PMMA-co-EHA cements for osteoprotheses

2011

Bárbara Joana Martins

Leite Ferreira

PMMA-co-EHA CEMENTS FOR

OSTEOPROSTHESES

Aveiro 2011

Bárbara Joana Martins

Leite Ferreira

PMMA-co-EHA cements for osteoprostheses

Dissertação apresentada à Universidade de Aveiro para cumprimento dos requisitos necessários à obtenção do grau de Doutor em Engenharia Biomédica, realizada sob a orientação científica do Doutor Rui Nunes Correia, Professor Associado com Agregação Aposentado do Departamento de Engenharia Cerâmica e do Vidro da Universidade de Aveiro.

o júri

presidente Prof. Dr. Fernando Manuel dos Santos Ramos

Professor Catedrático da Universidade de Aveiro

Prof. Dra. Maria Helena Mendes Gil

Professora Catedrática Aposentada da Faculdade de Ciências e Tecnologia da Universidade de Coimbra

Prof. Dra. Maria Helena Raposo Fernandes

Professora Catedrática da Faculdade de Medicina Dentária da Universidade do Porto

Prof. Dr. Nelson Fernando Pacheco da Rocha

Professor Catedrático da Universidade de Aveiro

Doutor Pedro Lopes Granja

Investigador Principal da Faculdade de Engenharia da Universidade do Porto

Prof. Dr. Rui Nunes Correia

agradecimentos Em primeiro lugar gostava de agradecer ao meu marido, Adão, pelo apoio, compreensão, carinho e incentivo dados ao longo da elaboração desta dissertação.

Aos meus pais pelo incentivo e apoio que sempre me deram, desde a minha entrada no ensino primário.

Ao Prof. Doutor Rui Nunes Correia pela sua orientação e amizade. Obrigada por me ter iniciado no estudo dos materiais para aplicações na área Biomédica.

A todas(os) as(os) colegas que durante este tempo passaram pelo Laboratório de Biomateriais (BioLab) do Departamento de Cerâmica e do Vidro e que de uma forma ou outra contribuiram para a realização deste trabalho. Um agradecimento muito especial às colegas (e amigas do peito) Ana Luísa Daniel, Nathalie Barroca e Poliana Lopes pela partilha de ideias, sugestões, pela ajuda na elaboração de alguns ensaios e sobretudo pela paciência e incentivo que sempre me deram.

À Professora Maria Helena Fernandes, da Faculdade de Medicina Dentária da Universidade do Porto, pela sua colaboração nos estudos de avaliação da biocompatibilidade in vitro dos materais.

À Engenheira Conçeição Costa pela aquisição dos espectros de difracção de raios-X.

À Ana Margarida Silva pelas sessões de microscopia electrónica de varrimento.

À Fundação para a Ciência e Tecnologia (FCT) pela bolsa concedida (Ref.ª SFRH/BD/17389/2004), fundamental para a realização desta dissertação. Por fim, à Universidade de Aveiro pelas óptimas condições de acolhimento que me proporcionou.

palavras-chave Cimento ósseo, PMMA, PMMA-co-EHA, parâmetros de cura, propriedades mecânicas, retenção de água, perda de peso, biocompatibilidade in vitro, HA, compósitos, enchimento cerâmico, bioactividade.

resumo O principal objectivo desta tese foi produzir novas formulações de PMMA-co-EHA e estudar a sua viabilidade como alternativa aos tradicionais cimentos ósseos de PMMA.

Assim sendo, foram inicialmente produzidos vários co-polímeros de PMMA-co-EHA e as suas propriedades mecânicas e comportamento in vitro foram avaliados. Os co-polímeros foram obtidos por polimerização radicalar e diversas formulações foram produzidas por substituição parcial do MMA (até cerca de 50 %) por EHA. Globalmente, os resultados sugerem que a substituição parcial do MMA por EHA diminuiu o módulo de elasticidade dos materiais e, consequentemente, aumentou a sua flexibilidade. Posteriormente, foram adicionados grânulos de PMMA pré-polimerizado (para se obter um cimento ósseo) às várias formulações de PMMA-co-EHA e as propriedades gerais dos cimentos resultante foram avaliadas. De um modo geral, os resultados obtidos revelaram que a substituição parcial do MMA pelo EHA levou a alterações benéficas dos parâmetros de cura (houve uma redução da temperatura máxima de polimerização e um aumento do tempo de cura), do comportamento in vitro (verificou-se um aumento da capaciadade de retenção de água) e das propriedades mecânicas (aumento da capacidade de flexão) dos novos cimentos. A resposta celular in vitro das novas formulações de PMMA-co-EHA foi comparada com a dos cimentos tradicionais. Para o efeito, foram avaliadas a adesão e proliferação celular de células tipo-osteoblastos MG63 e de células humanas provenientes de medula óssea. Os resultados revelaram que os dois tipos de células foram capazes de aderir e proliferar em ambas as formulações. A única excepção foi observada para a formulação preparada com maior percentagem de EHA, onde as poucas células que aderiram não conseguiram proliferar. Para além deste facto, verificou-se que o aumento da quantidade de EHA nos cimentos conduziu a uma crescente inibição do crescimento celular, sobretudo durante a primeira semana de cultura. Este facto foi relacionado com a crescente capacidade de retenção de água por parte das novas formulações e consequente libertação de alguns dos seus componentes tóxicos. Por último, os grânulos de PMMA foram parcialmente substituídos por partículas de HA e a influência desta substituição nos parâmetros de cura, nas propriedades mecânicas e no comportamento in

vitro dos compósitos resultantes foi também avaliada. A incorporação de HA

nos cimentos induziu uma série de alterações importantes nas suas propriedades finais: 1) diminuição da temperatura máxima de polimerização; 2) aumento significativo do tempo de cura; 3) aumento do valor do módulo elástico acompanhado de uma dimuição da sua força/tensão. Este último resultado foi relacionado com a baixa adesão interfacial entre os vários componentes e com uma distribuição heterogénea (possível aglomeração) das partículas de HA.

keywords Bone cement, PMMA, PMMA-co-EHA curing parameters, mechanical properties, water uptake capacity, weight loss, in vitro biocompatibility, HA, composites, ceramic filler, bioactivity.

abstract The main purpose of this thesis was to produce new formulations of PMMA-co-EHA and study its feasibility as being an alternative to traditional PMMA bone cements.

Thus, were originally produced several co-polymers of PMMA-co-EHA and its mechanical properties and in vitro behaviour were evaluated. The copolymers were obtained by radical polymerization and several formulations were produced by partial replacement of MMA (up to about 50%) for EHA. Overall, the results suggest that the partial replacement of MMA by EHA decreased the modulus of the materials and, consequently, increased its flexibility.

Then, PMMA commercial beads were added to PMMA-co-EHA formulations (to get bone cement) and the general properties of the resulting bone cements were evaluated. In general, the results revealed that the partial replacement of MMA by EHA led to beneficial changes in curing parameters (there was a reduction of the peak temperature and an increase of curing/setting time), in the in vitro behaviour (the water capacity increased) and in the mechanical properties (the bending strength increased) of new cements. The in vitro cellular response of new formulations of PMMA-co-EHA was compared with that of traditional PMMA bone cement. To this end, we tested the cell adhesion and proliferation of osteoblast-like MG63 cells and human cells from bone marrow. The results revealed that both types of cells were able to attach and proliferate in both formulations. The only exception was observed for the formulation prepared with the highest percentage of EHA, where a few cells that adhere failed to proliferate. Moreover, it was found that increasing the amount of EHA in cement led to an increasing inhibition of cell growth, especially during the first week of culture. This was related to increased water uptake capacity by the new formulations and consequent release of some of its toxic components. Finally, PMMA commercial beads were partially replaced by HA particles and the influence of this substitution on the curing parameters, the mechanical properties and in vitro behaviour of the resulting composites was also evaluated. Incorporation of HA into the bone cements induced a number of significant changes in its final properties: 1) decrease the peak temperature; 2) increase of curing time, 3) increasing the value of elastic modulus accompanied by decrease of the strength/tension. This last finding was related to poor interfacial adhesion between the various components of the bone cements and a heterogeneous distribution (possible agglomeration) of HA particles.

