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Polyhedral oligomeric silsesquioxane (POSS)–poly(ethylene glycol) (PEG) hybrids as

injectable biomaterials

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2012 Biomed. Mater. 7 035013

(http://iopscience.iop.org/1748-605X/7/3/035013)

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Biomed. Mater. 7 (2012) 035013 (9pp) doi:10.1088/1748-6041/7/3/035013

Polyhedral oligomeric silsesquioxane(POSS)–poly(ethylene glycol) (PEG)hybrids as injectable biomaterialsJohanna Engstrand, Alejandro Lopez, Hakan Engqvistand Cecilia Persson1

Department of Engineering Sciences, Division of Applied Materials Science, The Angstrom Laboratory,Uppsala University, Uppsala, Sweden

E-mail: cecilia.persson@angstrom.uu.se

Received 26 October 2011Accepted for publication 1 March 2012Published 11 April 2012Online at stacks.iop.org/BMM/7/035013

AbstractOne of the major issues with the currently available injectable biomaterials for hard tissuereplacement is the mismatch between their mechanical properties and those of the surroundingbone. Hybrid bone cements that combine the benefits of tough polymeric and bioactiveceramic materials could become a good alternative. In this work, polyhedral oligomericsilsesquioxane (POSS) was copolymerized with poly(ethylene glycol) (PEG) to forminjectable in situ cross-linkable hybrid cements. The hybrids were characterized in terms oftheir mechanical, rheological, handling and in vitro bioactive properties. The results indicatedthat hybridization improves the mechanical and bioactive properties of POSS and PEG. TheYoung moduli of the hybrids were lower than those of commercial cements and more similarto those of cancellous bone. Furthermore, the strength of the hybrids was similar to that ofcommercial cements. Calcium deficient hydroxyapatite grew on the surface of the hybridsafter 28 days in PBS, indicating bioactivity. The study showed that PEG–POSS-based hybridmaterials are a promising alternative to commercial bone cements.

1. Introduction

Injectable, in situ curing biomaterials are used in a varietyof applications such as spinal fracture stabilization, bone-void filling in tumour patients or for improving the anchoringof fixations in, for example, the tibial plateau [1, 2].Currently, there are two main types of commercially availablebone cement. Acrylic bone cements based on poly(methylmethacrylate) (PMMA) are widely used in clinics forprosthesis fixation and spinal stabilization [1]. Ceramiccements based on calcium phosphate (CP) or calcium sulfate(CS) are mainly used as bone-void fillers but have also beenevaluated for use in vertebroplasty and kyphoplasty [1, 3]. Bothtypes of cements present advantages and disadvantages. Thehigh Young’s modulus of PMMA compared to osteoporoticcancellous bone has been associated with additional fractures

1 Author to whom any correspondence should be addressed.

[4, 5]. Furthermore, the high polymerization temperaturesas well as residual monomer may be detrimental to thesurrounding tissue [6, 7]. On the other hand, advantages ofPMMA-based cements include their high mechanical strength,ease of injection and ability to rapidly set in situ via radicalpolymerization [8]. Ceramic cements, on the other hand,are brittle and have a low fracture toughness, which haslimited their use in load-bearing applications despite their goodbiological properties [9]: they can be osteoconductive, i.e., theyserve as a template for the formation of new bone [9].

The ideal bone cement should combine the benefits ofboth polymeric and ceramic systems; high strength, adequatestiffness and bioactivity are some of the most desired properties[1]. In fact, polymer–ceramic composites have been widelyinvestigated, both as in situ setting cements [10–12] andas scaffolds moulded outside the body [13–18]. However,these composite materials consist of fillers and matrix, andin the case of poor bonding between the two, the mechanical

1748-6041/12/035013+09$33.00 1 © 2012 IOP Publishing Ltd Printed in the UK & the USA

Biomed. Mater. 7 (2012) 035013 J Engstrand et al

properties of the composite would be negatively affected. Theliterature shows that nanosized particles bond better to thematrix compared to microsized particles, due to the highersurface-area-to-volume ratio, resulting in a higher surfaceenergy and activity [19]. A way to further improve theseproperties could be to covalently bind both types of materialsto form hybrids [20, 21]. However, most research today onhybrids is focusing on materials that have to be preparedoutside of the body, after which they can be implanted[20–27]. To the authors’ knowledge, only two in situ settinghybrid materials have been studied to date: both are PMMA-based cements with a hybrid polymer–silica system as thepowder phase, giving silica-containing cements with tightlinkage to the polymer matrix [28, 29]. However, as previouslymentioned, PMMA may give rise to excessive temperaturesduring curing. Furthermore, the hybrid materials studied hadas high or even higher elastic moduli than commonly usedPMMA formulations, which may cause an abnormal loaddistribution in vivo and lead to further complications, asmentioned above [4, 5].

