elastomeric high-mineral content hydrogel-hydroxyapatite composites for orthopedic applications

Elastomeric high-mineral content hydrogel-hydroxyapatitecomposites for orthopedic applications

Jie Song,1,2 Jianwen Xu,1,2 Tera Filion,1,2 Eduardo Saiz,3 Antoni P. Tomsia,3 Jane B. Lian,1,2

Gary S. Stein,2 David C. Ayers,1 Carolyn R. Bertozzi41Department of Orthopedics, University of Massachusetts Medical School, Worcester, Massachusetts 016552Department of Cell Biology, University of Massachusetts Medical School, Worcester, Massachusetts 016553Materials Sciences Division, Lawrence Berkeley National Laboratory, Berkeley, California 947204Departments of Chemistry and Molecular and Cell Biology, University of California; Howard Hughes MedicalInstitute; The Molecular Foundry, Lawrence Berkeley National Laboratory, Berkeley, California 94720

Received 4 April 2008; accepted 8 April 2008Published online 10 June 2008 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.a.32110

Abstract: The design of synthetic bone grafts that mimicthe structure and composition of bone and possess good sur-gical handling characteristics remains a major challenge. Wereport the development of poly(2-hydroxyethyl methacry-late) (pHEMA)-hydroxyapatite (HA) composites termed‘‘FlexBone’’ that possess osteoconductive mineral contentapproximating that of human bone yet exhibit elastomericproperties enabling the press-fitting into a defect site. Theapproach involves crosslinking pHEMA hydrogel in thepresence of HA using viscous ethylene glycol as a solvent.The composites exhibit excellent structural integrationbetween the apatite mineral component and the hydroxy-lated hydrogel matrix. The stiffness of the composite and theability to withstand compressive stress correlate withthe microstructure and content of the mineral component.The incorporation of porous aggregates of HA nanocrystalsrather than compact micrometer-sized calcined HA effec-tively improved the resistance of the composite to crack

propagation under compression. Freeze-dried FlexBone con-taining 50 wt % porous HA nanocrystals could withstandhundreds-of-megapascals compressive stress and >80%compressive strain without exhibiting brittle fractures. Uponequilibration with water, FlexBone retained good structuralintegration and withstood repetitive moderate (megapascals)compressive stress at body temperature. When subcutane-ously implanted in rats, FlexBone supported osteoblastic dif-ferentiation of the bone marrow stromal cells pre-seeded onFlexBone. Taken together, the combination of high osteocon-ductive mineral content, excellent organic-inorganic struc-tural integration, elasticity, and the ability to support osteo-blastic differentiation in vivo makes FlexBone a promisingcandidate for orthopedic applications. � 2008 Wiley Periodi-cals, Inc. J Biomed Mater Res 89A: 1098–1107, 2009

Key words: hydrogel; hydroxyapatite; elastomeric compo-sites; osteoblastic differentiation; bone marrow stromal cells

INTRODUCTION

It is estimated that one seventh of the US populationsuffers from musculoskeletal impairment, costing theUS economy >$215 billion a year.1–4 Well over 600,000orthopedic surgical procedures performed annually inthe US would require bone grafts.5 With a growing

and aging population, this number will continue toincrease. Three types of bone grafts, autogenic, allo-genic, and synthetic, are currently used clinically.6

Although considered as a gold standard, the limita-tions associated with the autografting procedureinclude donor site morbidity, the frequent needs for asecond operation and an inadequate volume of trans-plant material.7,8 Allografts, on the other hand, sufferfrom significant failure rates, poor mechanical stability,and the inherent risks for disease transmission andimmunological rejections.7,8 Synthetic bone grafts havethe potential to become the primary materials of choicein the reconstructive repair of skeletal defects if theyare properly engineered. Unfortunately, current syn-thetic bone substitutes (mostly gel foams or ceramics)have not yet achieved the level of sophistication struc-turally, mechanically, and biochemically to replaceautografts or allografts. Consequently, their clinicaluse (<10%) has lagged far behind autografts (�50%)and allografts (>40%).6

Additional Supporting Information may be found in theonline version of this article.Correspondence to: J. Song; e-mail: [email protected] grant sponsor: U.S. Department of Energy;

contract grant number: DE-AC03-76SF00098Contract grant sponsor: National Institutes of Health;

contract grant numbers: 5R01DE015633, 1R01AR055615Contract grant sponsors: Worcester Foundation for Bio-

medical Research, University of Massachusetts MedicalSchool (Clinical and Translational Science Pilot Project Pro-gram)

� 2008 Wiley Periodicals, Inc.