LIST OF TABLES ... IX LIST OF FIGURES ... X

CHAPTER 1 ... 1

INTRODUCTION ... 1

1.1 PMMA BONE CEMENT IN ORTHOPAEDICS ... 1

1.2 CHEMICAL COMPOSITION AND POLYMERIZATION OF PMMA BONE CEMENT ... 4

1.2.1 CHEMICAL COMPONENTS ... 4

1.2.2 POLYMERIZATION ... 4

1.2.3 PHYSICAL ASPECTS OF MIXING ... 8

1.3 COMMERCIALLY AVAILABLE BONE CEMENTS ... 10

1.4 THERMAL EFFECTS ON CURING ... 11

1.5 FACTORS THAT CAN AFFECT THE PROPERTIES OF PMMA BONE CEMENTS ... 13

1.5.1 POWDER/LIQUID RATIO ... 13

1.5.2 INITIATOR/ACTIVATOR RATIO ... 14

1.5.3 PMMA BEAD SIZE ... 15

1.5.4 MIXING METHODS ... 16

1.5.5 CROSSLINKING AGENTS ... 18

1.6 MAIN SIDE EFFECTS ... 19

1.6.1 ASEPTIC LOOSENING ... 19

1.7 ADDITIVES ... 23

1.7.1 RADIOPACIFIERS ... 23

1.7.2 ANTIBIOTICS ... 25

1.7.2.1 THE CHOICE OF THE ANTIBIOTIC ... 25

1.7.2.1 MECHANISMS OF ANTIBIOTIC RELEASE FROM PMMA BONE CEMENTS ... 28

1.7.2.2 INFLUENCE OF THE ANTIBIOTIC ON THE PROPERTIES OF THE CEMENT ... 29

1.8 IMPROVEMENT OF PMMA BONE CEMENTS ... 30

1.8.1 FIBER REINFORCEMENT ... 30

1.8.2 REINFORCEMENT WITH BIOACTIVE CERAMIC FILLERS ... 32

1.9. CONCLUSIONS ... 36

REFERENCES ... 37

CHAPTER 2 ... 51

PREPARATION AND CHARACTERIZATION OF NEW PMMA-EHA CO-POLYMERS: MECHANICAL PROPERTIES AND IN VITRO BEHAVIOUR ... 51

2.1 ABSTRACT ... 51

2.2. INTRODUCTION ... 52

2.3. MATERIALS AND METHODS ... 53

2.3.1 MATERIALS ... 53

2.3.2 METHODS ... 54

2.3.2.1 PREPARATION OF THE CO-POLYMERS ... 54

2.3.2.2.2 MECHANICAL PROPERTIES (THREE-POINT BENDING AND UNIAXIAL

COMPRESSION) ... 56

2.4. RESULTS AND DISCUSSION ... 57

2.4.1 WATER UPTAKE (WU), WEIGHT LOSS (WL) AND SURFACE EVALUATION ... 57 2.4.2 MECHANICAL PROPERTIES ... 61 2.4.2.1 BENDING (THREE-POINT) ... 61 2.4.2.2 UNIAXIAL COMPRESSION ... 64 2.5. MAIN CONCLUSIONS ... 66 REFERENCES ... 67 CHAPTER 3 ... 69

NEW PMMA-CO-EHA BONE CEMENTS (I): PREPARATION AND STUDY OF THE CURING PARAMETERS, MECHANICAL PROPERTIES AND IN VITRO BEHAVIOUR ... 69

3.1 ABSTRACT ... 69

3.2 INTRODUCTION ... 70

3.3. MATERIALS AND METHODS ... 73

3.3.1 MATERIALS ... 73

3.3.2 METHODS ... 73

3.3.2.1 FORMULATION OF THE CEMENTS ... 73

3.3.2.2 MILLING OF THE COMMERCIAL PMMA BEADS ... 75

3.3.2.3 CHARACTERIZATION OF THE MATERIALS ... 76

3.3.2.3.3 SURFACE EVALUATION ... 76

3.3.2.3.4 MECHANICAL BEHAVIOUR ... 77

3.4. RESULTS AND DISCUSSION ... 77

3.4.1 CURING PARAMETERS (PEAK TEMPERATURE AND SETTING TIME) ... 77

3.4.2 IN VITRO BEHAVIOUR (WATER UPTAKE AND LOSS OF WEIGHT) AND SURFACE EVALUATION ... 81

3.4.3 MECHANICAL PROPERTIES ... 86

3.5. CONCLUSIONS ... 89

REFERENCES ... 91

CHAPTER 4 ... 97

NEW PMMA-CO-EHA BONE CEMENTS (II): IN VITRO BIOCOMPATIBILITY ASSESSMENT BY MG63 AND HUMAN BONE MARROW CELL CULTURES ... 97

4.1 ABSTRACT ... 97

4.2 INTRODUCTION ... 98

4.3 MATERIALS AND METHODS ... 99

4.3.1 MATERIALS ... 99

4.3.2 PREPARATION OF THE PMMA-CO-EHA BONE CEMENTS ... 99

4.3.3 IN VITRO CELL CULTURES ... 100

4.3.3.1 MG63 OSTEOBLAST-LIKE CELLS ... 100

4.3.3.2 HUMAN BONE MARROW CELLS ... 100

4.3.4 BIOCHEMICAL AND MICROSCOPICAL ASSAYS ... 101

4.3.4.3 CONFOCAL MICROSCOPY AND SEM ... 102

4.3.4.4 STATISTICAL ANALYSIS ... 103

4.4 RESULTS ... 103

4.4.1 BEHAVIOUR OF MG63 OSTEOBLAST-LIKE CELLS ... 103

4.4.1.1 LIGHT MICROSCOPY ... 103

4.4.1.2 CELL VIABILITY/PROLIFERATION BY MTT REDUCTION ... 105

4.4.1.3 CONFOCAL LASER SCANNING MICROSCOPY (CLSM) ... 107

4.4.2 BEHAVIOUR OF HUMAN BONE MARROW CELLS ... 109

4.4.2.1 CELL VIABILITY/PROLIFERATION AND ALP ACTIVITY ... 109

4.4.2.2 SEM MICROGRAPHS ... 112

4.5. DISCUSSION ... 114

4.6. CONCLUSIONS ... 117

REFERENCES ... 118

CHAPTER 5 ... 121

NEW PMMA-CO-EHA BONE CEMENTS FILLED WITH HA: CURING PARAMETERS, IN VITRO BEHAVIOUR AND MECHANICAL PROPERTIES ... 121

5.1 ABSTRACT ... 121

5.2. INTRODUCTION ... 122

5.3 MATERIALS AND METHODS ... 123

5.3.1 MATERIALS ... 123

5.3.2 METHODS ... 124

5.3.3.1 DETERMINATION OF THE CURING PARAMETERS ... 125

5.3.3.2 MECHANICAL PROPERTIES ... 125

5.3.3.3 IN VITRO BEHAVIOUR (WATER UPTAKE AND LOSS OF WEIGHT) ... 126

5.3.3.4 SURFACE EVALUATION (BY SEM AND XRD) ... 126

5.4. RESULTS AND DISCUSSION ... 126

5.4.1 CURING PARAMETERS (PEAK TEMPERATURE AND SETTING TIME) ... 126

5.4.2 MECHANICAL PROPERTIES ... 128

5.4.3 WATER UPTAKE AND LOSS OF WEIGHT ... 132

5.4.4 SURFACE EVALUATION ... 134 5.4.4.1 SEM MICROGRAPHS ... 134 5.4.4.2 XRD PATTERNS ... 139 5.5 CONCLUSIONS ... 142 REFERENCES ... 143 CHAPTER 6 ... 145

