THE EFFECT OF GLASS-CERAMIC (GC) FILLED POLY(METHYL METHACRYLATE) BONE CEMENT
COMPOSITES
HAMIZAH BINTI ABDUL SAMAD
UNIVERSITI SAINS MALAYSIA
2011
THE EFFECT OF GLASS-CERAMIC (GC) FILLED
POLY(METHYL METHACRYLATE) BONE CEMENT COMPOSITES
by
HAMIZAH BINTI ABDUL SAMAD
Thesis submitted in fulfillment of the requirements for the degree of
Master of Science
June 2011
DECLARATION
I hereby declare that the thesis entitles “The Effect of Glass-Ceramic (GC) Filled Poly (Methyl Methacrylate) Bone Cement Composites” submitted for the Master of Science degree at the Universiti Sains Malaysia is my original work, except where otherwise stated. It also has not been previously submitted by me at another University for any degree.
Candidate’s Name : Hamizah Bt Abdul Samad
Signature :
Date :
Supervisor’s Name : Assoc. Prof. Dr. Ir. Mariatti Jaafar @ Mustapha
Signature :
Date :
ACKNOWLEDGEMENTS
First of all, I would like to express my sincere appreciation to my supervisor, Assoc. Prof. Dr. Ir. Mariatti Jaafar @ Mustapha for her supervision, concern and support to complete this research. I would like to convey my gratitude to Prof. Radzali Othman for his kindly advice and help throughout this research. I also want to express my special thanks to Dr Masakazu Kawashita from Tohoku University, Japan for his sharing respectful experiences in research with me regarding my lab work.
My special thanks are also extended to School of Materials and Mineral Resources Engineering and the Dean, Prof. Ahmad Fauzi Mohd Noor and all support staffs for preparing me a lot of technical and practical courses. I would like to express my great thanks to the technical staffs especially to Mr. Sharul Ami, Mr. Kemuridan, Mr.
Abdul Rashid, Mr. Mokhtar, Mr. Khairi, Mr. Faizal and Madam Fong Lee Lee for their assistance and technical support during the process of completing the research.
This appreciation is also dedicated to all my friends for their unlimited help and moral support. I send my deepest gratitude for my family for their love, pray and continued support. I would like to dedicate all my gratefulness to my lovely husband, Mohd Fauzi Alias, for his endless love, care and fully support to finish my M.Sc.
Finally, a great thanks to the Ministry of Science, Technology & Innovation (MOSTI) and Universiti Sains Malaysia for providing me scholarship, research grant and good facilities during my research study.
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TABLE OF CONTENTS
Page
ACKNOWLEDGEMENTS ii
TABLE OF CONTENTS iii
LIST OF FIGURES viii
LIST OF TABLES xii
LIST OF ABBREVIATIONS xiv
LIST OF SYMBOLS xvi
ABSTRAK xviii
ABSTRACT xix
CHAPTER 1: INTRODUCTION
1.1 Background Study and Problem Statement 1
1.2 Objectives 4
1.3 Outline of Thesis Structure 4
CHAPTER 2: LITERATURE REVIEW
2.1 Biomaterials 6
2.1.1 Introduction 6
2.1.2 Classification of Biomaterials: Based on Types of Biomaterials 9
2.1.2.1 Metallic Biomaterials 9
2.1.2.2 Polymer Biomaterials 10
2.1.2.3 Ceramic Biomaterials 12
2.1.2.4 Composite Biomaterials 12
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2.2 Bioceramics 13
2.2.1 Introduction 13
2.2.2 Classification of Bioceramics 14
2.2.3.1 Bioinert Ceramics 14
2.2.3.2 Bioresorbable Ceramics 15
2.2.3.3 Bioactive Ceramics 15
2.2.3 Applications of Bioceramics 16
2.3 Bioactive Glass and Glass-ceramics 19
2.3.1 Bioactive Glasses 20
2.3.2 Bioactive Glass-ceramics 21
2.3.2.1 Glass-ceramic Processing 21 2.3.2.2 Properties of Glass-ceramics 22 2.3.2.3 Commercial Bioactive Glass-ceramics 23 2.3.2.4 Mechanism of Bioactive Bonding 25 2.4 Bone Cement as Polymer Biomaterials 27
2.4.1 History 27
2.4.2 Functions and Compositions 28
2.4.3 Curing of Bone Cement 30
2.4.4 Viscosity and Handling Properties 33 2.4.5 Mechanical Properties of Bone Cement 35 2.5 Current Research on Incorporation of Bioactive Fillers into 39
PMMA Bone Cement
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CHAPTER 3: RAW MATERIALS AND METHODOLOGY
3.1 Raw Materials 44
3.1.1 Glass-ceramic materials 44
3.1.2 PALACOS®LV Bone Cement 44
3.1.3 Hydroxyapatite (HA) 46
3.2 Methodology 46
3.2.1 Fabrication of Bioactive Glass-ceramic 46 3.2.1.1 Batching of Glass Ceramic Composition 46
3.2.1.2 Melting 46
3.2.1.3 Heat treatment 47
3.2.2 Preparation of PMMA Bone Cement Composites 47 3.2.2.1 Mixing and Casting Process 47 3.2.2.2 Measurement of Temperature Changes 48
3.3 Testing and Characterization 49
3.3.1 Physical Characterization 49
3.3.1.1 Particle Size Distribution 49 3.3.1.2 Density Measurement (ASTM 2792-00) 49 3.3.1.3 Scanning Electron Microscopy (SEM) & Energy
Dispersive X-ray (EDX) Analysis 50
3.3.1.4 X-ray Diffraction (XRD) 51 3.3.1.5 X-ray Fluorescence (XRF) 52
3.3.2 Thermal Characterization 53
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3.3.2.1 Differential Scanning Calorimetry and Thermogravimetry
Analysis (DSC/TGA) 53
3.3.2.2 Dynamic Mechanical Analysis (DMA) 53
3.3.3 Mechanical testing 54
3.3.3.1 Single Edge Notch Bending Test (ISO 13586:2000) 54 3.3.3.2 Flexural Testing (ASTM D790-03) 55
3.3.3.3 Vickers Hardness 57
3.3.4 Evaluation of Bioactivity In Vitro Behavior 57
3.3.4.1 Preparation of SBF 58
3.3.4.2 Soaking in SBF 59
CHAPTER 4: RESULTS AND DISCUSSIONS
4.1 Introduction 60
4.2 Fabrication of Bioactive Glass-ceramic 60
4.2.1 Chemical Composition 60
4.2.2 Thermal Behavior 62
4.2.3 Phase Identification 63
4.2.4 Morphology 66
4.2.5 Evaluation of Physical Properties 68
4.2.5.1 Vickers Hardness 68
4.2.5.2 Density 69
4.2.6 Bioactivity Test 70
4.3 Characterization of Particulate Fillers and PMMA Powder 76 4.3.1 Particle Size and Particle Size Distribution Analysis 76
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4.3.2 Morphology Analysis 77 4.4 Properties of PMMA Bone Cement Composite 79