To fundamentally improve the quality and perform-ance of synthetic bone grafts, new generations of or-ganic/inorganic composites that mimic the structureand composition of bone,9 and possess useful surgicalhandling characteristics (e.g. straightforward and sta-ble fixation) are required. From a materials point ofview, bone is a composite of collagen, a gel-like proteintemplate that resists stretching forces, and the inorganicapatite crystals that resists compression. The bendingand compression strength of human bone is known topositively correlate to bone mineral content.10 Thequantity and quality of the deposited mineral (crystalsize, maturity and structural integration with the or-ganic matrices) influences the mechanical properties ofbone.11 Coupled with its osteoconductivity, hydroxy-apatite (HA), the major inorganic component of bone,has long been recognized as an important design ele-ment for tissue-engineered bone substitutes.12 Towardthis end, methods for incorporating a high percentageof HA with synthetic hydrogels with good structuralintegration are highly desired.

Poly(2-hydroxyethyl methacrylate), or pHEMA, andits functionalized derivatives have been used in a widerange of biomedical applications. With physical prop-erties similar to natural gel-like extracellular matrices,these hydrogel polymers have been explored as oph-thalmic devices,13 soft tissue engineering scaffolds,14

carriers for drug or growth factor delivery,15,16 dentalcements, and medical sealants.17,18 We previouslydeveloped a urea-mediated mineralization method thatcould lead to high-affinity integration of HA on the sur-face of pHEMA hydrogels.19–21 However, integrationof HA with pHEMA at a high mineral-to-gel ratiothroughout the 3D scaffold has not yet been achieved.Here we report the synthesis and characterization ofwell-integrated pHEMA-HA composites with mineral-to-organic matrix ratios approximating those of dehy-drated human bone,22,23 yet exhibiting elastomericproperties. Termed as ‘‘FlexBone,’’ these composites areengineered to withstand moderate (megapascals) com-pressive stresses and exhibit elastomeric propertiesunder physiological conditions to facilitate easy andstable surgical insertion to a site of skeletal defect.Compressive behavior of FlexBone as a function of itsmineral content and microstructures is investigated,and the ability for FlexBone to support the osteogenicdifferentiation of bone marrow stromal cells (BMSC)in vivo is also examined.

MATERIALS AND METHODS

Materials

The radical inhibitors in the commercial HEMA andethylene glycol dimethacrylate (EGDMA) from Aldrich

(Milwaukee, WI) were removed via distillation underreduced pressure and by passing through a 4 A molecularsieve column prior to use, respectively. Polycrystalline com-mercial HA powders (designated as ComHA) were pur-chased from Alfa Aesar (Ward Hill, MA) and used asreceived. The calcined HA powders (designated as CalHA)were obtained by treating ComHA at 11008C for 1 h. Priorto use, the CalHA powders were ground in a planetaryagate mill for 2 h and then passed through a 38 lm sieve toremove larger agglomerates. The microstructures and sizedistributions of these HA particles are shown in Figure 1.Cell culture media and supplements were purchased fromInvitrogen (Carlsbad, CA) and the fetal bovine serum waspurchased from HyClone (Logan, UT). All reagents for his-tochemistry were purchased from Sigma (St Louis, MO).

Preparation of FlexBone composites

The HA content of the FlexBone is defined as the weightpercentage of the HA incorporated over the total weight ofthe HA, monomer HEMA, and crosslinker EGDMA usedin any given preparation. In a typical procedure, freshlydistilled HEMA was mixed with EGDMA along with ethy-lene glycol (EG), water and aqueous radical initiators am-monium persulfate (I-1, 480 mg/mL) and sodium metasul-fite (I-2, 180 mg/mL) at a volume ratio of HEMA:EGD-MA:EG:I-1:I-2/100:2:35:20:5:5 (formulation 1). ComHA orCalHA powder was then added to the hydrogel mixture,thoroughly mixed by using a ceramic ball to break up thelarge agglomerates, and allowed to polymerize in a dispos-able syringe barrel or rigid PMMA tubing of a 7.0-mm or4.7-mm inner diameter to afford composites with HA con-tents varying from 37 to 50%. The resulting elastic materialwas either used as it was (as-prepared), thoroughlyexchanged with a large volume of water (fully hydrated),or freeze-dried. By altering the amount of EG and waterrelative to the HA, 70% of HA, a mineral content approxi-mating that of human bone,22,23 can be incorporated. Forinstance, a volume ratio of HEMA:EGDMA:EG:I-1:I-2/100:2:60:40:5:5 (formulation 2) was used to prepare compo-sites containing 70% CalHA with consistent properties. (Inthis article, however, only properties of composites con-taining up to 50% HA are discussed.) The resulting com-posites are denoted as ComHA-N-# or CalHA-N-#, whereN stands for the type of hydrogel formulation and #denotes the weight percentage of HA content. For instance,ComHA-1-50 represents FlexBone composite containing50% commercial HA that is formed using crosslinking for-mulation 1. Unmineralized pHEMA control was preparedusing formulation 1 in the absence of HA particles.

