint surface from subsiding when the bone is subjected to loading. Most commonly, the subchondral void is filled with autologous bone graft. However, this procedure has serious drawbacks, such as donor site morbidity and the limited immediate mechanical stability afforded by cancellous autograft. The second factor that complicates stabilisation of such fractures is the presence of osteoporosis, which increases the risk that a fracture will become displaced before it heals.

Distal radius Fractures of the distal end of the radius are common injuries. Treatment with closed reduction and plaster usually works well, especially for patients who have little initial displacement, limited comminution and good bone quality. However, fractures with severe comminution, bone loss at the fracture site or involving osteoporotic bone may be difficult to stabilise with plaster alone, and the risk for secondary displacement and malunion is high. The use of a bioactive cement may make it possible to fill the metaphyseal defect, improve fracture stability and reduce immobilisation time. In a biomechanical study, Yetkinler et al. (1999) tested the stability of intra-articular fractures of the distal radius stabilised with Kirschner wires or CPC. The fractures augmented with cement were significantly more stable and had a significantly higher strength when subjected to cyclic loading and when loaded to failure.

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The use of CPC to treat distal radius fractures has been studied for fresh fractures (Jupiter et al. 1997, Sanchez-Sotelo et al. 2000, Cassidy et al. 2003) and for fractures that had become displaced after primary treatment with closed reduction and plaster (Kopylov et al. 1999). In fresh fractures, percutaneous injection of the cement worked well in providing adequate filling of the fracture void. In patients who needed a secondary procedure after redislocation while wearing a plaster cast, an open cementing technique that includes evacuating the organised haematoma before injecting the cement was recommended. In a randomised multicentre study of patients with fresh distal radius fractures, cement augmentation combined with a short period of plaster immobilisation was compared with conventional external fixation. The group treated with cement augmentation had a significantly faster gain in grip strength and range of movement for the first 8 weeks after the injury, although there were no significant differences later, to a maximum follow-up of 1 year. Similar results were reported in a randomised study on redisplaced fractures of the distal radius where cement augmentation combined with 2 weeks of plaster immobilisation was compared with external fixation for 5 weeks. Proximal humerus Most fractures through the proximal humerus can be treated nonoperatively. However, for certain fracture types, including severely impacted valgus fractures, restoration of a good shoulder function seems to be associated with reduction and stabilisation. In order to reduce the risk for further impairment of the blood circulation to the humeral head caused by the trauma, it seems reasonable to utilise minimally invasive techniques for reduction and fixation. Based on published reports, it seems that reduction using a percutaneous technique can provide an acceptable position of the main fracture as well as of the fractured tuberosities. The commonly reported problem is instead secondary displacement due to inability to provide an adequate fixation during the course of healing. One reason for this frequent tendency for secondary displacement seems to be the lack of cancellous bone support due to the bone defect and the void created behind the reduced humeral head. A few years ago, Robinson and Page (2004) addressed this problem by filling the subchondral void with CPC. In a case series they treated 29 patients with severely impacted valgus fractures with open reduction, and filling of the space behind the humeral head with CPC. In addition, the fractures were fixed with either screws or a buttress plate. All fractures united and all the reductions were maintained. Notably, no patient had any signs of osteonecrosis of the humeral head at final follow-up at between 1 and 2 years. When comparing the functional results with similar case series treated non-operatively, the authors concluded that