4.4.1 Density on Composite 79
4.4.2 Setting Properties 80
4.4.2.1 Peak Temperature 82
4.4.2.2 Setting Time 84
4.4.2.3 Dough Time 86
4.4.3 Mechanical Properties of PMMA Bone Cement Composite 87
4.4.3.1 Flexural Properties 87
4.4.3.2 Single Edge Notch-Bending (SEN-B) 98
4.4.4 Thermal Properties 101
4.4.4.1 Thermogravimetric Analysis (TGA) 101 4.4.4.2 Dynamic Mechanical Analysis (DMA) 103 4.5 Bioactivity of Bone Cement Composites 107
CHAPTER 5: CONCLUSIONS AND RECOMMENDATIONS
5.1 Conclusions 112
5.2 Recommendations 114
REFERENCES 115
APPENDICES 127
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LIST OF FIGURES
Page Figure 2.1 Clinical uses of bioceramics 18 Figure 2.2 The SiO2-CaO-Na2O phase diagram 20 Figure 2.3 Temperature-time cycle for a glass-ceramic 22 Figure 2.5 Sequence of interfacial reactions involved in forming a bond
between bone and bioactive glasses 25 Figure 2.5 Schematic diagram of the component of a cemented total
hip joints replacement. ABC: Antiobiotic loaded bone cement 28 Figure 2.6 Polymerization of PMMA by an addition reaction. Note
that the MMA monomer reacts with a radical to form a secondary radical that can attack the double bond of
another MMA monomer 32
Figure 2.7 Illustration of exothermic temperature changes occurring in
acrylic bone cement during the setting process 35 Figure 3.1 Schematic illustration of SEN-B specimen used for
fracture toughness testing 54
Figure 4.1 Compositional dependence of apatite formation on the surfaces
of glasses in the Na2O-CaO-SiO2 system 61 Figure 4.2 TG/DSC curve of glass powder 63 Figure 4.3 XRD patterns of glass before heat treatment, revealing an
amorphous structure 64
Figure 4.4 XRD patterns of glass-ceramic after heat treatment at
850, 950, and 1000 oC 65
Figure 4.5 SEM micrographs of surface structure of glass-ceramic samples heat treated at different temperatures; (a) 850oC (b) 950oC and (c) 1000oC (W denoted as CaSiO3 phase
and S is Na2Ca3Si6O16 phase) 67
Figure 4.5 continuation 68
Figure 4.6 Vickers hardness of glass-ceramic heat treated at various
temperatures; 850, 950, and 1000oC 69 viii
Figure 4.7 Density of glass-ceramic heat treated at various temperatures;
850, 950 and 1000oC 70
Figure 4.8 SEM images of the glass-ceramic surface samples heat treated at 950oC (a, b) before immersion (c, d) after
immersion in SBF for a period of 1- day (e, f) 3- day and (g, h) 7- day at magnifications of 1 K X (a, c, e, g) and
10K X (b, d, f, h) 72
Figure 4.8 continuation 73
Figure 4.9 EDX analysis of chemical composition for glass-ceramic heat treated at 950°C; (a) before immersion (b) after
immersion in SBF for a period of 1 day (c) 3 day and (d) 7 day 74 Figure 4.10 TF-XRD patterns of glass-ceramic heat treated at 950°C
before and after immersion in SBF for 7 days 75 Figure 4.11 SEM micrographs of (a) PMMA microsphere particles with
magnification of 500X (b) HA particles with magnification of 50K X (c) glass-ceramic particles with magnification
of 6K X 78
Figure 4.11 continuation 79
Figure 4.12 Time-temperature graphs obtained for the control sample, glass ceramic- filled bone cement (GCBC) and
hydroxyapatite-filled bone cement (HABC) composite
for filler loading of 4 and 16 wt% during setting reactions 82 Figure 4.13 The effect of filler loading on the peak temperature of the
samples tested during setting reaction 83 Figure 4.14 The effect of filler loading on the setting time of the samples
tested during setting reactions 86 Figure 4.15 The effect of filler loading on the dough time of the samples
tested during setting reactions 87 Figure 4.16 Stress versus strain of selected PMMA bone cement
composites compared to a control sample 88
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Figure 4.17 Effect of filler loading on the flexural modulus of GCBC
and HABC samples at 4 to 16 wt% of fillers 90 Figure 4.18 Effect of filler loading on the flexural strength of GCBC
and HABC samples at 4 to 16 wt% of fillers 92 Figure 4.19 (a) SEM micrograph of fracture surfaces of commercial
PMMA bone cement at magnifications 100X (b) at
magnifications 500X (c) EDX result of ZrO2 93
Figure 4.19 continuation 94
Figure 4.20 SEM micrograph: (a) of fracture surfaces of GCBC4 and
(b) fracture surface of GCBC16 at magnifications 100 X 95 Figure 4.21 SEM micrograph: (a) of fracture surfaces of HABC4 and
(b) fracture surface of HABC16 at magnifications 100 X 96 Figure 4.22 SEM micrograph of particles dispersion at: (a) GCBC16
and (b) HABC16 with magnification of 1.00K X 97 Figure 4.23 SEM micrograph of particles agglomerations at:
(a) GCBC16 and (b) HABC16 with magnification of 1.00K X 98 Figure 4.24 Effect of filler loading on the fracture toughness of GCBC
and HABC samples at 4 to 16 wt% of fillers 99 Figure 4.25 SEM micrograph of ZrO2 particle shows the debonding
with the matrix PMMA 100
Figure 4.26 TGA analysis of control, glass-ceramic filled bone cement (GCBC) and hydroxyapatite filled bone cement (HABC)
with 4 and 16 wt% filler loadings 103 Figure 4.27 Storage modulus as a function of temperature for commercial
PMMA bone cement (control), GCBC and HABC composites
with filler content 4 and 16wt% 105
Figure 4.28 Tan delta as a function of temperature for commercial PMMA bone cement (control), GCBC, and HABC composites with
filler content 4 and 16 wt% 107
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Figure 4.29 SEM images of the GCBC samples with filler loading (a) 4 wt% (b) 8 wt% after soaking in SBF solution for 7
days at a magnification of 1K X 109
Figure 4.30 TF-XRD patterns of GCBC4 and GCBC8 cement samples
after immersion in SBF for 7 days 110
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LIST OF TABLES