In this work, a hybrid material was produced in orderto maximize the benefit from both polymeric and ceramicproperties. Polyhedral oligomeric silsesquioxane (POSS R©)functionalized with methacrylic groups were chosen asceramic nanoparticles for this purpose. In this form, POSShas the potential to act as a cross-linking core for thehybrids, due to its multi-functionality and hence improvethe properties, both through its nanometric size as wellas its capacity to covalently bind to another polymerizablematerial through the methacrylic groups. The first POSS wasisolated in 1946, starting from methyltrichlorosilane [30].These types of molecules are synthesized through hydrolysisand condensation reaction of organotrichlorosilanes, similar tothe sol–gel route used to synthesize pure silica networks [31].Their abbreviated chemical formula is (RSiO1.5)n, where R canbe almost any organic substituent and n is variable. If n equals8, 10 or 12 the structure has a specific cage-like structure. Theirsizes range between 1 and 3 nm in diameter and can be thoughtof as the smallest possible silica particles [32]. Unlike puresilica, the POSS cages can be modified with multiple organicsubstituents to tailor their compatibility with other polymersand organic components [31]. Previous research has shown thatthe incorporation of POSS into polymer matrices significantlyenhances various physical properties [31]. Furthermore, silicamay be beneficial for the bioactivity of the material [28, 33].

For the matrix, poly(ethylene glycol) (PEG) was choseninstead of the commonly used PMMA. PEG is a highlyhydrophilic polymer approved by the US Food and DrugAdministration (FDA) for use in biomedical applications[34, 35]. It is non-toxic, non-immunogenic, non-antigenic, andif released in vivo it can pass through the bloodstream withoutdegradation [34, 35]. PEG is relatively inexpensive and readilyavailable in various molecular weights and functionalized withdifferent end-groups, e.g., acrylic groups. Elastic moduli muchlower than PMMA have been reported for PEG [36, 37], whichsuggests that it may be adequate as a matrix material for thepurpose of adapting the mechanical properties to those of bone.

The purpose of this study was to synthesize and evaluatehybrid materials based on methacryl-POSS (POSS) and

Figure 1. Chemical structures of the PEGda and methacryl-POSSprecursors that were used to fabricate the hybrids.

poly(ethylene glycol)-α,ω-acrylate (PEGda) for potential useas bone cements. Both the POSS and PEGda studied containpolymerizable side groups that allow for their incorporation inthe hybrid during the setting of the cement, i.e., this reactioncould be considered a co-polymerization. The followingproperties of the hybrids were determined as they weredeemed of interest to the application: Young’s modulus(E), maximum compressive strength (σ max), strain at break(εmax), zero shear viscosity (η0), maximum polymerizationtemperature (Tmax) and setting time (ST). Additionally,in vitro bioactivity tests were performed and the materials werecharacterized using scanning electron microscopy (SEM),electron dispersive spectroscopy (EDS) and grazing incidencex-ray diffraction (GI-XRD).

2. Materials and methods

2.1. Materials

All chemicals were acquired from Sigma-Aldrich (Sigma-Aldrich, St Louis, MO, USA), unless otherwise specified, andused as received.

2.1.1. Sample preparation. Methacryl-POSS R© cage mixture(n = 8,10,12) (Hybrid Plastics, Inc., City, United States,pure) and PEGda (700 g mol–1, pure) were polymerized atroom temperature (RT) in the presence of a radical initiator,benzoyl peroxide (BPO), and a catalyst, N,N-Dimethyl-p-toluidine (DMPT) (figure 1), compounds commonly includedin PMMA bone cement formulations. The amounts of BPOand DMPT used were kept equimolar and similar to thoseused in commercial PMMA bone cement formulations [38].