Microstructural characterization

The microstructures of the composites were character-ized using environmental scanning electron microscopy(ESEM) on a Hitachi S-4300SEN microscope (Hitachi,Japan). The chamber pressure was kept �35 Pa to avoidcomplete sample dehydration and surface charging duringthe observation. The chemical composition was analyzed

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using energy dispersive spectroscopy (EDS) (Noran systemSIX, Thermoelectron, USA) attached to the ESEM.

Mechanical testing

To assess the compressive behavior of FlexBone in as-prepared, fully hydrated and freeze-dried states as a func-tion of the mineral microstructure and content, unconfinedcompression tests were performed on two different instru-ments, a Q800 Dynamic Mechanical Analyzer (DMA) anda high capacity MTS, to accommodate the needs for highsensitivities and high loading capacities, respectively. Allsamples were tested in accordance with ASTM D695 withthe exception of sample size and slenderness ratio (recom-mended ratio: 1:4 diameter-to-length) due to sample heightlimitation of the DMA instrument (�5 mm) and the con-cern over the significant error that preparing and testingextremely small-diameter samples may introduce. Shortercylinders were also used for the high capacity MTS due tothe concern over sample buckling under extremely highcompressive strains. All stress–strain curves presented arebased on engineering stress and engineering strain re-corded on each instruments, assuming a fixed cross-sectionof the material defined at the start of the test.

At least five specimens were tested for each sample. Foras-prepared and water-equilibrated samples, cylindricalspecimens with a diameter of 4.7 mm were transversely

cut into 5.0-mm long cylinders using a custom-machinedparallel cutter with adjustable spacing. Any visible rough-ness of the top and bottom surfaces of each specimen wasreduced by sandpaper. An L-square was used to makesure that these surfaces were parallel prior to testing, andthe final dimensions of each specimen were measured by adigital caliper. For freeze-dried samples, cylindrical speci-mens with the dimension of 7 mm 3 6 mm (diameter 3height) were used.

The compressive behavior of as-prepared and water-equilibrated FlexBone composites along with pHEMA con-trol, particularly their elasticity, was evaluated on a Q800DMA (TA Instruments) equipped with a submersion com-pression fixture. The instrument has an 18-N load cell, aforce resolution of 10 lN and a displacement resolution of1.0 nm. The as-prepared samples were compressed in aforce-controlled mode in ambient air, ramping from 0.01 to18.0 N at a rate of 3.0 N/min then back to 0.01 N at thesame rate. The samples fully equilibrated with water werecompressed in water at 37.58C, ramping from 0.01 to10.0 N at a rate of 3.0 N/min then back to 0.01 N at thesame rate. To evaluate the reversibility of the compressivebehavior, the controlled force cycle was repeated 10–40times consecutively for each specimen unless the materialfailed (major cracks developed) during the force ramping,at which point the test would be terminated. In all cases,we observed little further shift of loading–unloadingcurves beyond 10 cycles. For clarity, in figures that com-

Figure 1. Microstructures and size distribution of ComHA versus CalHA powders. (A) SEM micrograph of ComHApowders showing porous aggregates of polycrystalline HA. (B) Higher resolution SEM image of the circled area in (A)showing HA crystallites �100 nm in size. (C) Grinded CalHA powders. (D) Particle size distribution of the CalHA asdetermined by sedimentation measurements for particles with diameters below 10 lm. Both SEM micrograph and the sed-imentation measurement plot suggested a bimodal size distribution of CalHA powders with most particles sized 5 lm orbelow and the larger grains over 10 lm in size.

1100 SONG ET AL.


pare the compressive behaviors among different samples[Fig. 2(A,C)], the first 10 cycles of the stress–strain curvesfrom one representative specimen of each sample wereplotted. To demonstrate the reliability of the measurementsand the reproducibility of the compressive behavior ofeach sample, five specimens were examined for each sam-ple. Because of the observed high reproducibility, how-ever, only the first 10 cycles of the consecutive loading–unloading curves from three specimens for each sampleare presented in overlaid fashion as Supplemental Infor-mation (Supplemental Fig. 1 and Supplemental Fig. 2).

The compressive behavior of freeze-dried FlexBone com-posites, particularly their ability to withstand high com-pressive loads without exhibiting brittle fractures, wasevaluated in ambient air on a high-capacity MTS servo-hydraulic mechanical testing machine (MTS Systems Cor-poration) equipped with a 100 kN load cell and stiff,

nondeforming platens. The samples were loaded underdisplacement control at a rate of �0.015 mm/s up to 80–90% compressive strain, while the corresponding loadsand displacements were continuously monitored using thebuilt-in load cell and linear variable displacement trans-ducer (LVDT).