Page
Table 2.1 Classification of biomaterials types in medical devices
and dental with example of their applications 8
Table 2.2 Consequences of implant-tissue interactions 14 Table 2.3 Form, phase and function of bioceramics 17 Table 2.4 Composition (wt%) and mechanical properties of
bioactive glasses 19
Table 2.5 Mechanical properties of glass-ceramics 23 Table 2.6 Compositions of some bioactive glass-ceramics 24 Table 2.7 Chemical compositions of six commercial formulations of
bone cement 30
Table 2.8 Mechanical properties of human bone, PMMA
and prosthesis material 36
Table 2.9 Summary of relevant mechanical properties of acrylic
bone cements 37
Table 3.1 List of raw chemicals used for preparation of glass-ceramic 44 Table 3.2 Composition of PALACOS® LV bone cement 45 Table 3.3 List of raw chemicals in PALACOS® LV bone cement
with its functions 45
Table 3.4 Mixing composition of particulate filler-filled PMMA bone cement 48 Table 3.5 Nominal ion concentration of SBF compared to those
in the human blood plasma 58
Table 3.6 Reagents for preparation of SBF (pH 7.25) 58 Table 4.1 Calculated composition of raw chemicals according to a batch
of 100g of glass powder 61
Table 4.2 Chemical compositions of glass sample (wt%) 62 Table 4.3 Mean particles size of PMMA powder, HA and glass-ceramic
particle analyzed by HELOS particle size analyzer 76 Table 4.4 Effect of filler loading on the density of the PMMA
bone cement and bone cement composite 80 xii
Table 4.5 Values of the peak temperature, setting time, and dough time for PMMA bone cement modified with different
amounts of glass-ceramic and HA fillers 84 Table 4.6 Measurement of the average flexural strength and flexural
modulus of the samples tested 89
Table 4.7 TGA result of bone cement composite filled with
glass-ceramic and HA particles 102 Table 4.8 Tan delta and Tg values of the PMMA bone cement composite 107
Table 4.9 Summary results of PMMA bone cement composite
studies in previous works 111
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LIST OF ABBREVIATIONS
AW-GC : Apatite-wollastonite glass-ceramic BPO : Benzoyl peroxide
DMA : Dynamical mechanical analysis DMPT : N, N-dimethyl-4-toluidine DSC : Differential scanning calorimetry EDX : Energy dispersive x-Ray
FTIR : Fourier transform infrared GCBC : Glass-ceramic-filled bone cement
HA : Hydroxyapatite
HABC : Hydroxyapatite-filled bone cement HCA : Hydroxycarbonate apatite
HQ : Hydroquinone
L/P : Liquid to powder MMA : Methyl methacrylate MMT : Montmorillonite
MPS : α-methacryloxypropyltrimethoxysilane MWCNT : Multi walled carbon nanotube
NIH : National Institutes of Health
PC : Polycarbonate
PE : Polyethylene
PFF : Poly (propylene fumarate) PGA : Poly (glycolic acid) PLA : Poly (lactic acid)
PMMA : Poly(methyl methacrylate)
PP : Polypropylene
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PS : Polystyrene PVC : Polyvinyl chloride P/L : Powder to liquid SBF : Simulated body fluid
SEM : Scanning electron microscopy SEN-B : Single edge notch bending test SG-P : Silica glass powder
SI : Standard International SiC : Silicon carbide
TCP : Tricalcium phosphate TF-XRD : Thin film x-ray diffraction XRD : X-ray diffraction
XRF : X-ray fluorescence
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LIST OF SYMBOLS
% Percent
µm Micrometer
g Gram
g/cm3 Gram per centimeter cubic
GPa Gigapascal
h Hour
Hz Hertz
inHg Inches of mercury kcal Kilo calorie
kJ Kilo joule
kmol/ml Kilomole per milliliter
kV Kilovolt
L Litre
M Molarity
mA Milliampere
mg Milligram
min Minutes
ml Milliliter
mm Millimeter
mM Millimolar
mm/min Millimeter per minute mm2 Millimetre square mol % Mol percent MPa Megapascal
MPa.m1/2 Megapascal meter square xvi
N Newton
nm Nanometer
Nm Newton meter
oC Degree celcius
oC/min Degree per minute ppm Parts per million psi Pound per square inch
T Temperature
T Time
tamb Ambient temperature td Dough time
Tp Peak temperature tset Setting time
wt% Weight percent
θ Theta
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KESAN SERAMIK KACA (GC) TERISI KOMPOSIT POLIMETIL METAKRILAT (PMMA) SIMEN TULANG
ABSTRAK
Dalam kajian ini, komposisi seramik kaca telah dihasilkan berdasarkan kepada sistem kaca Na2O-CaO-SiO2 dan ia telah digunakan sebagai pengisi di dalam komersil simen tulang PMMA (PALACOS LV®). Dalam penghasilan serbuk seramik kaca, pertamanya serbuk kaca yang terhasil di analisa menggunakan DSC/TGA dan XRF, kemudian ia dipadatkan dan dirawat haba pada suhu antara 850 hingga 1000 oC. Keputusan XRD bagi seramik kaca yang dirawat haba pada suhu 950 oC telah menunjukkan sifat kristal wollastonite (CaSiO3) dan sodium kalsium silikat (Na2Ca3Si6O16) yang tinggi. Ia juga menunjukkan kebioaktifan yang tinggi, yang mana ia menghasilkan lapisan apatit selepas direndam di dalam SBF selama 7 hari. Kemudian, seramik kaca yang dirawat haba pada suhu 950oC digunakan sebagai pengisi di dalam simen tulang PMMA dengan 0, 4, 8, 12 dan 16 % berat pengisi dan keputusannya dibandingkan dengan komposit simen tulang terisi HA. Kesan pengisi terhadap sifat pengesetan, mekanikal dan terma telah dikaji. Didapati, suhu puncak dan masa doh simen tulang semasa pempolimeran menurun dengan meningkatnya peratus berat pengisi. Walaubagaimanapun, masa pengesetan tidak memberikan sebarang kesan dengan peningkatan peratus berat pengisi.
Keputusan menunjukkan kekuatan lenturan dan keliatan patah menurun, manakala modulus lenturan meningkat dengan meningkatnya peratus berat pengisi. Selain itu, kestabilan terma, Tg dan modulus penyimpanan komposit simen meningkat dengan peningkatan bahan pengisi. Kajian morfologi ke atas bioaktiviti simen komposit menunjukkan pertumbuhan apatit di atas permukaan sampel GCBC4 dan GCBC8.
