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Acta Biomaterialia 5 (2009) 356–362

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New PMMA-co-EHA glass-filled compositesfor biomedical applications: Mechanical properties and bioactivity

Poliana Lopes a,*, Marcelo Corbellini b, Barbara Leite Ferreira a, Nuno Almeida a,Marcio Fredel b, Maria Helena Fernandes a, Rui Correia a

a CICECO, Department of Glass and Ceramic Engineering, University of Aveiro, Campus Universitario de Santiago, 3810-193 Aveiro, Portugalb Department of Mechanical Engineering, Federal University of Santa Catarina-UFSC, Trindade/Caixa Postal: 476–Florianopolis/SC-88040-900, Brazil

Received 16 January 2008; received in revised form 19 May 2008; accepted 17 July 2008Available online 29 July 2008

Abstract

A bioactive glass of the 3CaO�P2O5–MgO–SiO2 system was incorporated as a filler into poly(methylmethacrylate)-co-(ethylhexylac-rylate) (PMMA-co-EHA) copolymer. The effect of filler proportion (0, 30, 40 and 50 wt.%) on the bending properties was evaluated anda maximum flexural strength of 29 MPa coupled with an elastic modulus of 1.1 GPa was obtained at an intermediate filler concentration(30 wt.%). These values are slightly higher than those usually reported for human cancellous bone. The in vitro bioactivity was assessedby determining the changes in surface morphology and composition after soaking in simulated body fluid (SBF, Kokubo solution).Inductively coupled plasma was used to trace the evolution of ionic concentrations in the SBF solution, namely Ca and P. X-ray diffrac-tion and scanning electron microscopy confirmed the growth of spherical calcium phosphate aggregates on the surface of composites,indicating that the composites are potentially bioactive.� 2008 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Keywords: Acrylic bone cements; Glass filler; Mechanical properties; Bioactivity

1. Introduction

Acrylic polymers have been extensively used, as bonecement, in the treatment of bone defects and prostheticfixation. In these applications the cement acts as an inter-mediary phase, fixing the implant to the bone, transmit-ting the applied force and body weight uniformly to thetissue and functioning as a load-bearing material [1,2].It is therefore very important that the cement is able tomaintain its mechanical properties over a long period oftime in vivo. Chaplin [3] suggested that the mechanicalproperties of poly(methylmethacrylate) (PMMA) bonecement implanted for 15–24 years appear to be compara-

1742-7061/$ - see front matter � 2008 Acta Materialia Inc. Published by Else

doi:10.1016/j.actbio.2008.07.012

* Corresponding author. Tel.: +351 234 370 354.E-mail address: polianap@ua.pt (P. Lopes).

ble to freshly made PMMA, with no indication of poly-mer deterioration.

Despite the successful application of acrylic polymers,some well-known drawbacks are associated with theiruse, such as non-bone-bonding capability, relatively lowmechanical strength and high curing temperatures [4,5].These characteristics can cause serious complications invivo, such as necrosis of the surrounding tissues and evenloosening of the implant [1,6–8]. In order to overcomethese problems, efforts have been made to improve theproperties of bone cements (e.g. their mechanical strengthand bioactivity) [9–15].

Incorporation of a second reinforcing phase into thecement, such as glass, may promote bioactivity, i.e. bonegrowth around the implant, resulting in increased longev-ity of the prosthesis [16]. Furthermore, the addition offiller to PMMA is considered to be a potential methodto improve its mechanical properties [17]. Hence, the idea

vier Ltd. All rights reserved.

Table 1Composition of composites produced (wt.%)

Samples* Solid component Liquid component

PMMA VH1 MMA EHA

C0M 66.6 0 16.7 16.7C3M 45.9 19.7 16.7 16.7C4M 39.3 26.3 16.7 16.7C5M 32.8 32.8 16.7 16.7

* BPO 2 wt.% of the solid component and DMPT 1 wt.% of the liquidcomponent.

P. Lopes et al. / Acta Biomaterialia 5 (2009) 356–362 357

of combining a bioactive glass with another material(polymer) to produce a composite with mechanical prop-erties comparable to those of natural bone and with abioactive character has been explored in a number ofworks [18–21].