The following procedure was used to prepare all samplesto achieve homogeneous solutions and to avoid entrappedair that could affect the mechanical properties. The amountsof PEGda and POSS were varied from 100 wt% PEGdato 100 wt% POSS in steps of 25 wt%, resulting in fivecompositions. The samples were named PEGdaxPOSSy where

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Biomed. Mater. 7 (2012) 035013 J Engstrand et al

(A) (B)

Figure 2. Photographs of (A) the dual syringe system including equal volumes of homogeneous PEGda/POSS prepolymer mixture and (B)simultaneous mixing and injection of the material into a standard Teflon R©mould.

Table 1. Sample compositions and dissolution time of BPO in the PEGda/POSS prepolymer mixtures.

Formulation PEGda (g) POSS (g) BPO (g) DMPT (μL) Dissolution time (h)

PEGda100POSS0 20 0 0.2 170 20PEGda75POSS25 15 5 0.2 170 20PEGda50POSS50 10 10 0.2 170 20PEGda25POSS75 5 15 0.2 170 72PEGda0POSS100 0 20 0.2 170 480

x is the wt% of PEGda and y is the wt% of POSS. Appropriateamounts of POSS and PEGda (table 1), were mixed thoroughlyfor 2 min at 64 000 revolutions per minute (rpm) usingan Ultra-Turrax R© (IKA, City, Germany). This PEGda/POSSprepolymer mixture was split in half and poured into separatevials. BPO was added to one of the vials and mixed for 30 swith the Ultra-Turrax R© at 64 000 rpm. DMPT was added tothe other vial. The BPO-to-DMPT molar ratio was kept equalto 1 for all samples. Both vials were placed on a shakingtable at a speed of 300 rpm until the BPO was completelydissolved in the PEGda/POSS prepolymer mixture (table 1).Upon dissolution, the final prepolymer mixtures were pouredinto separate compartments of a dual syringe, and injected intometal or Teflon R© moulds, to achieve the desired shape uponcuring (figure 2).

2.2. Methods

2.2.1. Compression test. Specimens for compression testswere made in Teflon R© moulds to achieve a cylindrical shapeof 6 mm in diameter and 12 mm in height, as stipulated in theASTM F451–08 [39] standard. Between five and nine sampleswere tested for each group. The specimens were tested after24 h of curing at RT, using an AGS-H universal materialstesting machine (Shimadzu, Kyoto, Japan) at a crossheaddisplacement rate of 20 mm min−1. Young’s modulus (E),maximum strength (σ max) and strain at break (εmax) wereobtained from the load-versus-displacement data. Statisticalanalysis was done in IBM SPSS Statistics 19 (IBM, Chicago,IL, USA) using ANOVA at a significance level of α =0.05. Tamhane’s post hoc test was used since homogeneityof variances could not be confirmed.

2.2.2. Rheometry. A parallel plate AR 2000 rheometer (TAInstruments, USA) was used with a steel geometry of 40 mm,

Table 2. Composition of Dulbecco’s D 8662 phosphate bufferedsaline, pH = 7.

Salt Amount (g L−1)

CaCl2 · 2H2O 0.133MgCl2 · 6H2O 0.1KCl 0.2KH2PO4 (anhydrous) 0.2NaCl 8Na2HPO4 (anhydrous) 1.15

a gap of 500 μm and a lower plate temperature of 37 ◦C. Anoscillation procedure was run in the stress sweep mode at aconstant frequency of 1 Hz to determine the linear viscoelasticrange. A flow experiment in a continuous ramp mode was thenperformed to determine the zero shear viscosity (η0).

2.2.3. Handling. The maximum polymerization temperature(Tmax) and ST were determined according to ASTM F451–08 [39]. The samples were cured in a Teflon R© mould thatincluded a thermocouple to measure the temperature inside thesamples. The maximum polymerization temperature and thetime when this temperature was reached were recorded. TheST was calculated as the midpoint between RT and tmax (timeat Tmax).

2.2.4. Bioactivity. The bioactivity test was made in accor-dance with ISO standard 23317 [40], with the exception ofusing commercially available Dulbecco’s phosphate bufferedsaline (PBS, Sigma Aldrich, St Louis, MO, USA) (table 2)instead of the recommended simulated body fluid (SBF).PBS contains more phosphates than SBF, while SBF con-tains other ions such as sulfate and carbonate. PBS has fre-quently been used instead of SBF to confirm bioactivity [41].

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Figure 3. Representative stress–strain curves in compression for allgroups.