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the results using this technique seemed to be better than those for nonoperative treatment. Calcaneus The relative benefits of surgical versus non-surgical treatment of calcaneal fractures are controversial. To some extent, the reluctance to carry out surgery on these fractures seems to be attributable to the technical difficulties of trying to restore the anatomy of the subtalar joints. Not only is it often a substantial technical challenge to reduce the multiple bone fragments, but in addition it is often difficult to maintain joint congruency during healing. The limited stability of the fracture fragments achieved with metal fixation is, to some extent, due to the cancellous bone defect that is regularly present under the subtalar joints. This defect is due to the crushed cancellous bone, similar to that seen in tibial plateau fractures, made worse by the reduced cancellous bone density commonly found in this part of the calcaneus. Whether bone graft is being used or not to fill this defect, restricted weight bearing is usually recommended for 8–12 weeks. The rationale for using an injectable, bioactive cement in calcaneal fractures is to fill the void under the subtalar joints, thereby providing an augmented construct. This should allow earlier return to full weight bearing and faster restoration of ankle and foot function without increasing the risk for secondary displacement of the fractured joint surfaces. In a biomechanical study, Thorardson et al. (1999) compared intra-articular calcaneal fractures fixed with standard internal fixation with those in which the osseous defect was filled with either bone graft or CPC. In the specimens treated with the combination of hardware and CPC, cyclic loading produced significantly less deformation. In a prospective clinical series by Schildhauer et al. (2000), 36 calcaneal fractures with joint depression were treated with a combination of internal fixation and CPC. Through the series, progressively shorter times to full weight bearing were prescribed. The last patients were allowed full weight bearing 3 weeks after surgery and showed no radiological evidence of loss of reduction. In another clinical series (Thordarson et al. 2005), 15 patients with displaced intra-articular calcaneal fractures were treated with a similar technique utilising a combination of conventional metal hardware applied laterally, while the defect below the posterior facet was filled with CPC following open reduction. The first 6 patients were allowed weight bearing at 6 weeks while the following 9 patients were already allowed to bear weight 3 weeks after the procedure. No patient had any visible loss of reduction based on radiographs until healing. Based on those two clinical series as well as on biomechanical studies, it seems reasonable to believe that augmentation with CPC will

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provide a more stable fixation when compared with metal fixation alone. However, so far no controlled randomised study has been published that compares the outcome when using internal fixation with and without augmentation with CPC in calcaneal fractures. Tibial plateau Fractures of the tibial plateau often involve depression of the articular fragments and instability that make surgery necessary to restore intra-articular anatomy and joint stability. Elevating the depressed articular fragments often reveals a metaphyseal void outlined by crushed cancellous bone. This void is commonly filled with autologous or allogenic bone graft or with preformed blocks or granules composed of sintered, highly crystalline hydroxyapatite (Bucholz et al. 1989). Although autograft has desirable osteoconductive and osteogenic properties, cancellous bone graft does not provide enough mechanical stability to allow full weight bearing until the fracture has healed. CPC and other injectable bioactive cements, however, can completely fill the defect to produce immediate stability, and they also provide enough compressive strength to facilitate earlier active joint motion and shorter time to full weight bearing than bone graft procedures (Yetkinler et al. 2001). In a series of 41 patients with an isolated tibial plateau fracture, Keating et al. (2003) allowed free weight bearing 6 weeks after surgery when using CPC to supplement internal fixation in tibial plateau fractures. Despite the early weight bearing, they reported substantial loss of reduction in only one patient, an elderly man with poor compliance. Metal was used to support the bone–cement construct, but in most cases less metal was required than would have been used for fixation without cement augmentation. In another randomised study (Larsson et al. 2004), the aim was to measure the stability of the elevated articular fragment until healing and also to measure actual weight bearing. A total of 30 patients with a lateral tibial plateau fracture were allocated to filling of the subchondral void with either autologous bone or CPC, while conventional hardware was used in all patients with no differences between groups. Radiostereometry, a radiological technique that allows three-dimensional description of movement with a high accuracy, was used to measure the stability of the reduced articular fragment until healing. Patients randomised to treatment with CPC were allowed full weight bearing at 6 weeks, compared with 12 weeks in the bone-grafted group, and the actual weight bearing was measured at each follow-up visit. Despite the fact that patients treated with CPC applied significantly more weight earlier on when compared with the bone-grafted patients, the articular fragment remained more stable with less subsidence until healing in the CPC group compared with those that were bone grafted.