MAIN CONCLUSIONS AND SUGGESTIONS FOR FUTURE STUDIES ... 145

6.1 MAIN CONCLUSIONS ... 145

PMMA Poly (methylmethacrylate)

THA Total hip arthroplasty

BPO Benzoyl peroxide

BaSO4 Barium sulphate

ZrO2 Zirconium dioxide

MMA Methyl methacrylate

HQ Hydroquinone

DMPT N,N dimethyl-p-toluidine

P/L ratio Powder/liquid ratio

BPO/DMPT ratio Initiator/activator ratio

IHQM 2, 5-diiodo-8-quinolyl methacrylate

TPB Triphenyl bismuth

BIEM 2-(2-bromoisobutyryloxy) ethyl methacrylate BPEM 2-(2-bromopropionyloxy) ethyl methacrylate SRC-PMMA Self-reinforced composite of PMMA

HA Hydroxyapatite

TCP Tricalcium phosphate

Bis-GMA Bisphenol-a-glycidyl methacrylate

AW-GC Wollastonite glass-ceramic

β-TCP β-tricalcium phosphate

SBF Simulated body fluid

α-TCP α- tricalcium phosphate

EHA 2-ethyl hexylacrylate

PBS Phosphate-buffered saline

ALP Alkaline phosphatase activity

SEM Scanning electron microscopy

BSA Bovine serum albumin

Table 1.1: Commercially available and most commonly used PMMA bone cementsUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU...10 Table 1.2: Commercially available antibiotic-loaded cementsUUUUUUU...27 Table 2.1: Proportions of MMA and EHA in the materials producedUUUUU54 Table 3.1: Composition of the liquid and solid phases of all the formulations studiedUUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU.74

Table 3.2: Setting times (tset) and maximum temperatures (Tmax) for all the

formulations studiedUUUUUUUUUUUUUUUUUUUUUUUUU..79 Table 5.1: Chemical composition of all the PMMA-co-EHA/HA composites preparedUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU124 Table 5.2: Setting times (tset) and maximum temperatures (Tmax) for all the

Figure 1.1: (a), (b) Components of a total hip arthroplasty; (c) Position of the cementU..UUUUUUUUUUUUUUUUUUUUUUUUUUUUUU3 Figure 1.2: Poly Polymerization of PMMA: (a) Reduction-oxidation process of benzoyl peroxide (BPO) caused by N,N dimethyl-p-toluidine (DMPT); (b) Initiation; (c) Chain propagation; (d) Termination by recombinationUUUUUUUUUUUUUUUUUUUUUUUUUUU...7 Figure 1.3: Polymerization temperature versus time (Tmax = maximum curing

temperature; Tamb = ambient temperature; DT = doughing time; WT = working

time; ST = setting time (when temperature reaches half of the difference Tmax −

Tamb)UUUUUUUUUUUUUUUUUUU... 9

Figure 2.1: (a) Evolution of the water uptake and (b) loss of weight during immersion in PBS (37 ºC)UUUUUUUUUUUUUUUUUUUUU...58 Figure 2.2: SEM micrographs of polymer surfaces before and after 14 days in PBS (37 ºC)UUUUUUUUUUUUUUUUUUUUUUUUUUUU...60 Figure 2.3: Typical force-displacement curves obtained for the polymers preparedUUUUUUUUUUUUUUUUUUUUUUUUUUUUU... .62 Figure 2.4: Elastic modulus obtained from three-point bending testsUUUU.63 Figure 2.5: (a) Force-displacement curves of the polymers prepared; (b) Elastic modulus obtained by compression testsUUUUUUUUUUUUUUUU..65

Figure 3.1: Particle size distribution of the PMMA commercial beads after millingUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU...75

Figure 3.2: Time-temperature diagram obtained for all the formulations studiedUUUUUUUUUUUUUUUUUUUUUUUUUUU... 78

Figure 3.3: (a) Evolution of the water uptake (WU) and (b) loss of weight (LW) of all the formulations during the immersion time in PBS (37 ºC)... 83

Figure 3.5: Typical stress-strain curves obtained for all the formulationsUU...87

Figure 3.6: Elastic modulus of all the formulations prepared, obtained by three-point bending testsUUUUUUUUUUUUUUUUUUUUUUUUUU89 Figure 4.1: Light microscopy images (x 20) from the viability/proliferation MTT assay of MG63 osteoblast-like cells for (a) 2 days and (b) 12 daysUUUU ..105

Figure 4.2: Cell viability/proliferation of MG63 osteoblast-like cells cultured on the surface of the materials for 12 days, estimated by the MTT assayUUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU106 Figure 4.3: CLSM images (200x) of the MG63 osteoblast-like cells cultured for 2, 7 and 12 daysUUUUUUUUUUUUUUUUUUUUUUUUUUU..109

Figure 4.4: (a) Cell viability/proliferation, estimated by the MTT assay and (b) ALP activity of human bone marrow cells cultured for 21 daysUUUUUU..111 Figure 4.5: SEM micrographs of human bone marrow cells cultured for 21 days on PMMA-co-EHA cementsUUUUUUUUUUUUUUUUUUUUU..114

Figure 5.1: Time-temperature diagram obtained for all the composites studiedUUUUUUUUUUUUUUUUUUUUUUUUUUUUUU...127

Figure 5.2: Typical stress-strain curves obtained for the composites (selected specimens)UUUUUUUUUUUUUUUUUUUUUUUUUUUUU130

Figure 5.3: Elastic modulus, obtained by three-point bending, of all the PMMA-co-EHA/HA compositesUUUUUUUUUUUUUUUUUUUUUUU..132

Figure 5.4: (a) Evolution of the water uptake and (b) loss of weight of the PMMA-co-EHA/HA composites during the immersion time in PBS (37 ºC)UUUUU134

Figure 5.6: XRD patterns of the composites before immersion, after 1 day and after 14 days in PBSUUUUUUUUUUUUUUUUUUUUUUUU...142

Chapter 1

Introduction

1.1 PMMA bone cement in orthopaedics

Bone cements are materials used to repair damaged or diseased areas of bones or to fix osteoprostheses [1]. Most of the commercial bone cements used in dentistry and orthopaedic surgery are based on poly (methylmethacrylate) (PMMA) and as a group they are called acrylic- based cements [1,2].

Since its introduction by Sir John Charnley [3] in 1958, acrylic bone cement has been used extensively for the fixation of joint replacements (including hip, knee and ankle). Today, approximately one-half million hip replacement surgeries are performed every year worldwide – and 70% of them using bone cements [4]. Orthopaedic surgery is responsible for giving long-term pain relief and restoring mobility, functionality, and high quality of life to millions of patients over the world [4,5].

In recent years, the application of bone cement in orthopaedic surgery has been expanded to vertebroplasty, a procedure in which cement is injected percutaneously into the vertebral body in order to stabilize fractures that occur primarily as a result of osteoporosis. Another variation of this procedure is khyphoplasty, during which a balloon is inserted percutaneously into the vertebral body, inflated to restore the height of the compressed vertebrae and subsequently filled with injected bone cement to stabilize the fracture. These newer applications of bone cement have been successful in relieving pain and restoring vertebral strength and function [2,6,7].