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THE EFFECT OF GLASS-CERAMIC (GC) FILLED
POLY(METHYL METHACRYLATE) BONE CEMENT COMPOSITES
ABSTRACT
In this study, a composition of glass-ceramic was fabricated based on the Na2O-CaO- SiO2 glass system and was used as filler in commercial PMMA bone cement (PALACOS LV®). In producing the glass-ceramic powder, firstly the glass powders were analyzed using DSC/TGA and XRF, then it was compacted and heat treated at temperatures between 850 to 1000oC. XRD result of glass-ceramic heat treated at 950oC shows high crystallization of wollastonite (CaSiO3) and sodium calcium silicate, (Na2Ca3Si6O16) in the glass composition. It also exhibits a high bioactivity which formed apatite after soaking in SBF for 7 days. Next, glass-ceramic heat treated at 950oC were used as a filler in the PMMA bone cement with filler loading of 0, 4, 8, 12, or 16 wt%
and compared with HA composites. The effect of filler loadings on the setting, mechanical, and thermal properties were evaluated. It is found that the peak temperature and dough time during the polymerization of bone cement decreased with increasing filler loading. However, setting time did not show any significant trend. Result shows the flexural strength and fracture toughness decreased, and the flexural modulus increased as the filler loading increased. Besides, the thermal stability, Tg and storage modulus of cement composite increased with increasing filler loading. Morphological studies of the bioactivity of cement composite revealed the growth of apatite deposited on the GCBC4 and GCBC8 surface sample.
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CHAPTER 1 INTRODUCTION
1.1 Background Study and Problem Statement
Self-curing polymethylmethacrylate (PMMA) bone cements have been in the market for more than 50 years since their introduction by Sir John Charnley in 1958 (Charnley, 1960). It was first used in dental applications followed by the use in orthopaedic surgery for the fixation of total joint replacement such as for hip and knee prosthesis. In orthopaedics surgery, PMMA bone cement functions to transfer body weight and service loads from the prosthesis to the bone. PMMA bone cement has also been used to increase the load carrying capacity of the prosthesis-bone cement-bone system (Lewis, 1997; Kuehn et al., 2005a). Commercial bone cements are prepared by mixing powder and liquid components with proportion of powder to liquid (P/L) equal to 2. The powder component consists of PMMA or PMMA-based copolymers, and a polymerization initiator, usually benzoyl peroxide (BPO). The liquid component consists of methyl methacrylate (MMA) monomer, accelerator (usually N-N-dimethyl- p-toluidine (DMPT)) and hydroquinone (HQ) as an inhibitor (Lewis, 1997; Hasenwinkel, 2004; Kuehn et al., 2005a). In the operation theatre, the powder and liquid parts are mixed for 2-3 minutes until a dough mixture is obtained and then applied to the desired bone cavity. Due to a rapid polymerization reaction, bone cement hardens in the ensuing 3-5 minutes (Serbetci et al., 2002).
The main adverse effect of bone cement application is a strongly exothermic reaction at the bone and cement interface during the setting period. Maximum
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temperatures in the range of 80 oC to 124 oC have been reported and these values could damage living tissue (Pascual et al., 1996). In addition, bone-PMMA bone cement interface is known as one of the weak-link zones in the prosthesis-bone cement-bone construct because it does not bind or adhere to bone and has poor mechanical properties.
The lack of ability to bind to bone sometimes results in the widening of the intervening fibrous tissue layer between bone and PMMA cement, causing aseptic loosening of the cement (Shinzato et al., 2000; Kamimura et al., 2002). On the other hand, PMMA has been demonstrated to be biocompatible and easy to shape in vivo, allowing its use as a bone substitute in reconstructive surgery of the knee and in vertebroplasty. However, high shrinkage during curing, and the release of monomer to the surrounding tissue and again, the ability to bond directly to bone, pose several potential risks that lead to prosthesis loosening with time due to tissue necrosis, interfacial failure, and cement failure (Goto et al., 2005).
Therefore, in an effort to improve their mechanical, thermal, handling and biocompatibility properties, investigations have been carried out on many different types of bone cements. Various approaches have been proposed and reported in the literature and one of them is bioactivation of PMMA bone cement by the incorporation of bioactive fillers in bone cement. The introduction of a bioactive phase in the PMMA matrix was suggested in order to enhance the quality of the bone-cement interface and to improve the setting and mechanical properties of the cement (Gilbert et al., 1995; Dalby et al., 2002). The in vivo studies of Kwon et al. (1997) found that there is new bone formation adjacent to the interface between the implant and surrounding bone as the amount of hydroxyapatite (HA) particles is increased. They also found that the
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interfacial shear strength of the implanted specimens has a significant increase compared with the cement without HA.
Goto et al. (2008) reported that when using titania as filler in PMMA bone cement, lower peak temperature than for the unfilled cement were obtained. Besides HA and titania, Fujita et al. (1998) evaluated the bonding strength of the bioactive bone cements with higher percentage of apatite-wollastonite glass-ceramic powder. They found that bioactive bone cement had a higher bonding strength after surgery. The rationale for incorporating bioactive filler into PMMA cement had been also reported by Vallo (2000), and Dalby et al. (2002). From the literature, the cements showed good mechanical properties and excellent osteoconductivity by forming a biologically active bone-like apatite layer on their surfaces. However, trials using various fillers in bone cement produced unsatisfactory result due to deterioration of the mechanical properties after adding large weight percent (wt%) of the bioactive particles that caused difficulty in handling of the bone cement. The lack of bioactivity of the composite cement was also affected when the wt% of added bioactive particles is too small (Mousa et al., 2000).
In this study, a glass-ceramic composition (55SiO2, 35CaO, 10Na2O and 3P2O5 (wt%)) was developed and characterized. Trials to incorporate this glass-ceramic particle as filler into commercial PMMA bone cement (PALACOS® LV) that possesses favorable physical, mechanical, thermal and bioactivity properties was carried out.
Different weight percent (wt%) of the fillers were used and as compared, incorporation of commercially HA filler into PMMA bone cement also being investigated in this study.
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1.2 Objectives
The objectives of project are listed as below:
1) To evaluate bioactivity of glass-ceramic filler in PMMA bone cement.
2) To study the effect of the incorporation different weight percent of glass- ceramic and HA fillers on the setting, mechanical, thermal and bioactivity properties of PMMA bone cement composites.
1.3 Outline of Thesis Structure
Chapter 1:
Introduction of PMMA bone cement and problem statement has been briefly explained in this chapter. The objectives of the study also have been stated.
Chapter 2:
This chapter reviews the literature on biomaterials and bioceramics field. In addition, literature on PMMA bone cement as polymer biomaterials and highlights on various studies and published works on incorporation of bioactive fillers into PMMA bone cement has been summarized in this chapter.