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Patients treated with CPC also had less pain during the course of healing. In another recent randomised multicentre study by Russell and others (2004), 120 patients with a tibial plateau fracture were randomised to filling of the subchondral defect with either CPC or bone graft. The authors came to the very strong conclusion that due to fewer complications and better stability of the articular fragment they discouraged future use of bone graft for this indication in favour of bioresorbable CPC. Based on these randomised studies and other studies reporting clinical series (Lobenhoffer et al. 2002), it seems that augmentation of tibial plateau fractures with CPC in the subchondral void, instead of bone graft, is a very suitable indication (Fig. 17.2). Even so, it might be reasonable to wait for studies with longer follow-up, i.e. patients followed until the material has totally, or nearly totally, been degraded and replaced with normal host bone, before making a more definite statement on the role for CPC in tibial plateau fractures. Hip fractures Surgical treatment is the preferred option in almost all patients with hip fracture. The surgeon should aim to produce fracture stability that will allow unrestricted weight bearing directly after surgery. This is because most patients who experience a hip fracture are elderly and have limited strength in their upper extremities which makes it difficult for them to reduce the weight carried by the injured hip. Unfortunately, the bone is often weakened by osteoporosis, which hampers the ability to achieve stability. In addition, necrosis, non-union and complications after internal fixation of displaced femoral neck fractures are frequent because of disturbed blood circulation to the femoral head. As a result, primary prosthetic replacement is often the treatment of choice. However, even though deficient blood flow is the primary cause of complications, there are reports indicating that good primary stability can improve outcomes after internal fixation (Rehnberg and Olerud 1989). Stankewich et al. (1996) studied the use of CPC to augment comminuted femoral neck fractures fixed with cannulated screws. Under cyclic loading, the augmented specimens were significantly stiffer than specimens fixed with screws alone, and they failed at higher loads. In a prospective, randomised clinical study (Mattsson and Larsson 2006), 118 patients with displaced femoral neck fractures were allocated to treatment with closed reduction and internal fixation with two cannulated screws alone or in combination with CPC augmentation. The cement was used to augment the bone around the screw threads to enhance the holding properties in the femoral head but also to fill the fracture void. Examination with radiostereometry showed significantly better stability in the cement-augmented

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(a) (b)

(c)

17.2 Images from a 79-year-old osteoporotic man. Lateral tibial plateau fracture as seen on conventional radiographs (a) on computed tomography scan (b) and after percutaneous reduction and fixation with screws and CPC (c). Unrestricted weight bearing from 3 weeks after surgery.

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fractures with less overall movement, less distal migration of the femoral head fragment and less varus angulation during early rehabilitation, although there were no significant differences between the groups later on (Mattson et al. 2003). In fact, there was a trend towards more re-operations in the augmented group towards the end of the 2-year follow-up. This finding was interpreted as indicating that the cement inserted at the fracture void might have caused an additional impairment of the vulnerable circulation to the femoral head leading to an increased risk of healing complications. The authors therefore recommended that the cement should not be used for filling of the fracture void when fixing femoral neck fractures. Trochanteric fractures, which are almost as common as femoral neck fractures, rarely present any healing difficulties. The main concerns instead are related to mechanical problems when attempting to achieve stability. Two-part fractures seldom present any problems, but multifragmentary fractures, especially those without posteromedial support, often are technical challenges. In two biomechanical studies (Elder et al. 2000, Yetkinler et al. 2002), trochanteric fractures with a detached minor trochanter were created and fixed with a sliding screw device alone or in combination with CPC to fill the posteromedial defect. Cyclic loading revealed that the specimens augmented with cement had significantly higher stiffness, stability and strength, and less shortening at the fracture site. In addition, the strain on the medial bone surface in the augmented specimens was closer to normal than the strain in the specimens fixed with the metal device alone. In a few clinical trials it has been shown that augmentation with PMMA when dealing with unstable trochanteric fractures can provide improved stability and low complication rates (Schatzker et al. 1978, Muhr et al. 1979, Bartucci et al. 1985). However, drawbacks with PMMA, including fear of disturbed fracture healing and difficulties if revision surgery becomes necessary, has made surgeons reluctant to use PMMA for fracture augmentation. As CPC has different properties to PMMA, it might be that the CPC could be useful for augmentation of unstable trochanteric fractures (Goodman et al. 1998). In a prospective, randomised multicentre clinical trial (Mattsson et al. 2005), patients with a trochanteric fracture that included a detached posteromedial fragment were treated with surgical fixation using a conventional sliding screw system alone or in combination with CPC (Fig. 17.3). The purpose of augmentation was to restore a mechanically competent posteromedial arch that enabled a more efficient transfer of load between fracture ends while reducing the risk for secondary fracture displacement and cutting out of the lag screw. Radiostereometry was used to measure fracture movement until 6 months after surgery, when all fractures were healed. The fractures augmented with cement were significantly more stable at all times, with reduced overall movement, less distal displacement