The focus of this thesis is the application of bone cement to implant fixation. Approximately 50% of total joint replacements utilize bone cement to achieve implant fixation and currently the survival probabilities of cemented total joint arthroplasties average at least 90% after 15 years, especially for hip and knee implants in patients over age 50 [2,5]. In this application, the primary functions of bone cement are to stabilize the prosthesis and transfer loads between the implant and the bone. PMMA does not chemically bond with either interface, but rather serves as a grouting or space-filling material [5,8]. Figure 1.1 shows the components of a total hip arthroplasty (THA) and the location of bone cement at the interface. For ultimate success of the total joint arthroplasty, the interfaces and the cement itself must withstand the loading of the hip [1,4].

Figure 1.1: (a), (b) Components of a total hip arthroplasty; (c) Position of the cement [9].

Commercial bone cements are supplied as two component systems, consisting of a powder and a liquid, which are mixed in the operating room, delivered to the implant site prior to introduction of the prosthesis and polymerized in situ. Rheological and setting characteristics of the mix as well as mechanical properties, porosity and residual monomer concentration are critical to the clinical success of the cement [1,2]. Some of these properties, along with the in

vitro and biological response will be discussed.

1.2 Chemical composition and polymerization of PMMA bone cement

1.2.1 Chemical components

Most commercially available cements consist of a two component system: a solid (roughly 70 % of the total) and a liquid. The solid (or powder) part consists of pre-polymerized beads, usually of poly methylmethacrylate (PMMA) (or sometimes a copolymer of PMMA/other acrylics or PMMA/styrene), benzoyl peroxide (BPO) (used as an initiator of the polymerization), barium sulphate (BaSO4)or zirconium dioxide (ZrO2) (opacifying agents to allow imaging under

X-rays) and, optionally, a colorant (such as clorophyllin) and antibiotics. The PMMA beads usually are between 10 and 150 µm in diameter and are added to mitigate the high curing temperature and the shrinkage of the monomer. The liquid component contains the methyl methacrylate (MMA) monomer (sometimes mixed with other acrylic monomers), a retarder (such as hydroquinone, HQ, to avoid polymerization of the monomer during storage) and N,N dimethyl-p-toluidine (DMPT), used as the activator of the initiator [1,2,5].

1.2.2 Polymerization

The reaction by which PMMA and other vinyl polymers are formed is known as free radical or addition polymerization, characterized by three major kinetic stages: initiation, propagation or growth, and termination. Initiation in free radical polymerization encompasses the set of reactions that generate the primary radicals which give rise to the active centers for polymer chain growth. Radicals can be formed by several different methods including the collision of two monomer molecules of sufficient energy or the decomposition of an initiator

molecule by means of heat, light or chemical reaction. In the case of cold-cured PMMA bone cements, primary radicals are generated by reaction between the initiator (BPO) and the activator (DMPT) [1]. DMPT causes the decomposition of BPO in a reduction-oxidation process, Figure 1.2 (a), and the produced radicals can induce polymerization by addition to the C=C double bond in the MMA – Figure 1.2 (b) [1,8]. It is believed that only the benzoyl free radicals initiate polymerization, the others, as well as part of the benzoyl, undergoing side reactions [10]. The new radical reacts with another MMA molecule in the same way as the initiator fragment did. Another radical is always formed when this reaction takes place over and over again, adding monomer molecules to the growing chains– Figure 1.2 (c). A significant consequence of this propagation phase is an increase in viscosity, due to the increasing concentration of polymer molecules and increasing molecular weight of the growing chains [1,2]. Termination is the final step in the polymerization process, whereby propagation ceases with the annihilation of radicals, either by disproportionation or by recombination of the radical-carrying chains. Recombination is the simplest way, wherein the two unpaired electrons join to form a bond – Figure 1.2 (d).

C O C O O

+

₊

-

•

+

•

+

Figure 1.2: Poly Polymerization of PMMA: (a) Reduction-oxidation process of benzoyl peroxide (BPO) caused by N,N dimethyl-p-toluidine (DMPT); (b) Initiation; (c) Chain propagation; (d) Termination by recombination. Adapted

from Ref. [1]. H2C C CH3 C O O CH3 • CH2 C CH3 C O O CH3 wwwwww • wwwwww C CH3 C O H2C wwwwww O CH3 CH2 C wwwwww CH3 C O O CH3 C H H C CH3 C O O CH3 •

Free radical vinyl polymerization CH2 C CH3 C O O CH3

[

]

1.2.3 Physical aspects of mixing

When the liquid and solid components of bone cements are mixed together, both physical and chemical phenomena take place. The physical part of setting includes: 1) solvation of the powder, as well as the BPO; 2) diffusion of liquid into the organic powder and swelling; 3) polymer-polymer diffusion from the liquid to the solid phase; 4) monomer evaporation. Simultaneously, the redox reaction between BPO and DMPT occurs and consequently the initiation step of polymerization begins [2,10].

This sequence of physico-chemical events is macroscopically identified by a succession of four stages that lead to complete hardening (curing). As shown by Kühn [8], they are:

I) mixing phase: the period during which an homogeneous dough is formed; the cement has low viscosity and the beads are being wetted by the monomer;

II) waiting phase: the dough is still sticky and adheres to the surgical glove; in this phase the polymerization initiates with a consequent increase in viscosity (decrease in polymer chains mobility);

III) working phase: the dough is no longer sticky and can be kneaded; there is propagation of the polymerization, heat generation and further increase in viscosity;

IV) setting phase: the dough becomes a hard cement; the peak temperature is attained, with termination of chain growth and mobility.

This entire process occurs over the course of six to twelve minutes, depending on the particular cement composition and the ambient temperature [2]. The phases the cement dough goes through are shown in Figure 1.2.

Figure 1.3: Polymerization temperature versus time (Tmax = maximum curing

temperature; Tamb = ambient temperature; DT = doughing time; WT = working

time; ST = setting time (when temperature reaches half of the difference Tmax −

1.3 Commercially available bone cements

Table 1.1: Commercially available and most commonly used PMMA bone cements. Adapted from ref. [1] and [8].

Bone cement Manufacturer Market

C-ment® 1 E.M.C.M.B.V. Central Europe, G

C-ment® 3 E.M.C.M.B.V. Central Europe, G

Cemex® Isoplastic Tecres South Europe; I

Cemex® RX (LV) Tecres South Europe; I

Cerafix® LV Ceraver Osteal South Europe; France CMW® 1 radiopaque De Puy, J & J Worldwide

CMW® 2 De Puy, J& J Worldwide

CMW® 3 De Puy, J & J Worldwide

DuracemTM 3 Sulzer Central Europe, CH

Durus® H Macmed Orthopedics South Africa

Endurance® De Puy, J & J USA

Osteobond® Zimmer USA

Osteopal® Merck Worldwide

Palacos® LV Schering Plough Worldwide

Palamed® Merck Worldwide

Subiton RO Prothoplast Argentina

Surgical Simplex® P Stryker Howmedica Worldwide, USA

Zimmer® Zimmer USA

CH: Switzerland; G: Germany; I: Italy; USA: United States of America; J & J: Johnson & Johnson.

1.4 Thermal effects on curing

During the polymerization of PMMA, the carbon double bond in the MMA monomer cleaves and is replaced by a single bond. This process releases 52 KJ mol-1MMA per bond. Therefore, the polymerization reaction leads to a large

rise in the temperature of the material during curing [2,10]. High temperatures cause quick setting and may cause damage to the surrounding tissue. The rate of polymerization, the temperature rise on setting and the setting time are sensitive to ambient temperature. Thus, when the temperature of the operating room increases, the polymerization rate also increases and the dough hardens quicker [1,5]. For example, the bone cement Surgical Simplex® P has a setting time of 9 minutes at room temperature of 24 ºC, 12 minutes at 21 ºC and 15 minutes at 18 ºC [11]. Also the temperature of the powder and liquid components, the temperature of the implant and of the mixing equipment can markedly affect the setting time and temperature. If the cement components are stored at temperatures lower or higher than that of the room, sufficient time (12 to 24 h) must be allowed for them to reach the appropriate ambient operating room temperature before they are mixed, otherwise setting time will be correspondingly lengthened or shortened [12]. If an implant is used while still

warm from the autoclave, setting time will be reduced. Mixing equipment still warm from storage or autoclaving will also induce a shorter setting time than expected [1].