Chapter 3:
This chapter describes the detail of raw materials, chemicals and equipments that have been used to synthesize glass-ceramic and PMMA bone cement composites.
Experimental and characterization methods have been explained in this chapter.
4
Chapter 4:
Chapter 4 consists of results from the experiments and presented in charts, tables and micrographs. The results obtained from the experiments have been evaluated and discussed thoroughly.
Chapter 5:
Several conclusions of the present study are discussed in this chapter and a few suggestions and recommendations are proposed for future studies.
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CHAPTER 2 LITERATURE REVIEW
2.1 Biomaterials 2.1.1 Introduction
According to Black (1992) biomaterials can be defined as a material used in a medical device, intended to interact with biological systems. Over the years, various definitions of biomaterials have been proposed. For example, a biomaterial can be simply defined as a synthetic material used to replace part of a living system or to function in intimate contact with living tissue (Park & Bronzino, 2002). The other definition most commonly accepted is from the National Institutes of Health (NIH) which describes a biomaterial as:
“any substance (other than a drug) or combination of substances, synthetic or natural in origin, which can be used for any period of time, as a whole or as part of a system which treats, augments, or replaces any tissues, organ, or functions of the body”
(Williams, 1987).
A material that can be used for medical application must possess a lot of specific characteristics, of which the first and foremost requirement is biocompatibility.
Biocompatibility is the ability of a material to perform with an appropriate biological host response in a specific application (Williams, 1987). It means that, it should be non- toxic and non-carcinogenic, cause little or no foreign-body reaction, and be chemically stable and corrosion resistant. The biomaterial also should possess adequate physical and mechanical properties to serve as augmentation or replacement of body tissues. For
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practical use, a biomaterial should be able to formed or machined into different shapes, relatively cheap, and be readily available.
Biomaterials have been widely used in application such as (Davis, 2003):
(1) orthopaedics – total joint replacements (hip, knee), bone cements, bone void fillers, fracture fixation plates, and artificial tendons and ligaments;
(2) cardiovascular applications - heart valves, pacemakers, artificial heart and ventricular assist device components, stents, and blood substitutes;
(3) ophtalmics – contact lenses, corneal implants and artificial corneas, and intraocular lenses;
(4) other applications- dental implants, cochlear implants, tissue screws and tacks, burn and wound dressings and artificial skin, tissue adhesives and scalants, drug- delivery systems, and sutures.
In general, biomaterials can be broadly categorized into the following categories:
metals, ceramics, polymers, and composites. Table 2.1 illustrates some of the biomaterials types and their applications for these four groups of synthetic materials used for implantation.
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Table 2.1: Classification of biomaterials types in medical devices and dental applications (Binyamin et al., 2006; Davis, 2003)
Classification Biomaterial Examples of applications
Metal 316L stainless steel Surgical instruments, orthopedic fixation devices, stents
Ti and Ti-containing alloys Fracture fixation, pacemaker encapsulation, joint replacement
Nickel-Titanium Alloy
(Nitinol)
Stents, orthondotic wires, bone plates Platinum and platinum-
containing alloys
Electrodes Polymer Polytetrafluoroethylene
(Teflon, Gore-Tex) Vascular grafts, catheters, introduces Poly(ethylene terephthalate)
(polyester, Ethibond, Dacron)
Vascular graft, drug delivery, non- resorbable sutures
PMMA Bone cement, intraocular lenses, dental restorations
Polyurethane Cathethers, tubing, wound dressing, heart valves, artificial hearts
Silicone rubber
(polydimethylsiloxane)
Cathethers, feeding tubes, drainage tubes, introduces tips, flexible sheaths, gas exchange membranes
Polycarbonate Major component in renal dialysis cartridge, heart-lung machine,trocars, tubing interconnectors
Hydrogels (poly(ethylene
oxide)), poly(ethylene glycol), poly(vinyl alcohol), etc.)
Drug delivery, wound healing, hemostasis, adhesion prevention, contact lenses, extracellular matrices, reconstruction
Polyamides (nylon) Non-resorbable sutures
Polypropylene (i.e., prolene) Non-resorbable sutures, herni mesh Ceramic
and glasses Alumina Joint replacement, dental implants, orthopaedic prostheses
Carbon Heart valves, biocompatible coatings, electrodes, dental implants
Hydroxyapatite Implant coatings, bone filler Bioglass Metal prosthesis coating, dental
composites, bone cement fillers
Porcelain Dental restorations
Composites BIS-GMA-quartz/silica filler PMMA-glass filler
Dental restorations
Dental restorations (dental cements)
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2.1.2 Classification of Biomaterials: Based on Types of Biomaterials 2.1.2.1 Metallic Biomaterials
Metals are inorganic materials that have unique atomic arrangements and bonding characteristics leading to enhanced mechanical, thermal and electrical properties. Their excellent electrical and thermal conductivity, fair biocompatibility and mechanical properties like high stiffness, high ductility and good wear resistance make them very ideal for a variety of medical applications especially for load bearing properties (Binyamin et al., 2006). One of the advantages of using metals as biomaterials is their availability and relative ease of processing from raw ore to finished products.
Although they have excellent mechanical properties, metallic materials can have serious corrosion problems in an in vivo environment. The consequences of corrosion are the disintegration of the implant material per se, which result in releasing toxic metal ions to the body and also weakening the implants. Thus, corrosion resistance is a primary criterion in selecting metals for biomedical implants (Desai et al., 2008; Donglu, 2006).
Metallic biomaterials have been used mainly for the fabrication of medical devices for the replacement of hard tissue such as total hip and knee prostheses and for fracture healing aids such as bone plates and screws, pins and spinal fixation devices.
Besides orthopaedic, there are other markets for metallic implants and devices, including oral and maxillofacial surgery and dental implants (Niinomi, 2008). Some metals have also been used for repairing soft tissues as part of cardiovascular surgery as vascular stents, as pacemaker leads, and catheter guide wires. Besides that, surgical instruments, dental instruments, needles, staples, and implantable drug pump housings are also made from metallic materials (Davis, 2003).
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Metallic biomaterials have the longest history among the various biomaterials.
The main material groups that dominate biomedical metals are stainless steel, cobalt- based alloy, titanium alloys, and shape memory alloys such as nickel-titanium alloy known as nitinol (Pelton et al., 2000; Niinomi, 2002; Bartel et al., 2006; Frosch &
Sturmer, 2006). Generally, these materials are popular primarily because of their ability to bear significant loads, withstand fatigue loading, and undergo plastic deformation prior to failure. They also exhibit good biocompatibility, which does not cause serious toxic reactions in the human body.