The samples for the bioactivity tests were made in cylindri-cal metal moulds of 6 mm in diameter and 2 mm in height.For each group three samples were made. The samples werecured for 5 h at RT before being removed from the moulds,immersed in 15 ml of PBS and stored in a sealed 15 ml fal-con tube in an oven (Carbolite, Hope Valley, UK) at 37 ◦Cfor 28 days. Two reference groups were used. Firstly, oneof each PEGdaPOSS sample was stored in 15 ml of double-distilled water; secondly, PMMA discs cut from a PMMA rodof 6 mm in diameter were stored in PBS. After the storageperiod, all samples were rinsed in ethanol and dried in vac-uum for 24 h before analysing their surface by SEM (LEO1550, Carl Zeiss, Oberkochen, Germany) and EDS. Prior tothe SEM/EDS analysis, a thin gold/palladium coating wassputtered onto the surface. As a reference a pure hydroxyap-atite (HA) sample with a known calcium-to-phosphate ratio of1.67 was analysed with EDS to confirm the calibration of theinstrument. An uncoated PEGda25POSS75 sample was alsoanalysed with GI-XRD using Cu K alfa radiation with a beamincidence angle of 2◦ and a 2θ scan from 25◦ to 33◦ (step size:0.03◦, step time: 20 s, scan rate: 0.09 deg min−1), where themost intense peaks for HA (PDF#01–086-0740) are found inorder to confirm any presence of such crystals.

3. Results

3.1. Compression test

Figure 3 shows the representative stress–strain curves ofthe cements. E and σ max peaked at 819 and 76 MPa forPEGda25POSS75 and PEGda50POSS50, respectively. Onthe other hand, the lowest combined values were given bythe homopolymers, PEGda0POSS100 and PEGda100POSS0,which gave E = 254 MPa, σ max = 18 MPa and E = 28 MPa,σ max = 12 MPa, respectively (table 3). The strain at break ofthe materials varied between 9.0% and 32.7% and increasedwith the amount of PEG (table 3).

The statistical analysis showed a significant effect ofthe composition on all three variables: modulus, strength

Figure 4. Viscosity versus shear rate for the PEG/POSS prepolymermixtures at 37 ◦C. The extrapolation of the curves to zero sheargives the zero shear viscosity.

Table 3. Results from the compression test. Young’s modulus (E),maximum compressive strength (σ max), strain at break (εmax) andnumber of samples (N) tested are shown in the table.

Formulation E (MPa) σ max (MPa) εmax (%) N

PEGda0POSS100 254 ± 19 18 ± 2 9.0 ± 1.6 5PEGda25POSS75 819 ± 39 50 ± 22 9.9 ± 4.0 9PEGda50POSS50 456 ± 7 76 ± 18 20.0 ± 3.2 6PEGda75POSS25 147 ± 3 57 ± 31 25.3 ± 7.2 5PEGda100POSS0 28 ± 0.4 12 ± 4 32.7 ± 6.0 5

and strain. The multiple comparison tests showed that forE there were statistically significant differences between allgroups ( p < 0.01); for σ max there were statistically significantdifferences between PEGda0POSS100 and PEGda25POSS75( p = 0.02), PEGda0POSS100 and PEGda50POSS50 ( p =0.005), PEGda25POSS75 and PEGda100POSS0 ( p = 0.007),and PEGda50POSS50 and PEGda100POSS0 ( p = 0.002); forεmax there were statistically significant differences betweenall groups ( p < 0.05) except between PEGda0POSS100 andPEGda25POSS75, PEGda50POSS50 and PEGda75POSS25,and PEGda75POSS25 and PEGda100POSS0.

3.2. Rheometry

The zero shear viscosity is the value of viscosity when noshear forces are applied and was obtained by extrapolatingthe viscosity versus shear rate curves to zero shear [42]. Theflow experiment indicated that all samples had Newtonian flowbehaviour within the range of shear rates shown in figure 4.This is because the viscosity remains at an approximatelyconstant value regardless of the shear rate, which also suggeststhat the rate of injection would not affect the viscosityof the PEGda/POSS prepolymer mixtures. The η0 of thePEGda/POSS prepolymer mixtures decreased from 1.22 to0.05 Pa.s when the amount of PEGda was increased from 0%to 100%.

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Biomed. Mater. 7 (2012) 035013 J Engstrand et al

Table 4. Handling properties of the PEG–POSS hybrid cementsduring curing, determined according to the ASTM F451–08standard [39].