One of the efforts to adjust the setting time and improve the handling properties of bone cements involves pre-chilling their liquid constituents, prior to their mixing with the powder. Lidgren et al. compared effects of ambient temperatures (4 ºC vs. 21 ºC) on high viscosity, high molecular weigh antibiotic-containing polymethacrylate bone cements mixed by hand or vacuum. If components were pre-chilled at 4 ºC, mixing became easier and handling characteristics were also improved. Pre-chilling did not cause any differences in thermal or mechanical parameters [13].

Hansen and Jensen [14] made comparative studies of nine commercial bone cements whose components were stored at room temperature or chilled to 5 ºC and mixed either manually or under vacuum. They found that pre-chilling and vacuum mixing prolonged the setting time and preserved a lower viscosity during the handling period.

Lewis [15] also studied the influence of the storage temperature of constituents prior to mixing (4 ºC vs. 21 ºC) and the mixing method (hand mixing vs. vacuum mixing) on the uniaxial tension-compression fatigue performance and porosity of Palacos® R acrylic bone cement. The results showed that although the mixing method exerted a significant influence on the fatigue performance and porosity, the effect of the storage temperature on either of these parameters was not significant.

More recently, Sullivan and Topolesky [16] studied the effects of the initial component temperature (-14 ºC, 6 ºC and 23 ºC) on the rheological (apparent viscosity) and handling characteristics (setting time, working time, and peak exotherm temperature) of an high viscosity cement, Palacos® R. By adjusting the initial component temperature, Palacos® R was able to mimic the flow characteristics of a medium viscosity bone cement (Simplex®, used as a benchmark) at room temperature. Lowering the initial component temperature

increased the working and setting times; however, a significant effect on the peak exotherm temperature was not observed.

1.5 Factors that can affect the properties of PMMA bone cements

To design bone cements with predictable intra-operative and post-operative behaviour, we must understand how cement formulations affect the polymerization and the properties of the end product. In this section, the effect of several factors on curing parameters (such as the setting time and maximum temperature of polymerization) and material properties are detailed.

1.5.1 Powder

/

A parameter that has a very strong effect on the working time of the cement is the powder/liquid (P/L) ratio, which is taken as the ratio between the weight of the powder (g) and the volume of the liquid (mL). Most of the commercial bone cement formulations have a P/L ratio of 2 [1,8].

Meyer et al. [17] reported the effect of changes in P/L (from 3.0 to 1.5) on the setting properties of acrylic bone cement. They showed that although the working time tended to remain unchanged, there was a definitive rise in maximum temperature when the P/L ratio decreased, due to the fact that it is the polymerization of the liquid that causes the heat release.

Pascual B. and co-workers [18] tested the influence of P/L, from 2.0 to 1.86, on a new formulation of acrylic bone cements. Reduction on P/L led to higher

temperature peaks and shorter setting times, whereas the residual monomer content increased slightly. Polymerization shrinkage was slightly greater (because of the introduction of higher proportions of monomer in the formulation) and the mechanical properties were similar to those obtained with conventional P/L ratios.

In general, it is known that the use of smaller P/L ratios favours the wetting and swelling of PMMA beads by the liquid component and subsequently, improves the handling characteristics of the mixture in the dough state. The main goal of this modification is improved handling and workability, with enhanced adhesion between phases. As the P/L ratio increases, the peak temperature decreases. These results can be understood in terms of the relative amounts of initiator and monomer present at various P/L ratios and the role played by unreactive particles in absorbing heat [10].

More recently, special attention has been given to the P/L ratio of PMMA bone cements commonly used in vertebroplasty [19,20]. This is because clinicians practicing vertebroplasty commonly alter the mixture of powder-to-monomer recommended by the manufacturer, in an effort to decrease viscosity and increase working time [19]. This procedure could explain one of the main problems in this type of surgery: outflow of bone cement into the spinal and venous system: bone cement that is too fluid or slow setting is more likely to escape its site of injection. Therefore, understanding the factors that can affect the properties of bone cements is extremely important.

1.5.2 Initiator

/

Altering the initiator/activator (BPO/DMPT) ratio has a significant effect on setting time, maximum temperature of polymerization and strength.

Unfortunately, adjusting BPO and DMPT concentrations to delay setting reduces the overall strength [1,20].

The rate of radical formation is dependent on the concentrations of activator and initiator. Faster radical formation activates more monomers that act as nucleation sites for polymer chain growth and has a number of downstream effects: 1) it will speed up the overall polymerization process, decreasing setting time; 2) more individual polymer chains will form simultaneously, decreasing the average molecular weight and affecting the strength and tensile properties of the cement [20].

To determine the proper BPO/DMPT ratio, a system that uses two liquid solutions system was developed [21]. The use of two-part liquid solutions rather than conventional liquid/powder cements allows for better mixing and consistency among samples. Also, it reduces porosity caused by trapping of air while mixing (air pockets act as fracture initiation sites). Both parts contained dissolved PMMA powder and MMA monomers in a 4:5 ratio. BPO was added to the first solution and DMPT to the second. Samples with several concentrations of BPO and DMPT were prepared for comparative testing. It was reported that this two-solution cement compositions were comparable to Simplex® P bone cement with regard to polymerization exotherm and setting time, but with higher flexural strengths and moduli. Residual monomer content was significantly affected by both the individual concentrations of BPO and DMPT (and their molar ratios). Although the new system presented significantly higher residual monomer content than Simplex® P, this can be attributed to their higher initial monomer concentration rather than a lower degree of conversion [22].

1.5.3 PMMA bead size

The average diameter and size distribution of PMMA beads play an important role in the curing properties. Aside from its structural role as a component of the

cement matrix, PMMA beads serve as a heat sink, dissipating energy released by the exothermic polymerization of MMA. Samples containing PMMA particles with larger mean diameters and widespread distributions of particle size had lower peak polymerization temperatures and longer setting times [20,23].

Pascual B. et al. [24] prepared formulations with PMMA particles in the range 10-60 µm average diameter and a relatively wide size distribution. Their results indicated that the use of PMMA particles of 50-60 µm average diameter and size distribution of 10-140 µm significantly changes the peak temperature and setting time, in comparison with commercial systems CMW® and Rostal, without any noticeable mechanical deterioration. Pascual and co-workers found that as the PMMA average bead size decreases, the peak curing temperature increases and the average curing time decreases.

1.5.4 Mixing methods

Mixing methods have a large impact on bone cement properties [20]. They are categorized as hand mixing, centrifugation, vacuum mixing and combined mechanical mixing [1,5]. Depending on the mixing technique, air may become trapped in the cement mixture, increasing porosity. As previously reported, this weakens the cement and provides interfaces for fatigue cracks to develop [20]. In hand mixing the powder is added to the liquid (which may or may not have been chilled to a temperature usually between -15 to 6 ºC) in a polymeric (usually polypropylene) bowl. These components are stirred with a polypropylene spatula for a period of time between 45 and 120 s [5]. This type of mixing method can introduce a significant amount of air into the mixture and a relatively high degree of trapped porosity in the set cement - as high as 10 % or greater [2,20].