2.1.2.2 Polymer Biomaterials
Polymers are the most widely used materials in biomedical applications. They have addressed neurological, cardiovascular, ophthalmic, and reconstructive pathologies with implantable devices designed to sustain or enhance human life. They have also been found useful in temporary therapies such as hemodialysis and coronary angioplasty. In addition, polymers are also used extensively in dentistry as composite (resin-ceramic), implants, dental cements, and denture bases and teeth (Davis, 2003).
The advantage of using polymers as biomaterials, is their manufacturability. Polymers are easy to fabricate into various sizes and shapes (rod, film, fiber, sheet, etc) compared to metals and ceramics. They are also light in weight and have a wide range of mechanical properties for different applications. The range of polymer biomaterials applications can be classed into types; synthetic and natural polymers (Donglu, 2006).
Synthetic polymers are the majority of the polymer biomaterials that have been widely used in making various medical devices, such as disposable supplies, implants,
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drug delivery systems and tissue engineering scaffolds. Synthetic polymers, then can be divided into two types: synthetic non-biodegradable polymers and synthetic biodegradable polymers. Although most synthetic non-biodegradable polymers were originally developed for non-biomedical uses, they are widely used as biomaterials mainly because of the necessary physical-mechanical properties they have. There are still no newly engineered biomaterials that can replace those non-degradable polymers.
A good example is PMMA bone cement which has been used for fixation of artificial joint since 1943 and is still being widely used clinically nowadays (Kuehn, 2005).
Example of others non-biodegradable polymers include polyvinyl chloride (PVC), polyethylene (PE), polypropylene (PP), polystyrene (PS), polycarbonate (PC), polyesters, polyamides (nylon), polyurethanes, and polysiloxanes (silicone) (Donglu, 2006).
Synthetic biodegradable polymers have attracted much attention in the last decade because they offer the advantage of being able to be eliminated from the body after fulfilling its intended use. Therefore, the second surgery can be avoided. This polymer is becoming more and more important in biomaterials and for the regeneration of tissues and organs. Example of this kind polymers include polyamino acid, poly (propylene fumarate) (PFF) and aliphatic polyester, such as poly (glycolic acid) (PGA), and poly (lactic acid) (PLA) (Donglu, 2006).
Commonly encountered natural polymers are proteins, collagen, chitin and chitosan, hyaluronic acid, heparin and DNA. These materials are used as biomaterials largely because their structures are similar to the human tissue they intend to replace.
These are important classes of biomaterials because of their biodegradation 11
characteristics and they are easily to find abundantly. However, the use of naturally occurring polymers often has some problem that provokes immune reaction of the host tissue. Therefore, many of them have to be chemically modified before being used as biomaterials.
2.1.2.3 Ceramic Biomaterials
Ceramics are non-metallic, refractory, polycrystalline compounds and usually inorganic material, which have some typical properties which are extremely hard, chemically stable, good wear resistance, and high durability that make them good materials as inert materials and useful for medical applications. But, ceramics are limited by their relative brittleness, high melting temperature and low electrical and thermal conductivity. Examples of ceramics include silicates, metallic oxides, carbides, sulfides, refractory hydrides, selenides and carbon structures such as diamond, graphite and pyrolized carbons. They are produced under a high temperature heat treatment process called firing. Ceramics used for the body are called bioceramics. Bioceramics used in fabricating implants typically can be classified as inert, bioactive and biodegradable or resorbable (Billotte, 2003; Binyamin et al., 2006; Navarro et al., 2008). The details of these bioceramics materials will be discussed in Section 2.2
2.1.2.4 Composite Biomaterials
Composite materials are combinations of two or more distinct constituent materials or phases on a macroscopic scale and in which mechanical properties are significantly altered in comparison with the homogenous constituents (Lakes, 1993).
Composite materials offer some advantages which include control over material bulk 12
properties and improvements in surface properties. The bulk properties of composite materials depend upon the volume fraction and the shape of the heterogenities. The principal inclusion shape categorized as the particle, fiber, and lamina. Particles and fiber reinforcements have been used to improve properties of biomaterials. For example, rubber used in catheters, where rubber gloves are usually reinforced with very fine particles of silica to make the rubber stronger and tougher. In dental composite materials, glasses or ceramic particles are blended in a polymeric organic resin matrix with interfacing silane coupling agents. Composite such as graphite fibers in epoxy resin can be as strong as steel when loaded in the fiber direction but much lighter. However, this material is compliant when loaded transversely to the fibers (Bhat, 2005).
2.2 Bioceramics 2.2.1 Introduction
Park (2008) stated that bioceramics are ceramic materials that are used to make devices for the replacement, repair and reconstruction of diseased, damaged or “worn out” parts of living systems or to function in intimate contact with living tissues. In general, bioceramics show better biocompatibility with tissue response compared to polymer or metal biomaterials (Bilotte, 2003). Other than biocompatibility, ceramic materials have the following excellent properties: (a) non-toxic, (b) non-carcinogenic, (c) non-allergic, (d) non-imflammatory, and (e) biofunctional for its lifetime in the host.
However, despite the excellent biocompatibility of bioceramics, the problems that occur in conventional ceramics also exist in bioceramics. The primary drawbacks of bioceramics are their brittleness, low strength, and inferior workability. Consequently,
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bioceramics are very sensitive to notches or microcracks because they do not deform plastically (Bilotte, 2003).
2.2.2 Classification of Bioceramics
In general, bioceramics can be classified into three types based on their tissue response in the body. These are bioinert, bioactive, and bioresorbable (Thamaraiselvi &
Rajeswari, 2004). The implant – tissue response are listed in Table 2.2.
Table 2.2: Consequences of implant-tissue interactions (Hench & Wilson, 1993) Implant-tissue
Reaction
Consequence Example Bioinert Tissue forms a non-adherent
fibrous capsule around the implant Alumina, Zirconia and Carbon Bioactive Tissue forms an interfacial bond
with the implant
Hydroxyapatite (HA), Bioactive glass
Bioactive glass-ceramics
Bioresorbable Tissue replace implant β-tricalcium phosphate (β-TCP), carbonated hydroxyapatite, calcium carbonate
2.2.2.1 Bioinert Ceramics
Bioinert ceramics are biocompatible materials that maintain their mechanical and physical properties after implantation. This bioinert material undergoes little or no chemical reactivity, even after long term of exposure to the physiological condition and therefore, shows minimal interfacial bonds with the living tissues (Bhat, 2005).
Examples of this type of materials include alumina (Al2O3), zirconia (ZrO2), pyrolitic carbon, and silicon nitrides. Bioinert ceramics are very popular in orthopaedics and commonly used for structural support applications. They are also known to have
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excellent wear properties and are therefore useful for gliding functions (Binyamin et al., 2006; Li & Hastings, 1998).