Formulation Tmax (◦C) ST (s)

PEGda0POSS100 26.9 8528.9 73

PEGda25POSS75 34.7 4635.0 46

PEGda50POSS50 40.6 6333.1 64

PEGda75POSS25 62.3 5660.9 53

PEGda100POSS0 54.2 2357.6 22

3.3. Handling

Table 4 shows the values of Tmax and ST of the PEG–POSS hybrid cements during curing. The general trend(table 4) was that increasing the amount of PEGdaincreased Tmax and decreased ST. The composites withhigh amounts of PEGda had extremely fast polymerizationwith relatively high polymerization temperatures, whereasthe composites containing �50% of PEGda exhibitedpolymerization temperatures that were similar to or lower thanthe body temperature.

3.4. Bioactivity test

It was observed that crystals grew on the surface of all samplescontaining PEGda (figures 5(B)–(E)), but not on the surfaceof PEGda0POSS100 (figure 5(A)). The crystals grown on thehybrids had a plate-like spherical morphology and coveredmost of the surface. Although small differences can be seen infigures 5(B)–(E), the amount of surface coverage was similarfor all three hybrids. The size of the spheres formed on allsamples differed from around 1 μm up to almost 20 μm.Mostly small and randomly distributed crystals were observedon PEGda100POSS0 (figure 5(E)); however, some clusters,with similar morphologies as the larger spheres on the hybrids,could be seen. In contrast, all the control samples storedin water (representative micrograph in figure 5(F)) and thePMMA samples stored in PBS exhibited no crystal growth(figure 5(G)).

The EDS analysis showed a Ca/P ratio of 1.30,1.42 and 1.51 for PEGda25POSS75, PEGda50POSS50 andPEGda75POSS25, respectively. The reference gave a Ca/Pratio of 1.63, indicating that the instrument had an error ofapproximately 2.5%.

The GI-XRD showed an intense peak around 32◦, whichis the most intense peak for HA. Note that some backgroundnoise is present due to the low incidence angle (figure 6).

4. Discussion

New hybrid cements based on PEGda and POSS wereproduced. The hybrid cements were injectable and in situcurable in the presence of BPO and DMPT.

The Young modulus of cancellous bone typically liesbetween 20–750 MPa [43, 44], whereas in standard PMMAcements, it ranges between 1.5 and 3.7 GPa [37, 45].This difference has been associated with additional, adjacentfractures in the spine [5]. The hybrid materials prepared hereinhad E varying from 147 to 819 MPa and compressive strengthsvarying from 50 to 76 MPa. Hence, they have a Young’smodulus that is lower than that of PMMA and CP/CS cementsand more similar to that of cancellous bone. Furthermore,their ultimate compressive strength was higher than that ofcancellous bone and similar to that of PMMA and some of thestronger CP/CS cements, showing their potential use as bonecements. Interestingly, the hybrid PEGda50POSS50, whichhad E = 455 MPa, also had σ max > 70 MPa as is required byASTM F451 standard [39] for prosthesis fixation. The highYoung’s modulus and ultimate compressive strength observedfor the PEGda25POSS75 and PEGda50POSS50, respectively,illustrate the result of hybridizing two materials that possesspoor mechanical behaviour into load-bearing hybrid materials.These properties are the result of strengthening the PEG matrixby introducing POSS cage cross-links. Even though the PEGhas a high strain at break, it has a low mechanical strength,whereas the POSS cages act as both physical nanofillers andeight-armed cross-linking agents for PEG to form chemicallybonded hybrids with higher values of E and σ max than PEGand POSS alone. The high values of E and σ max exhibitedby samples PEGda25POSS75 and PEGda50POSS50 might bethe result of limited chain mobility, due to the chemical cross-linking and due to the presence of bulky POSS, which can act asanchoring groups between the PEG chains. If the concentrationof PEGda increases relative to POSS, there could be a lowerprobability for the PEGda to react with the POSS, resulting ina lower cross-linking density and giving lower values of E andσ max. On the other hand, pure cross-linked POSS resulted in abrittle network that was incapable of absorbing much energybefore fracturing and had low values of E and σ max. Due tothe higher flexibility of the PEG segments compared to themore brittle POSS, the strain at break of the cements increasedwith the amount of PEGda, from 9% for pure POSS to 33%for pure PEGda, which is similar to or higher than the yieldstrain of cancellous bone [46]. The resulting high strain atbreak indicates that these materials can withstand deformationbeyond that of bone.