In centrifugation mixing the hand-mixed dough is immediately poured into a syringe (from which the nozzle is detached) that is then promptly placed in a centrifuge and spun with a speed of 2300 to 4000 rpm for 30-180 s [1,5].

In 1983, Demarest et al. [25] first introduced the concept of mixing the powder and liquid components of bone cements under partial vacuum as a means of reducing this porosity [2]. Other studies, subsequently, demonstrated that an effective vacuum mixing can reduce the porosity in bone cements and improve their mechanical properties (in particular their flexural and compression strength as well the modulus of elasticity) [26-29].

Another study [13] showed that vacuum mixing improved fracture strength, maximal deflection, modulus of elasticity and hardening, when compared with hand mixing. The fatigue life was ten times longer, the setting time was delayed and the maximum peak temperature decreased. Radiographic analysis showed that not only the microporosity but also the macroporosity were reduced. After 2 months in Ringer’s solution the mechanical properties deteriorated slightly but the differences between the mixing procedures remained unchanged.

The reduction in porosity of acrylic bone cements and the associated improvement in mechanical properties has made vacuum mixing the leading method for clinical use [1,2]. There are currently a wide variety of commercially available vacuum mixing systems including the Advanced Cement Mixing System (Stryker Instruments, Kalamazoo, MI), the MixOR System and Vortex TM Vacuum Mixer (Smith & Nephew, Memphis, TN), the Optivac ® (Biomet Merck, Sjobo, Sweden) and the HiVac TM Bowl and Syringe Systems (Summit Medical, Gloucestershire, UK) [2]. There are no generic steps in vacuum mixing. Usually, 5 to 100 KPa of vacuum is applied for 15 to 150 s to the dough during the mixing stage. In some instances, air within the powder is evacuated before mixing. The mixing apparatus, the vacuum level, the vacuum application time and the mixing frequency differ among procedures [1].

1.5.5 Crosslinking agents

The extent to which monomers react is expected to have an important effect on the physico-mechanical properties and biocompatibility of bone cements. Incorporation of a crosslinking agent causes an insoluble network to form during the course of polymerization. It is known that crosslinking may improve physical properties, such as modulus of elasticity, hardness, heat distortion, solvent resistance, shrinkage and glass transition temperature. Different crosslinking agents may have different effects on the mechanical properties depending on the inherent structure of the crosslinking agent. Crosslinking agents containing longer and more flexible chains may aid in achieving a higher degree of crosslinking without impairing mechanical properties [30].

In 1997, Deb et al. [30] evaluated the effect of three crosslinking agents on the curing parameters and mechanical properties of acrylic bone cements. The crosslinking agents were dimethacrylates containing a range of chain lengths and degrees of flexibility. For the parent formulation of the bone cement, PMMA powder was used, and the properties and kinetics of curing were compared for the same system with and without crosslinking agents. It was observed that with the addition of crosslinking agents: 1) the curing parameters were not greatly affected but the dough time increased with increasing concentration of crosslinking agents; 2) the tensile strength of the cement with small amounts of crosslinking agents increased, in general, in comparison with the parent cement.

In 2001, the same author evaluated the effect of crosslinking agents on the mechanical properties of acrylic bone cements containing radiopacifiers [31]. The results indicated that the incorporation of dimethacrylate crosslinking agents, such as polyethylene glycol dimethacrylate, with radiopaque agents in commercial acrylic bone cements improves the tensile strength and reduces brittleness.

1.6 Main side effects

The main side effects of PMMA bone cements can be summarized as aseptic loosening, local temperature rise and release of toxic molecules.

1.6.1 Aseptic loosening

Stability of the prosthesis within the surrounding bone is the decisive factor for flawless functioning and longevity of the implants. Osteolysis is a multifactorial process stemming from host, prosthesis and surgical factors. Billions of wear particles are generated at material interfaces and are dispersed along the joint space, into bone and adjacent soft tissue, inducing an inflammatory reaction that leads to osteclast activation and finally to osteolysis. Over time osteolysis may progress to aseptic loosening and failure of the implant. Usually patients only become symptomatic when implant loosening, implant failure or periprosthetic fractures occur [32].

Aseptic loosening is the main cause of failure of cemented total hip arthroplasties [33]. According to Bartl R. et al. [32], the main factors involved in periprosthetic osteolysis and aseptic loosening are:

- Wear debris induced osteolysis: integration of the prosthesis into the surrounding bone can be hindered by a “foreign body reaction” induced by macrophages absorbing small particles, mainly polyethylene, PMMA and metallic debris, leading to activation of RANKL and OPG which then trigger osteoclastic activity. Finally, osteolysis and bone loss around the implant occur; - Micromovement of surfaces: implants that do not achieve adequate initial fixation will exhibit micromotion in response to load. The greater the area of friction the more osteoclasts are activated causing osteolysis around the implant which leads to fatigue failure at interfaces. When the distance between bone

and implant exceeds 150 µm, connective tissue membranes are formed between implant and bone as well as between implant and cement. These membranes hinder the osteo-integration of the prosthesis. Many biochemical mediators are involved: cytokines, prostaglandins, metalloproteases and collagenases;

- Inappropriate mechanical load and stress shielding: insertion of an implant leads to new biomechanical relationships between various regions of the surrounding bone and the implant. As a consequence of stress shielding, bone apposition and higher bone density occur in regions around the implant receiving high loads, whereas regions receiving lower stress loading react with bone loss. Appropriate load transmission is thus an essential factor in maintaining bone volume. Optimal load transfer is influenced by the design and stiffness of the implant. Bone loss around the implant due to stress shielding can account for up to 50 % of the former bone stock in underloaded regions. Finally, periprosthetic fractures can occur;

- Post-operative immobilization: the post-operative decrease in weight bearing results in local immobilization osteoporosis. Overall the post-operative bone loss mainly occurs in the first 6 months and can reach up to 50 % of the former bone stock;

- Operative trauma: thermal and mechanical necrosis caused by surgical procedure, type of bone cement and cementing techniques could alter bone quality.

Microstructural evaluation revealed that the newly formed bone adjacent to cement in loosened prostheses, is less mature than bone adjacent to stable prosthesis. The lattice parameters of bone apatite do not exhibit significant variations as compared to normal bone tissue. However, bone trabeculae at the interface with loosened prosthesis often present a strongly demineralised lamellar structure [34].

1.6.2 Local temperature increase

In commercial formulations, the powder part of the bone cement is pre-polymerized and this prevents severe polymerization reactions. Still, the highly exothermic polymerization of MMA cause a noticeable increase in the local temperature. The maximum curing temperature varies from 80 to 124 ºC depending on the L/P ratio, the chemical composition of the solid and liquid components, size of the PMMA particles, the concentrations of BPO and DMPT and the presence of additives [35]. Peak temperatures are about 25-30 ºC higher than the limit levels (50-60 ºC) and are considered, by some authors, a cause of thermal damage to tissues. It was reported that bone tissue degenerates when it is subject to 47 ºC for not less than 1 minute [36].

As previously mentioned, in section 1.5.3, PMMA bead size is an important factor for peak temperature. As the PMMA particle size decreases, the surface volume ratio increases. Therefore, the amount of polymer dissolved in the monomer increases, leading to higher viscosity of the dough. The increase in viscosity leads to a more difficult transfer of heat, causing an increase in the curing temperature. This also leads to a faster setting. Higher amounts of either the initiator or the activator increase the polymerization temperature and decrease the setting time. On the contrary, as the polymer/monomer ratio increases, both the polymerization temperature and the monomer residue decrease. This seems to be an advantage, were it not for the rise in viscosity, so that workability and penetration of bone cement into bone trabeculae become more difficult [1].