2.2.2.2 Bioresorbable Ceramics
Bioresorbable ceramics refer to materials that, upon placement within the human body, would start to dissolve and slowly be replaced by advancing tissues. In other words, resorbable implants are designed to degrade gradually with time and be replaced with natural tissues (Bilotte, 2003). It leads to tissue regeneration instead of replacement.
The rate of degradation varies from one material to another. The advantage of this type of implant is that it will be replaced by normal functional bone, thus eliminating any long term biocompatibility problems. However, during the remodeling process, the load bearing capacity of the implant could possibly be weakened and resulted in mechanical failure. Therefore, the resorption rates of the material should be matched with the repair rates of body tissues (Hench & Wilson, 1993).
2.2.2.3 Bioactive Ceramics
Hench and Anderson (1993) define bioactive materials as a material that elicits a specific biological response at the interface of the material which results in the formation of a bond between the tissues and the material. When a bioactive material is implanted into the human body, it will interact to some extent with the surrounding bone or other tissue. An ion-exchange reaction between the bioactive implant and surrounding body fluids results in the formation of a bone-like apatite layer on the implant that is chemically and crystallographically equivalent to the mineral phase in the bone, which promotes the bonding between the natural tissues and the material (Liu et al., 2008).
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Typical examples of conventional bioactive ceramics used in orthopaedic surgery are synthetic HA, Bioglass®, Ceravital®, and A-W Glass-ceramic (Hench, 1998; Ratner et al., 2007).
The ability for the formation of this apatite layer on the implanted substrate in the body environment is essential for the direct bonding to living bone. An estimate of the potential for apatite layer formation on a ceramic material is carried out by in vitro testing. Kokubo and his colleagues developed a simulated body fluid (SBF) similar with regard to inorganic ions to the human body plasma (Kokubo et al., 1990; Kokubo &
Takadama, 2006). Materials that form apatite in SBF are expected to form apatite in the body and bond to living bone; therefore, SBF has been widely used to estimate the in vivo bone bioactivity of various types of bioactive materials (Kamitakahara et al., 2009).
2.2.3 Applications of Bioceramics
Bioceramics are produced in a variety of forms and phases, and serve many different functions in the repair of the human body, which are summarized in Table 2.2.
Most applications of bioceramics relate to the repair of the skeletal system, composed of bones, joints, and teeth, and to augment both hard and soft tissues. These repairs become necessary when the existing part becomes diseased, damaged, or just simply worn out.
There are many other applications of bioceramics including pyrolotic carbon coatings for heart valves and special radioactive glass formulations for the treatment of certain tumors (Carter & Norton, 2008). In other situations, bioceramics are used as reinforcing components in a composite, combining the characteristics of both components into a
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new material with enhanced mechanical and biochemical properties. Figure 2.3 shows a number of clinical uses of bioceramics (Hench & Wilson, 1993; Ishikawa et al., 2003).
Ceramics are also widely used in denstistry as restorative materials, gold porcelain crowns, glass-filled ionomer cements, endodontic treatments, dentures, and so forth and the materials used in these applications are called dental ceramics. Ceramics and glasses have been used for a long time outside the body for a variety of applications in the health care industry. Eye glasses, diagnostic instruments, chemical ware, thermometers, tissue culture flasks, chromatography columns, lasers and fibre optics for endoscopy are commonplace products in the industry (Hench & Wilson, 1993).
Table 2.3: Form, phase and function of bioceramics (Hench & Wilson, 1993)
Form Phase Function
Powder Polycrystalline Glass
Space filling, therapeutic treatment, regeneration of tissues
Coating Polycrystalline Glass
Glass-ceramic
Tissue bonding, thromboresistance, corrosion protection
Bulk Single crystal Polycrystalline Glass
Glass-ceramic
Composite (multi-phase)
Replacement and augmentation of tissue, replace functioning parts
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Figure 2.1: Clinical uses of bioceramics (Hench & Wilson, 1993)
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2.3 Bioactive Glasses and Glass-ceramics 2.3.1 Bioactive Glasses
The first and most thoroughly studied bioactive glass is known as Bioglass® 455S (Hench, 1991). Bioglass® 45S5 is a multicomponent oxide glass where the main composition SiO2, Na2O, CaO and P2O5. The majority of bioactive glasses and glass- ceramics are based on these four components and all current bioactive glasses are silicates. There are three key compositional features to these bioactive glasses that distinguished them from traditional soda-lime-silica glasses: a) less than 60 wt% SiO2, b) high Na2O and CaO contents, and c) high CaO/P2O5 ratio. These compositional features make their surface highly reactive when exposed to an aqueous medium such as the body fluids (Davis, 2003). The 45S5 composition and several typical bioactive glasses are given in Table 2.4.
Table 2.4: Composition (wt%) and mechanical properties of bioactive glasses (Cao &
Hench, 1996)
Component 45S5 Bioglass®
45S5.4F Bioglass®
45B15S Bioglass®
52S4.6 Bioglass®
55S4.3 Bioglass®
SiO2 45 45 30 52 55
P2O5 6 6 6 6 6
CaO 24.5 14.7 24.5 21 19.5
Na2O 24.5 24.5 24.5 21 19.5
CaF2 9.8
B2O3 15
Structure Glass and Glass- ceramic
Glass Glass Glass Glass
This work is studied by Hench and co-workers and summarized in the ternary SiO2-Na2O-CaO diagram as shown in Figure 2.2. It illustrates the compositional dependence of bone bonding and soft tissue bonding for the SiO2-Na2O-CaO glasses.
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Composition in the middle of the diagram (region A) forms a bond with bone and is defined as bioactive bone bonding boundary. When the concentration of SiO2 in the glass network exceeds 55% the rates of reaction decrease, and bonding to bone is very slow. Silicate glasses within region B behave as almost bioinert materials and elicit formation of a fibrous capsule at the implant-tissue interface. Glasses within region C are resorbable and disappear within 10-30 days of implantation. Compound of glasses within region D are not technically interesting and therefore, have not been tested as implants (Cao & Hench, 1996).
Figure 2.2: The SiO2-CaO-Na2O ternary phase diagram (Cao & Hench, 1996)
The main advantage of the bioactive glasses is the rapid surface reaction that brings about fast connections for tissue bonding and their primary disadvantages are mechanical weakness and low fracture toughness due to an amorphous two-dimensional glass network. The bending-tensile strength of most composition of bioactive glass vary between 40-60MPa, which make them unsuitable for load-bearing applications and find use as coatings on metals, in low-loaded or compressively loaded devices, in the form of powders or as the bioactive phase in composites (Hench & Wilson, 1993).