Injectable biomaterials require an initial zero shearviscosity that is low enough for them to flow through standardneedles [42]. On the other hand, it should be high enoughto avoid leakage towards surrounding tissues [47]. The η0 ofthe PEGda/POSS prepolymer mixtures increased with higherconcentrations of POSS with respect to PEGda. Nevertheless,all the PEGda/POSS prepolymer mixtures were injectable at37 ◦C since they had viscosities between 0.05 and 1.22 Pa.s,which were much lower than those of other injectablebiomaterials found in the literature (e.g., oligo(trimethylenecarbonate) 2400 Da with a viscosity of η0 = 123 Pa.sis injectable through an 18 1

2 G 1.5′′ needle [42]). The lowη0 values of the PEG–POSS hybrids might represent a risk forleakage and must be further investigated, although it shouldbe noted that these values were measured before any type of

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Biomed. Mater. 7 (2012) 035013 J Engstrand et al

(A) (B)

(C) (D)

(E)

(G)

(F)

Figure 5. Micrographs showing representative surfaces after 28 days in PBS of (A) PEGda0POSS100, (B) PEGda25POSS75,(C) PEGda50POSS50, (D) PEGda75POSS25 and (E) PEGda100POSS0. (F) Representative micrograph of a sample stored in water (pictureshows sample PEGda75POSS25). (G) Representative micrograph of a PMMA sample. The cracks observed in some figures are probablydue to microcracks that propagated during vacuum drying.

Figure 6. GI-XRD with an incidence angle of 2◦ for sample PEGda25POSS75 (PDF#01–086-0740).

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Biomed. Mater. 7 (2012) 035013 J Engstrand et al

setting reaction had been initiated; waiting times before aninjection are common in the clinic.

Another issue associated with the handling of commercialPMMA cements is the maximum polymerization temperature,which lies between 40 and 100 ◦C and may harm thesurrounding tissues [48]. The cements investigated in thisstudy exhibited Tmax < 62 ◦C and for the formulationscontaining 50% PEGda or less, the Tmax did not exceed37 ◦C. Hence, the PEG–POSS hybrids possess lower or similarmaximum polymerization temperatures to commercial PMMAbone cements. In addition, an advantage of the potential use ofthe current PEG–POSS hybrids as bone cements would be thepossibility of avoiding the mixing step during the operation.This is because the material is based on two liquid phases thatcan be combined in a dual syringe and mixed simultaneouslyduring the injection. On the other hand, the initial viscositywas low and the setting times were not longer than 2 min,which is a relatively short time for this type of application.The short ST is not necessarily a disadvantage given that thedual syringe eliminates the need for a working time. However,if the injection is delayed, the cement might set in the mixingnozzle, preventing part of it from being injected. Therefore theST needs to be further optimized.

A material that exhibits HA growth in PBS is more likelyto bond to bone and other tissues than a material that does notexhibit HA growth [49]. PMMA bone cements are bioinertmaterials that do not exhibit HA growth on their surface, whichwas confirmed in this study, and they do not chemically bond tobone in vivo [50]. In load-bearing applications, the substitutematerial should transfer the load. If the mechanical propertiesof the substitute material and the tissue to be augmentedsignificantly differ from each other, the interface plays animportant role, and a good bonding could be the key forreliable mechanical performance. The crystals that grew onPEGda25POSS75, PEGda50POSS50 and PEGda75POSS25all had a flake-like spherical morphology indicating that a CPphase was formed on their surfaces [51, 52]. A precipitatecould also be seen on PEGda100POSS0 indicating that the CPphase was also formed here, whereas no growth was observedon PEGda0POSS100 or on the PMMA samples. Furthermore,the crystal layers were much more dense and covered a higherpercentage of the surface of the hybrid samples than thePEGda100POSS0, suggesting that the hybridization improvesthe bioactivity of PEGda and POSS. The EDS results showedan increasing Ca/P ratio with increasing PEGda content.Furthermore, all samples had a Ca/P ratio lower than theCa/P ratio for HA (1.67). However, the GI-XRD analysis ofPEGda25POSS75 showed a clear peak at around 32◦, wherethe highest intensity HA peak is situated, indicating that theprecipitate on the surface most likely is calcium-deficient HA.Previously, it has been shown that silanol (Si–OH) groupscan act as nucleation sites for HA [49]. Each POSS moleculecontains eight silicon atoms that are coordinated by threeoxygen atoms and one methacrylic functional group. Also,the Si–O–Si bridges constitute the edges of the POSS cagestructure (figure 1). Ahola et al [53] showed that the Si–O–Sibridges in silica xerogels synthesized via the sol–gel methodcan undergo hydrolysis to form free Si–OH groups in SBF. As