It is known that sol–gel glasses of the system studied inthis research show the formation of apatite in simulatedbody fluid (SBF), indicating that these composites can bebioactive [22]. However, their mechanical properties havenot been disclosed and the effect of this filler is not known.Studies reported in the literature concerning the mechani-cal properties of acrylic bone cement reinforced by bioac-tive ceramic have revealed a large variation in bendingbehaviour [9,16,23]. In order to achieve an improvementin mechanical properties, it is generally believed that theinterfacial adhesion between the filler and the matrix isan important factor [24,25].

The aim of this study was to prepare materials withthe desired mechanical properties and in vitro bioactivity.For this purpose, composites with different amounts ofglass were produced and the effect of composition onthe bending properties and apatite forming ability wasevaluated.

2. Materials and methods

2.1. Materials

Methylmethacrylate (MMA) and ethylhexylacrylate(EHA) were purchased from Aldrich Co. MMA was puri-fied by washing repeatedly with sodium hydroxide solutionsaturated with sodium chloride, in order to remove theinhibitor. PMMA beads (Aldrich Co.) were milled in arotor mill (Retsch ZM 200) to a mean particle size of160 lm and used as a solid component. Benzoyl peroxide(BPO, Merck) initiator and N,N-dimethyl-p-toluidine(DMPT, Fluka) activator were used as received for thepolymerization reaction. The glass (VH1) with nominalcomposition 20.15 MgO, 29.85 SiO2, 27.12 CaO and22.88 P2O5 (wt.%) was melted in a platinum crucible at1550 �C for 2 h, quenched in water and milled in an agatemill to a powder with an average particle size of 8.7 lm.

2.2. Preparation of composites

The composites were prepared by mixing the solid(BPO, VH1 and PMMA) and liquid (MMA, EHA andDMPT) components, employing a solid/liquid ratio of2:1. MMA and EHA monomers were used in the propor-tion of 1:1, with the BPO initiator and DMPT activatorin a concentration of 2 wt.% (solid phase) and 1 wt.%(liquid phase), respectively. The final nominal compositionof the composites is presented in Table 1, where C0M = 0%glass; C3M = 30%; C4M = 40% and C5M = 50% glass,referring to the wt.% of the solid component. The mono-mers were first mixed in a glass recipient and then BPOwas dissolved in this liquid mixture. The solid precursors

(PMMA and VH1, pre-mixed in an agate mortar) wereafterwards added to the liquid monomers. Finally, DMPTwas added and the mixture was hand-mixed for 1 min. Thedough was poured into a mould, between two 60 � 60 mm2

rectangular glass plaques sealed with silicone elastomer.The mix was cured in air at room temperature (23 ±2 �C), and removed from the mould to produce slabs withthe approximate dimensions 40 � 20 � 5 mm3.

2.3. Mechanical properties

The various composites were tested by bending. All thesamples were prepared in the same way to nullify any influ-ence of the preparation technique upon the mechanicalproperties. Bending test specimens were produced by cut-ting the original composite slabs into beams of 40 � 8 �5 mm3 using a band saw (Struers Secotom-10). The sur-faces were ground with 1200 grit (FEPA) silicon carbidegrinding paper using a grinding machine. Three-pointbending tests were performed at room temperature in aZwick/Roel Z020 machine at a crosshead speed of1 mm min�1, following a approach similar to thatdescribed by Daglilar [4] and Puska [26,27]. Six or sevensamples were tested for each composition.

The elastic modulus (E) was calculated by extractingdata from the initial linear portion of the force vs. displace-ment curve, using the relationship

E ¼ ð1=4ÞL3b�1h�3F ðDyÞ�1

where L is the distance between supports, F is force, b andh are the width and height of the specimen, respectively,and Dy is displacement.

Values were expressed as the mean value ± standarddeviation and compared using one-way analysis of variance(ANOVA). Differences between the composites were thencompared using Bonferroni’s post-hoc means comparison;p values less than 0.05 were considered significant.

Fracture surfaces of test specimens were examined byscanning electron microscopy (SEM) in a Hitachi S4100microscope in order to identify the nature of the fractureand any differences between the compositions. Porositywas measured using a liquid displacement method (Archi-medes’ principle), with distilled water as the displacementliquid due to its ease of penetration into the pores andnon-induction of damage in the material.