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Bioactive glasses may be produced in various forms depending on the repair function they will serve. One of the most successful uses of bioactive glass is as replacement for the ossicles (tiny bones) in the middle ear and to repair the bone that supports the eye. Cone-shaped plugs of bioactive glasses also have been used in oral surgery to fill the defect in the jaw created when a tooth is removed. In powder form, bioactive glasses are used in the treatment of periodontal disease and for the treatment of patients with paralysis of one of the vocal cords (Carter & Norton, 2007).
2.3.2 Bioactive Glass-ceramics 2.3.2.1 Glass-ceramic Processing
James, (1995) defined that glass-ceramics are materials obtained by controlled crystallization of certain glasses. Bioactive glass-ceramics have been developed to improve the mechanical performance of bioactive materials, or to introduce other interesting properties such as the machinable glass-ceramic Bioverit®. The formation of glass-ceramics is influenced by two important factors which are nucleation and growth of small crystal (< 1µm in diameter) and uniform size distribution. It is estimated that about 1012 to 1015 nuclei per cubic centimeter are required to achieve such small crystals.
In addition to the metallic agents already mentioned, Pt groups, TiO2, ZrO2 and P2O5 are widely used as nucleating agents. The nucleation of glass is carried out at temperatures much lower than the melting or glass transition temperature, at which the melt viscosity is in the range of 1011 to 1012 Poise for at least 1 to 2 h. To obtain a more microcrystalline phase, the glass is further heated to an appropriate temperature for maximum crystal growth. In this process, deformation of the products, phase transformation within the crystalline phases, or re-dissolution of some of the phases
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should be avoided. The crystallization is usually more than 90% complete when grain sizes are 0.1 to 1 µm, which are much smaller than in conventional ceramics. Figure 2.3 is a schematic representation of the temperature –time cycle for a glass-ceramic.
Melt & forming
Growth
Nucleation
Room temperature
Temperature
Time
Figure 2.3: Temperature-time cycle for a glass-ceramic
2.3.2.2 Properties of Glass-ceramics
Glass-ceramics have several desirable properties compared with glasses and ceramics. The thermal coefficient of expansion is very low. Due to the controlled grain size and improved resistance to surface damage, glass-ceramics can have at least double the tensile strength (from 100 to 200 MPa). The resistance to scratching and abrasion of glass-ceramics is similar to that of sapphire. The modulus of elasticity is of the order of 100 GPa, and the compressive strength is about five times the tensile strength, as given in Table 2.5.
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Table 2.5: Mechanical properties of glass-ceramics (Park, 2008)
Properties Bioglass® Ceravital® A-W Glass- ceramic® Young’s modulus (GPa)
Tensile strength (MPa) Compressive strength (MPa) Bending strength (MPa) Hardness (Vickers)
Fracture toughness (MPa.m1/2)
35 200
42 160-190
458 2.0
100-159 400 500 130 294 4.6
118 - 1080
215 680 3.34
A negative characteristic of the glass-ceramic is its brittleness. In addition, limitations on the compositions used for producing a biocompatible glass-ceramics hinder the production of glass-ceramic which has substantially higher mechanical strength. Thus, glass-ceramics cannot be used for making major load-bearing implants such as joint implants. However, they can be used as fillers for bone cement, dental restorative composites, and coating material (Billotte, 2003).
2.3.2.3 Commercial Bioactive Glass-ceramics
Several kinds of glass-ceramics compositions are bioactive and their behaviour in the body is very similar to that of bioactive glass which has an ability to form a strong interfacial bond with hard and soft tissues. There are three examples of well-known bioactive glass-ceramics that have been developed for implantation: machinable glass- ceramic (Bioverit® I), Ceravital® and A-W Glass-ceramic® (Carter & Norton, 2007).
Table 2.6 shows compositions of some bioactive glass-ceramics.
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Table 2.6: Compositions of some bioactive glass-ceramics (Cao & Hench, 1996; Park, 2008)
Type SiO2 CaO Na2O P2O5 MgO K2O
A-W Glass-Ceramic® 34.2 44.9 - 16.3 4.6
Ceravital® 40-50 30-35 5-10 10-15 2.5-5 0.5-3
Bioverit® I 29.5-50 13-28 - 8-18 6-28 -
All type of bioactive glass-ceramic composition in weight percent (wt%). In addition, Al2O3 (0-19.5), Na2O/K2O (5.5-9.5), F (2.5-7), Cl (0.01-0.6) and TiO2 (additions) are present in Bioverit® I. A-W Glass-ceramic® has CaF2 (0.5%).
A-W Glass-ceramic® is produced by crystallization of a glass of composition as can be seen in Table 2.6. The crystalline phases are oxyfluroapatite [Ca10(PO4)6(OH1F2)] and β-wollastonite (CaO-SiO2) and also content a residual glassy matrix. A-W Glass-ceramic® has excellent mechanical properties and forms a bond with bone that has very high interfacial bond strength. This type of glass has been used successfully in hundreds of patients for replacing part of the pelvic bone and in vertebral surgery (Hench & Kokubo, 1998). Ceravital® has been successfully used clinically in middle ear surgery to replace damaged bone. In this application the mechanical properties of the material are sufficient to support the minimal applied loads. To control the dissolution rate, Al2O3, F, and Cl are added in Ceravital® glass-ceramic. Bioverit® I is a mica-apatite glass-ceramic and known as machinable bioactive glass-ceramic. The key to the development of Bioverit® I was to form a phase separated base glass consisting of three glassy phases and to control the nucleation and crystallization by heat treating the glass.
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2.3.2.4 Mechanism of Bioactive Bonding
Bonding of bone to bioactive glasses and glass-ceramics involves 11 reaction stages summarized in Figure 2.4. The first five reaction stages that occur on the surface of bioactive glass and glass-ceramic do not depend on the presence of tissues. They occur in distilled water, tris-buffer solutions or SBF, and have been well studied using Fourier transform infrared (FTIR) spectroscopy, Auger electron spectroscopy, and electron microprobe analysis. These reactions result in a hydroxycarbonate apatite (HCA) crystal layer forming on the implant surface. Stages 6-11 are necessary for the implant to bond to tissues.
Increasing Time
11- Crystallization of matrix 10-Cellular attachment
9- Differentiation of stem cells
Surface reaction stages
8- Attachment of stem cells 7- Action of macrophages Log t
6- Adsorption of biologic moieties in HCA layer 5- Nucleation and crystallization of hydroxyl carbonate apatite (HCA)
4- Precipitation of amorphous calcium phosphate 2, 3- Dissolution and repolymerization of surface silica 1- Sodium hydrogen ion exchange
0- Initial glass surface
Figure 2.4: Sequence of interfacial reactions involved in forming a bond between bone and bioactive glasses (Cao & Hench, 1996)
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