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(b)

17.3 Image from an 82-year-old osteoporotic women. Unstable trochanteric fracture (a) before and (b) after internal fixation with a sliding screw device and CPC for augmentation. Unrestricted weight bearing immediately after surgery.

and less varus angulation when compared with fractures fixed with the metal device alone (Mattsson and Larsson 2004). It was concluded that, for severely unstable trochanteric fractures, augmentation with CPC might be an attractive option for better fracture stability and also for reduced pain during the early phase of the rehabilitation.

17.5.2 Spine Vertebroplasty Vertebral compression fractures in osteoporotic patients are relatively common, and are often responsible for persistent and severe pain. Injection of PMMA cement into the vertebral body under fluoroscopic guidance has been shown to give considerable pain relief and possibly structural strengthening of osteoporotic vertebral bodies (Deramond et al. 1998, Belkoff et al. 1999, Cortet et al. 1999, Barr et al. 2000, Grados et al. 2000, Majd et al. 2005).

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Early results suggest that CPC may be useful for vertebroplasty procedures as an alternative to PMMA. The mechanical properties of CPC, with the highest strength being in compression, make it suitable for this indication, as the material mainly will be subjected to compression after injection into the vertebrae. A problem that has been observed when using CPC for vertebroplasty, as well as when using PMMA, is the risk of leakage, either locally outside the vertebrae being injected or into a vessel, with possible risk of severe embolic complications. Filling of the vertebrae with cement can be carried out using two different techniques. With so-called percutaneous vertebroplasty (PVP), the technique was described already in the 1980s, no attempt is made to restore the vertebral body height and thus correct the kyphotic deformity of vertebral collapse. More recently, a procedure called kyphoplasty (KVP) has been developed to address this issue as well as to reduce the risk of leakage of cement outside the vertebra. An inflatable balloon is inserted into the centre of the fractured vertebra using the trans- or extrapedicular approach. The balloon is then inflated to impact cancellous bone circumferentially around the balloon and thus create a void while at the same time reducing the deformity. The central void is then filled with cement. Theoretically at least, a preformed cavity in the vertebral body will reduce the risk for extrusion of CPC outside the vertebrae and is thus advantageous. However, it still has to be shown if the risk of leakage of CPC outside the vertebral body can be reduced when using the KVP technique compared with PVP. Pedicle screw augmentation Transpedicular screws have been shown to provide a very good anchorage for any spinal device aimed for stabilisation of a spinal segment. The pullout strength for pedicle screws is, however, very dependent on the bone quality. When dealing with osteoporotic vertebrae, the holding properties can be expected to be much lower compared with vertebrae with normal bone density. Due to the limited diameter of the pedicle, it is not possible to increase the size of the screw or the threads too much to improve the holding strength. Another alternative would therefore be to reinforce the bone that surrounds the pedicle screw, i.e. augmentation. A few studies, mainly biomechanical but also a few clinical reports, have shown promising improvement of the holding power and reduced risk for screw loosening in low density vertebrae, when using cement augmentation. So far, the focus has been on the use of conventional PMMA. The problem with PMMA in this indication seems to be the increased difficulty of screw removal due to the strong bond formed between the screw and the cement. During extraction there is a risk that, when torque is applied, this will cause damage to the bone surrounding the cement. Due to the size of the screw with the

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attached cement mantle still on, it will be very hard to remove through the narrow pedicle without causing damage. As CPC has different mechanical properties compared with PMMA, this might become a useful alternative whenever augmentation around pedicle screws is considered necessary as the extraction torque, based on experimental studies, will cause a disruption between the screw and the CPC. Extraction will be almost like screw extraction from good cortical bone.