358 P. Lopes et al. / Acta Biomaterialia 5 (2009) 356–362

2.4. Bioactivity in vitro

The assessment of in vitro bioactivity was carried out bysoaking pieces of 5 � 5 � 5 mm3, mounted vertically in10 ml of SBF, in sterile polyethylene containers maintainedat 37 �C. The SBF solution was prepared by dissolvingNaCl, NaHCO3, KCl, K2PO4�3H2O, MgCl2�6H2O, 1.0 MHCl, CaCl2, Na2SO4 and (HOCH2)3CNH2 in ultrapurewater [28]. The solution had a similar composition to thatof human plasma (Table 2) and was previously filteredthrough a Milipore 0.22 lm system. Soaking periods are1, 3, 7, 14 and 21 days. After immersion, the specimenswere removed from the fluid and the ionic concentrationsof P and Ca were determined for each period by inductivelycoupled plasma spectroscopy. Formation of the apatitelayer was identified by X-ray diffraction (XRD) and SEM.

3. Results and discussion

3.1. Mechanical properties

Fig. 1 shows representative nominal stress–strain bend-ing curves for the matrix and the composites. The unfilledmaterial exhibits a ductile curve, whereas the presence ofglass decreases the strain to fracture, increases the elasticmodulus in the low stress regime and promotes the appear-ance of a plateau, which is supposed to result from a filler/matrix interfacial disruption process.

The strength at maximum load, the elastic modulus andthe porosity for each composition are plotted in Fig. 2.One-way ANOVA revealed that the differences in the meanbending strength were significant (p < 0.05) except thosebetween C0M and C5M (p > 0.05). The elastic modulusfor the glass-reinforced composites were considered to be

Table 2Ionic concentrations (mM) of SBF and human blood plasma

Na+ K+ Mg2+ Ca2+ Cl� HCO3� HPO4

2� SO42�

Plasma 142.0 5.0 1.5 2.5 103.0 27.0 1.0 0.5SBF 142.0 5.0 1.5 2.5 147.8 4.2 1.0 0.5

Fig. 1. Stress–Strain curves of composites.

similar (p > 0.05) but significant when compared to thatof C0M (p < 0.05).

The mechanical values for all composites were slightlyhigher than those reported for cancellous bone: theYoung’s modulus of cancellous bone ranges from 0.05 to0.5 GPa, and its bending strength ranges from 10 to20 MPa [29]. A 30 wt.% addition of glass brings the flexuralstrength to the 30 MPa level and the elastic modulus toabove 1 GPa while retaining a 4% ultimate strain (Fig.1). For materials containing proportions of filler higherthan 30 wt.%, the strength decreased steadily with increas-ing glass content, thus attaining similar values to those ofthe matrix material for 50 wt.% filler. Porosity was reducedwith increasing amount of particulate filler.

The observed behaviour can be explained through thedegree of adhesion between the matrix and the filler[25,30]. During the exothermic polymerization reaction,thermal stresses are generated due to the differences inthe thermal expansion coefficients of the composite compo-nents. The shrinkage of the PMMA matrix is greater thanthe shrinkage of the glass particles, resulting in a circumfer-ential tensile stress adjacent to the glass [12]. This leads to aweak bonding between the matrix and the filler and, hence,polymer–glass detachment can occur (Fig 3). Previousstudies have indicated that poor interface adhesionbetween the constituents of composites can be responsiblefor their decreased mechanical properties [9,17,31].

The increase in glass content can aggravate this prob-lem, owing to inhomogeneous distribution and agglomera-tion. When the filler begins to aggregate, it behaves likevoids in the cement, thereby weakening it; therefore thestrength is reduced even if the filler is strong enough toincrease the elastic modulus of the material [10,24,32].

The addition of filler produced a significant change inthe fracture surface of the materials. SEM micrographs(Fig. 4) shows a macroporous unreinforced matrix, whilethe reinforced polymer exhibits a reduction in porositywith the increase in the proportion of glass: the porositydecreased from 9.8 ± 2.5% for the unfilled composite to7.4 ± 1.7% for the composite filled with 50 wt.% glass. Sim-ilar results were achieved by the incorporation of glassspheres within bone cement [16]. Pores observed in thecured composites may be attributed to air entrapped dur-ing mixing, monomer evaporation over polymerisation,shrinkage around particles and agglomerates, coupled withweak interfacial bonding.

Although in the present case the filler has shown pooradhesion with the thermoplastic matrix, no strength degra-dation is verified in the composites relative to the matrix, sothat the beneficial effect of the filler in eliminating macrop-orosity over the interfacial decohesion process occurs up to50 wt.% filler content.