In some cases, PMMA bone cements with high setting temperatures may be desirable. Some surgeons treat the giant cell tumours of bone tissue by using the technique of aggressive curettage trough a large bone window followed by acrylic cement reconstruction [37]. The damage to the cells due to heat may be beneficial in reducing the rate of tumour recurrence [38].

1.6.3 Release of toxic molecules

MMA is toxic to bone. Its release to blood also causes a severe drop in blood pressure, due to the effect of MMA on blood vessels. The content of monomer in the cured cement is related to the monomer composition, glass transition temperature, polymerization temperature and type and concentration of initiator and activator used [1]. The proportion of non-converted residual monomer remaining in the polymerized cement is in the range of 2 to 6% just after hardening, decreases to approximately 0.5% or less after 2-3 weeks and thus remains for years [8].

Unreacted MMA leaks from the cement mantle into the surrounding tissues, causing toxic effect and impairing bone remodeling [5,39,40]; it also acts s as a plasticizer, influencing the mechanical properties of the cement [39,41].

Studies carried out in patients with hip metal prosthesis stabilized by acrylic bone cement pointed to statistically significant changes in clotting and fibrinolysis systems. It was reported that monomer or polymer release causes a tendency toward hypercoagulation and intravascular clotting. Fibrinolysis activation associated with these changes is secondary, resulting from, among others, the mechanism of action of free fribin monomers [42]. Bone cement affects the activity of the lysosomal enzymes in peripheral blood granulocytes. It was reported that after the endoprosthesis stabilization operation, statistically significant changes occur in the activity of lysosomal marker enzymes. Lability of lysosomal membranes appears with permeation of hydrolases into supernatant [43].

1.7 Additives

The main additives usually found in bone cements are radiopacifiers - such as barium sulphate (BaSO4) and zirconium dioxide (ZrO2) – and antibiotics.

1.7.1 Radiopacifiers

Since 1972, radiopaque materials such as BaSO4 or ZrO2 have been added to

the bone cement in order to provide radio-opacity. Usually, the addition of about 10% (w/w) of a radiopaque material to the powder part of PMMA-based cement, provides sufficient opacity for x-ray imaging [1]. Otherwise, the areas which bone cement occupies can be determined by using magnetic resonance imaging (MRI) [44]. Additional opacifiers are often added for interventional procedures such as vertebroplasty, in which visibility is key [20].

Addition of radiopaque agents to bone cements may have some disadvantages. Some studies have shown that the adition of radiopaque agents to PMMA enhance the macrophage-osteoclast differentiation and therefore they may contribute to bone resorption and aseptic loosening [45,46]. Furthermore, these agents evoke a significant pathological response in the surrounding tissue. Barium sulphate has been shown to intensify the release of inflammatory mediators in response to PMMA particles [47]. On the other hand, some commercial bone cements containing zirconium oxide (such as Palacos®, Implast® and Sulfix-6®) present a certain degree of radioactivity. These x-ray contrast media could remain in the body for decades, as components of the bone cement. Although no specific study has been reported, this situation seems to increase cancer incidence [1,48].

Radiopacifiers also have a significant effect on the mechanical properties of acrylic bone cements, depending on their size and morphology. It has been shown that while the addition of BaSO4 produces a decrease in tensile strength

of about 10% [49-51], the addition of ZrO2 does not affect this parameter. The

fracture toughness, which is unaffected by the inclusion of BaSO4, is increased

around 20% by the addition of ZrO2 [51]. Moreover, it has been shown that both

inorganic radiopacifiers enhance the fatigue crack propagation resistance [52]. Taking into consideration all these drawbacks, several alternatives to the traditional radiopaque agents were developed and studied [53-58]. One of these was based on the introduction of an iodine-containing methacrylate, 2, 5-diiodo-8-quinolyl methacrylate (IHQM), in the liquid phase [53]. The radiopacity of the resulting cements was confirmed and it was found to be good even for cements containing 5 wt% of IHQM in the liquid phase. In relation to the radiolucent cement the addition of barium sulphate produced a decrease in tensile strength, while not affecting the fracture toughness and improving the crack propagation resistance. When IHQM was used, although the tensile strength and the fracture toughness increased, the fatigue crack propagation resistance remained as low as in the radiolucent cement. Furthermore, it was observed that the incorporation of this monomer also lowered the peak temperature and slight increased the setting time [54].

Organo-bismuth compounds such as triphenyl bismuth (TPB) were also studied as radiopacifiers by Deb et al. [55]. TPB was incorporated in the bone cement matrix by two methods: 1) blending – addition to the polymer phase; 2) dissolution in the methyl methacrylate. The results showed that TPB at concentrations of 15% and 25%, by weight of polymer, did not affect the polymerization temperature and setting time. Furthermore, the addition of TPB via the dissolution method provided a statistically significant increase in the strain to failure in comparison with commercial acrylic cements containing barium sulphate, thus reducing the brittleness of the cement. The decay on mechanical properties after conditioning in water was also much less pronounced in the homogeneous TPB cements than in cements with BaSO4.

More recently, two bromine containing monomers, 2-(2-bromoisobutyryloxy) ethyl methacrylate (BIEM) and 2-(2-bromopropionyloxy) ethyl methacrylate (BPEM), were synthesized and characterized as being good candidates to be used as radiopacifiers [56]. The addition of BPEM decreased the maximum

temperature and increased the setting time, when compared with the radiolucent cement. It also decreased the glass transition temperature, increased the thermal stability, reduced the polymerization shrinkage and increased the compressive strength of the resultant material [57].

1.7.2 Antibiotics

The use of PMMA, like other biomaterials in the human body, entails the risk of attracting infectious microorganisms [59]. Although surgical operating rooms have sterile conditions some bacteria can pass through the protective barriers and contaminate the body tissues during the surgery. In order to prevent postoperative infections, small quantities of antibiotics can be added to bone cements [1]. This has been a topic studied by many research groups. Two most important points in the study are the release of the antibiotics from the cement and their effect on the final properties of the material.

1.7.2.1 The choice of the antibiotic

Antibiotics for incorporation in bone cements should have a broad antibacterial spectrum, including Gram-positive and Gram-negative pathogens, sufficient bactericidal activity, high specific antibacterial potency, low rate of primary resistant pathogens, minimal development of resistance during therapy, low protein binding, low sensitizing potential, marked water solubility facilitating its release from the bone cement and last, but not least, chemical and thermal stability [60,61]. Over the years several antibiotics have been evaluated by in

Most of the 18 different antibiotic-loaded bone cements currently available on the market, contain gentamicin as a sulphate – Table 1.2 [8]. Gentamicin shows a good release from bone cements, has a broad antimicrobial spectrum and good water solubility [75]. Several in vitro and in vivo studies have, however, indicated bacterial growth on antibiotic-loaded bone cements [60,76,77], with increased occurrence of gentamicin-resistant strains, thus necessitating improvements of the existing gentamicin-loaded bone cements. The increasing bacterial resistance to gentamicin has prompted renewed interest in the addition of further antibiotics to bone cements, such as tobramycin and cefuroxime [78,79]. From past experience, it is likely, however, that it will be only a matter of time until the bacteria develop a mechanism of resistance to overcome any new antibiotic which is incorporated in bone cements. Therefore, some research studies and industrial developments include the use of combinations of antibiotics (multidrug targeting). Multidrug targeting is assumed not only to be more powerful but also to prevent the emergence of resistant strains through the synergistic action of two antibiotics [80]. In Europe, one multidrug-loaded bone cement containing gentamicin and clindamycin, Copal – see Table 1.2, is commercially available. Combination of gentamicin and clindamycin in bone cement formulation has a theoretical antimicrobial effect on more than 90% of the bacteria common to infected arthroplasties. The release of gentamicin seems to be enhanced by the release of clindamycin in this cement [75]. This may be an effect of the extra antibiotic, which acts as a soluble additive that leaves a network of voids behind, enhancing further release [72]. Moreover, bone cements loaded with gentamicin and clindamycin or fusidic acid are more effective in preventing biofilm formation than bone cements with gentamicin alone [81].