mentioned earlier, the POSS cages are synthesized in a similarmanner, which suggests that hydrolysis of Si–O–Si bridges canoccur in PBS or in vivo and explains the presence of HA on thesurface of the hybrid samples. This effect was not observed forthe PEGda0POSS100 and is probably due to the hydrophobicnature of POSS [54], which could prevent the water frompenetrating the network and hydrolyzing the Si–O bond. Tosummarize, the presence of PEG gives a hydrophilic featureto the PEG–POSS network, which allows water to approachthe surface and to penetrate it in order to hydrolyze the Si–Obonds into Si–OH groups that serve as HA nucleating sites.It has also been shown that PEG can form a complex withdivalent metal cations, like calcium [55, 56]. This complexcan induce the formation of CP on the surface of PEG, whichexplains why some precipitates can be seen on the PEGdahomopolymer sample.

Although the hybrid material studied seems promising forbone augmentation, there are still some unknown factors thatcan influence its properties as well as its behaviour in vivo.Firstly, a high radiopacity is needed for some applicationssuch as vertebroplasty, and radiopacifying agents would needto be included. The addition of radiopacifiers would probablyaffect the mechanical and biological properties as well as theST of the materials. Secondly, the activator/initiator systemcomponents could be toxic if high concentrations are releasedin vivo during the reaction. However, after the completion ofthe reaction these molecules would be inactive. Thirdly, thepresence of unreacted PEGda could pose a problem due to itshydrophilic character, which would facilitate its release intothe bloodstream before the materials have completely cross-linked. Fourthly, the ST has to be optimized, for example, bylowering the concentration of the radical initiator, the ST couldbe prolonged. Another way of altering the ST could be throughthe addition of inert particles or using other molecular weightsof PEGda. Finally, the in vitro biocompatibility needs to beinvestigated. Materials made from the combination of PEGdaand POSS have not yet been tested for similar applications;however, in vitro tests from POSS-containing materials haveshown good results [57, 58]. This together with the fact thatPEG is approved by the FDA for use in biomedical applicationsgives reason to believe that these materials could achieve anappropriate response in vivo. Cell reactions as well as all ofthe above-mentioned parameters need to be further evaluatedbefore concluding if this hybrid material can substitute alreadyavailable materials.

5. Conclusions

Injectable PEG–POSS hybrid cements were successfullyproduced and characterized in terms of their mechanical,rheological, handling and in vitro bioactivity properties. Theresults showed that the hybrids had improved properties withrespect to PEG and POSS alone.

• Young’s moduli of the hybrids were lower than those ofPMMA and CP/CS cements and more similar to those ofcancellous bone.

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Biomed. Mater. 7 (2012) 035013 J Engstrand et al

• The maximum compressive strength of the hybrids washigher than that of cancellous bone and most commercialCP/CS cements.

• The strain at break of the hybrids varied from about 9%to 33% and increased with the amount of PEG.

• The PEGda50POSS50 samples had a maximumcompressive strength higher than that required forprosthesis fixation while showing an adequate Young’smodulus.

• The PEGda/POSS prepolymer mixtures were injectableat 37 ◦C with zero shear viscosities below 1.22 Pa.s.

• The hybrids exhibited polymerization temperaturesbetween 28 and 62 ◦C and set in approximately half aminute to one and a half minutes.

• The hybrids exhibited a dense calcium-deficient HAlayer after 28 days in PBS, unlike PMMA and thehomopolymers alone, indicating that the hybrids havebioactive surfaces.

In conclusion, this study showed that the PEG–POSS-basedhybrid materials are a promising alternative to both the PMMAand the CP/CS cements for bone augmentation. These newhybrids could offer a possibility to reduce the risk of bonefractures and implant loosening due to the poor mechanicalmismatch between the bone cement and the augmented tissue.Nevertheless, the materials need to be further optimized andevaluated in terms of their radiopacity, handling properties andin vitro biological response before clinical evaluations.

Acknowledgments

The authors would like to acknowledge funding fromthe European Union through the SPINEGO project (ERG268134), VINNOVA through a VINNMER project (2010-02073), the Carl Trygger Foundation and Swedish ResearchCouncil.

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