It is also noted that increasing the filler content beyond30 wt.% has no significant effect on the elastic modulus.This confirms that some form of microstructural degrada-tion should occur during processing composites with higheramounts of glass particles since, other factors being equal,

Fig. 2. Nominal bending strength, elastic modulus and porosity vs. wt.% glass.

Fig. 3. Composites (a) C4M and (b) C5M; the arrows indicate a void formed by poor adhesion at the filler–matrix interface.

Fig. 4. Fracture surfaces of composites (a) C0M and (b) C5M.

P. Lopes et al. / Acta Biomaterialia 5 (2009) 356–362 359

higher filler fractions are expected to result in progressivelyhigher stiffness.

3.2. Bioactivity

The bioactivity test showed that for all filler amounts thecomposite developed an apatite layer on its surface afterimmersion in an SBF solution. Further information onthe structure of the apatite layer was obtained by XRDand SEM measurements. Fig. 5 shows the X-ray diffractionpatterns of composites C5M and C3M, which exhibited thecharacteristic peaks at 2h = 26 and 32�, attributable tohydroxyapatite [14,33]. The composite C4M presented anXRD pattern similar to that of C5M. These results indicate

that apatite crystals are deposited on the surface of materi-als, being faster for composites with higher glass content.The appearance of only two peaks derived from an apatiticphase may be because of the small amount, small size andlow crystallinity of formed calcium phosphate [34]. How-ever, the intensity of these peaks improves with increasingimmersion time, due to the growth on the composite sur-face of an apatite layer of enhanced crystallinity with time.For C3M the broad band characteristic of hydroxyapatitewas only detected after 21 days of immersion.

The SEM studies of the composite during immersion inSBF for various periods are shown in Fig. 6. It can be seenfrom these micrographs that a dense apatite layer, com-posed of spherical particle aggregates [22,35], was formed

Fig. 5. DRX patterns of composite (a) C5M and (b) C3M.

Fig. 6. SEM micrographs: (a) C3M, 7 days; (b) C3M, 14 days; (c) C3M, 21 days; (d) C5M, 7 days; (e) C5M, 14 days; and (f) C5M, 21 days.

360 P. Lopes et al. / Acta Biomaterialia 5 (2009) 356–362

on the surface of the composite within 7 days for C5M andC4M (SEM micrographs not shown for C4M). For com-posite C3M, its surface was covered with apatite only after14 days of immersion.

The variations in Ca and P concentrations relative tosoaking times in SBF after immersion of the compositesare shown in Fig. 7. The calcium and phosphorous ionsare required for the generation of the calcium phosphatelayer on the surface of the composite. During the growthof this layer, consumption of Ca and P from SBF occurs,resulting in the observed decrease in the concentrationsof these ions in the solution. The increase in Ca and Pfor C4M is thought to result from the detachment of thephosphate deposit [36].

4. Conclusions

The addition of 30 wt.% of 8.7 lm (mean particle size)glass powder particles to a PMMA-co-EHA matrix resulted

in significant increases in flexural strength and elastic mod-ulus, increasing the mechanical performance of the materialto the upper level of the values reported for cancellous bone.The incorporation of glass reduces macroporosity duringprocessing but is likely to introduce other discontinuities,namely microvoids at the particle–matrix interfaces. In thisway, particle concentrations of 40 and 50 wt.% do not rep-resent improvements in mechanical properties.

These composites are bioactive in acellular medium, asshown by the XRD and SEM analyses, confirming thepresence of a mineralized deposit after immersion inSBF. The in vitro formation of an apatite layer on the sur-face of the materials is an indication of the potential bone-bonding ability of the developed composites. Possibleimprovements in the material reside in the processing stage,through changes in the proportions of co-monomers(which will affect the matrix stiffness), in powder contentand particle size, and in the use of an interfacial bondingagent.

Fig. 7. Variations of ionic concentrations of the composites during the soaking period: (a) phosphorous, (b) calcium.

P. Lopes et al. / Acta Biomaterialia 5 (2009) 356–362 361

Acknowledgements

B.L.F. and P.L. thank the Portuguese Foundation forScience and Technology (FCT) for their doctoral researchGrant (SFRH/BD/17389/2004) and (SFRH/BD/27961/2006), respectively. This work was supported by Universityof Aveiro–Research Institute (Project CTS 05). Theauthors also thank Conceic�ao Costa (X-ray Diffraction),Marta Ferro (Electron Microscopy) and Flavia Almeida(Mechanical Properties) for valuable technical and profes-sional assistance.

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