Multidrug targeting may be effective in preventing resistance but using it is a difficult option in bone cements, as the release of the different antibiotics depends on many factors. For example vancomycin has a high molecular weight and shows poor release because it is trapped in the cement matrix [82]. Also, combinations of antibiotics must be carefully selected due to known cross-resistances. For this reason, the combination of gentamicin and tobramycin is

not advisable. And, there is always the possibility of an antagonistic effect in the different ways antibiotics act upon the bacterial life cycle [75].

Table 1.2: Commercially available antibiotic-loaded cements. Adapted from ref.

[1] and [8].

Bone cement Manufacturer Market

Antibiotic Simplex® (=AKZ) Howmedica Worldwide

Allofix® -G Sulzer Central Europe, CH

Cemex® -Genta HV Tecres South Europe; I

Cemex® -Genta LV Tecres South Europe; I

Cerafixgenta® Ceraver Osteal South Europe; France CMW 1 Gentamicin DePuy, J & J Worldwide

CMW 2 Gentamicin DePuy, J & J Worldwide CMW 2000 Gentamicin DePuy, J & J Worldwide CMW 3 Gentamicin DePuy, J & J Worldwide

Copal® Merck Worldwide

Genta C-ment® 1 E.M.C.M. B.V. Central Europe, G Genta C-ment® 3 E.M.C.M. B.V. Central Europe, G

Osteopal® G Merck Worldwide

Palamed® G Merck Worldwide

Subiton G Prothoplast Argentina

CH: Switzerland; G: Germany; I: Italy; J & J: Johnson & Johnson.

1.7.2.1 Mechanisms of antibiotic release from PMMA bone cements

It is generally accepted that antibiotics incorporated in PMMA cements are released to some extent, but the corresponding mechanisms are still debated. PMMA is capable of taking up very small quantities of solvent fluid into its outermost layers [73]. This fluid slowly releases the antibiotic into the surrounding tissue. The sustained release of antibiotics from bone cements is largely influenced by the penetration of dissolution fluids into the polymer matrix, which requires a certain porosity of the bone cement [69,70]. Initially, large amounts of antibiotic are released [83], as it is easily available at the surface of the cement. A point on which all studies have agreed is that the period of maximum antibiotic release is limited to the first few hours or days after implantation. The maximum effectiveness of the antibiotic might thus be expected to occur during this period of time [84]. Most, if not all, of the antibiotic is released from the surface [85,86]. Studies using methylene blue and gentamicin diffusion through or into acrylic bone cement discs show that the bulk is essentially impermeable [83]. In fact, many in vitro and in vivo tests have shown that only small amounts (3-15%) of the antibiotics incorporated in bone cements are actually released [83,86,87]. The release of antibiotics from PMMA bone cement is, therefore, largely a surface phenomenon.

The porosity of the polymer matrix depends on air entrapment, while wetting and stirring the cement powder, and on effects of monomer evaporation [28].

Cements with greater porosity would allow more antibiotic release. Therefore, methods of cement preparation designed to improve mechanical properties by decreasing the porosity (such as vacuum mixing and/or centrifugation) could have adverse effects on elution characteristics. In commercially available formulations the antibiotic is evenly dispersed throughout the cement. This uniform dispersion cannot be achieved when adding the antibiotic by hand to the bone cement. An uneven dispersion may reduce the mechanical properties of the cement and cannot always ensure sufficient antibiotic release [88].

1.7.2.2 Influence of the antibiotic on the properties of the cement

Some studies reported that small amounts of antibiotics in acrylic bone cements had no influence on the compressive and tension strengths. Mechanical tests demonstrated that adding a small amount of antibiotic did not reduce strength below acceptable standards nor change handling characteristics [62,89].

But contrary results were also reported. It was claimed that the addition of antibiotics caused a reduction in fatigue life [90] and that large amounts of antibiotics decrease the compressive and tensile strength of PMMA bone cements [91].

1.8 Improvement of PMMA bone cements

A wide variety of modifications of conventional PMMA bone cements were and still are currently being investigated. The main purpose of the studies is the improvement of either the mechanical or the biological/clinical properties.

1.8.1 Fiber reinforcement

As previously mentioned, PMMA was first successfully used for implant fixation of hip prostheses in 1958 and since then it has been used also for stabilization of knee, shoulder, elbow and others. However, in the first applications many of the prosthesis were removed because fractures of the surrounding PMMA cement were significant [1,3].

Efforts for improving the properties of bone cement include dispersing small quantities (typically 1-2 vol%) of reinforcing materials such as graphite [92,93], carbon [94,96], aramid [95,97], polyethylene [99], PMMA [100-102], ultra-high molecular weight polyethylene [103], titanium [104,105] or steel [106] fibers in the cement matrix.

In 1975, Knoell and co-workers [92] reported that adding small amounts of graphite fibers would enhance the mechanical properties of PMMA cements. The results of Robinson and co-workers [93] suggest that graphite reinforcement of plain cement produced a 32 % increase in fracture toughness. In 1979 Wright and Trent [95] studied aramid fiber reinforcement of bone cement. By adding up to 7 % of aramid fibers with average length of 1.3 cm, the tensile strength could be increased from 30.8 to 42.8 MPa and the fracture toughness from 1.53 to 2.85 MPa m1/2.

Saha and Pal [96] observed that increased deformation due to creep in 24 h was reduced from 70 to 45 % by fiber reinforcement with 6 mm long and 8 µm

thick chopped carbon fibers. Stress relaxation in reinforced specimens was greater compared to unreinforced cement. The same authors [97] found that adding 4 % aramid fibers to 1 % carbon fiber reinforced cement provided best improvement in compressive strength and adding 2 % aramid fibers provided best improvement in compressive elastic modulus.

Pourdeyhimi et al. [98] studied Kevlar 29® fiber reinforcement and observed improved fracture toughness, even at very low fiber contents. By adding up to 7 % Kevlar 29® fibers, the flexural strength increased from 67.4 to 82.6 MPa, flexural modulus increased from 1.28 to 1.30 GPa, fracture toughness increased from 1.47 to 2.61 MPa m1/2.

Wagner et al. [99] studied the use of high-performance polyethylene fibers surface-activated by oxygen plasma for better adhesion in order to reinforce PMMA bone cement, but no significant reinforcement was observed.

Gilbert and co-workers [100,101] studied a self-reinforced composite of PMMA (SRC-PMMA) consisting of high strength, high ductility PMMA fibers embedded in a PMMA matrix. The tensile strength, tensile modulus and tensile strain to failure were significantly greater for SRC-PMMA, as compared to commercial PMMA cements.

Fibers of PMMA were also developed by Wright et al. [102] by extrusion of PMMA melt. The effects of melt viscosity on fibre properties were examined and the processing parameters optimized. Fibers with ultimate tensile strength values in the range of 60 to 225 MPa, elastic modulus values from 1.5 to 3.5 GPa and strain to failure from 10 to 40 % were obtained. Use of the fibers in bone cement reinforcement was proposed.

Kotha et al. [106] investigated the changes the fracture properties and temperatures generated in the ASTM F451 tests, by the addition of 316L stainless steel fibers with volume fractions of 5-15 % by volume and aspect ratios of 19, 46 and 57. Increasing the volume fraction of fibers resulted in significant increases in the fracture toughness of the steel-fiber-reinforced cement, up to 2.63 times the control value. No clear trend in the fracture toughness was discerned for increasing aspect ratios of the reinforcements. A