17.5.3 Bone augmentation Problems related to internal fixation of fractures in osteoporotic bone have usually been addressed by developing new metal implants. An alternative strategy is to develop methods that will enhance the strength of the weak cancellous bone surrounding the metal implant, as previously mentioned for pedicle screw augmentation. The most common technique is to inject the cement into the predrilled hole after which the screw is inserted while the cement is still not cured. Once the cement hardens, the screw becomes embedded in mechanically strong cement, whereby the fixation strength can increase dramatically. When compared with screws without augmentation, the relative improvement in holding power following such cement augmentation seems to be most pronounced in bone with low density, i.e. in bone where improvement in holding power is most needed (Eriksson et al. 2002, Andreassen et al. 2004). A few studies have also shown that PMMA increases the power and fixation strength of screws when inserted in osteoporotic bone. However, PMMA has not gained wide acceptance for this indication. One reason is probably the difficulties encountered when removing the cement if revision surgery becomes necessary. Augmentation around screws with CPC might, on the other hand, become a useful technique, especially when dealing with low quality bone, without the drawbacks associated with PMMA. In a biomechanical study by Eriksson et al. (2002), the holding properties for different screws were assessed when inserted into synthetic bone with or without augmentation. The study showed that augmentation with CPC around the implants increased the holding strength significantly, especially in low density bone. At extraction it was also shown that the CPC-augmented specimens failed between the screw surface and the surrounding cement, while for the PMMA-augmented specimens the failure occurred between the cement mantle and the surrounding bone. As a consequence, a rather large void might be created when extracting a PMMAaugmented screw if the breakage, when the extraction torque is applied, occurs at the bone–cement interface. For the CPC-augmented screws, the extraction torque will, in a much more predictable way, create loosening between the screw and the surrounding cement, i.e a situation similar to screw extraction from bone alone.

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With the use of locking plate constructs, results have improved compared with conventional plates and screws when dealing with weak bone. However, CPC augmentation can probably further improve the screw fixation by adding to the overall strength of the bone that surrounds the implant; an important improvement that might prevent cutting out, as angle stable screws do not primarily fail through pull-out but through cutting out through the bone.

17.6

Future trends

Calcium phosphate compounds such as CPC will most certainly be part of the future armoury when dealing with fracture treatment. It is reasonable to believe that we have so far only seen the very beginning when it comes to clinical applications. As a bone graft substitute, CPC has already been established for filling of voids in several types of metaphyseal fractures. So far, the best documented indication seems to be for augmentation of tibial plateau fractures. One problem with synthetic compounds like CPCs and the rapid development seen over the last decade, is that the technical development of new compositions and products is much faster than the clinical evaluation process. This means that, by the time valid clinical studies have been completed, the front line for the development of new materials will have already advanced way ahead of the products being examined in clinical trials. One limitation for all CPCs is the relatively low mechanical strength in shear, bending and tension. In order to improve the overall mechanical strength, and then bending and shear in particular, different types of fibre reinforcement have been studied (Gorst et al. 2006, Buchanan et al. 2007, Pan et al. 2007). In several experiments it has been shown that such a cement–fibre composition caused a very significant improvement especially in the flexural strength, but also in the modulus of elasticity. It seems reasonable to believe that, in the near future, we will see not only pure CPCs but also composite products with such fibre reinforcement as might be required to improve the ultimate strength. Conventional CPC has small pores which means that resorption will occur as a surface phenomenon. In certain indications it might be desirable to have a compound with larger pores even though larger pores will also reduce the mechanical strength of the cement. The presence of larger pores will provide a larger surface exposed to various biological activities, and macropores will therefore, for instance, increase the speed of resorption. Different techniques have been tried to obtain calcium phosphate compounds with macropores (Nilsson et al. 2004). One such technique is the inclusion of soluble particles in the cement that will dissolve quickly leaving large pores. Another technique is to combine calcium phosphate with

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calcium sulphate. As the calcium sulphate will dissolve and disappear much faster than the phosphate, macropores will be formed within the calcium phosphate. In the future, different new compounds aiming to create CPCs with macropores intended for specific indications will most probably be available. One of the most attractive features of calcium phosphate compounds, apart from providing mechanical strength when used for fracture fixation, is the potential use for controlled release of therapeutic or bioactive agents (Weir and Xu 2008). Several experimental studies have shown the potential use for delivery of, for instance, antibiotics and growth factors, drugs of great potential interest when used in fracture treatment (Kisanuki et al. 2007, Niikura et al. 2007, Stallmann et al. 2008, Tamini et al. 2008).

17.7

References

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