Isolation and in vitro expansion of rat BMSC

All animal procedures were conducted in accordance withthe principles and procedures approved by the University ofMassachusetts Medical School Animal Care and Use Com-mittee. BMSC were isolated from long bones of 4-week oldmale Charles River SD strain rats.24 Briefly, marrow wasflushed from femur with a syringe containing MEM. Afterlysing red blood cells with sterile water, the marrow cells

Figure 2. Compressive behavior and composition of as-prepared versus fully hydrated FlexBone. (A) Compressivebehavior of as-prepared FlexBone and pHEMA control as a function of mineral microstructure and content. Ten consecu-tive load-controlled loading-unloading cycles (3.0 N/min, 0.01 N to 18.0 N to 0.01 N) were applied to each specimen inambient air using a Q800 DMA equipped with a compression fixture. (B) EDS of the cross-sections of FlexBone showingthe removal of residue S-containing radical initiators upon equilibrating the as-prepared sample with water. (C) Compres-sive behavior of fully hydrated FlexBone and pHEMA control at body temperature as a function of mineral microstructureand content. Ten consecutive load-controlled loading-unloading cycles (3.0 N/min, 0.01 N to 10.0 N to 0.01 N) wereapplied to each specimen in water using a Q800 DMA equipped with a submersion compression fixture. The hydratedFlexBone containing CalHA started to fail approaching >30% compressive strain during the first force ramping (denotedby *), thus did not continue with additional loading cycles.



were centrifuged and resuspended in minimum essential me-dium (MEM) supplemented with 20% FBS, 0.2% penicillin-streptomycin and 1% l-glutamine, and passed through a ster-ile metal filter. Cells were expanded on tissue culture plates(10 million cells per 100-mm plate initial seeding density)with media changes on day 4 and every other day thereafterbefore they were lifted off for plating on FlexBone.

Subcutaneous implantation of FlexBone compositein rats with and without preseeded BMSC

Thin half discs (7 mm in diameter, 1 mm in thickness)of FlexBone containing 40% ComHA (ComHA-1-40) weresterilized in 70% ethanol, re-equilibrated with sterile waterbefore being seeded with BMSC and used for subcutane-ous implantation in rats. Fifty microliters of BMSC suspen-sion (in culture media described above) was loaded on thesurface of thin disks of FlexBone to reach 5000-cells/cm2

or 20,000-cells/cm2 seeding density. The cell-seeded Flex-Bone were incubated at 378C in humidified environmentwith 5% CO2 without additional media for 6 h to allowcell attachment to the FlexBone substrate. Additionalmedia were then added and the cells were cultured on thesubstrates for 2 days before being used for implantation.Four sets of samples were used for each cell seeding treat-ment. Thin discs of FlexBone without preseeded BMSCwere also used for implantation as controls.

Rats were anesthetized by intraperitoneal (IP) injectionof ketamine/xylazine (50 mg/5 mg per kg). They wereshaved and swabbed with betadine before two 1/4 inbilateral skin incisions were made over the rib cage forinsertion of the FlexBone discs with and without pre-seeded BMSC. The skin was closed with surgical staplesand buprenorphine (0.02 mg/kg) was given subcutane-ously. The rats were sacrificed by CO2 inhalation and cer-vical dislocation at day 14 and day 28 for the retrieval ofFlexBone. After removing the fibrous tissue encapsulation,the retrieved FlexBone was fixed in 4% paraformaldehyde(0.1M phosphate buffer, pH 7.4) for 5 h at 48C before beinganalyzed by SEM, XRD, and histology.

X-ray powder diffraction

The crystalline phases of the mineral in the FlexBonecomposites before and after subcutaneous implantation inrats were evaluated by XRD with a Siemens D500 instru-ment using Cu Ka radiation. Phases were identified bymatching the diffraction peaks to the JCPDS files.

Histochemical staining of explanted FlexBone foralkaline phosphatase activity

The 4% paraformaldehyde-fixed FlexBone explants wereequilibrated in cacodylic buffer overnight, then in 30% su-crose solution (pH 7.3) for 2 days before being frozen-sec-tioned on a Bright Cryostat (Model OTF; Bright InstrumentLtd., Huntigdon, UK). Frozen-sectioning was repeateduntil reaching the depth of 100–200 lm away from the sur-face where the BMSC were initially seeded. The 12-lm fro-

zen sections were held on adhesive slides using frozen sec-tioning tape. Histological staining for ALP activity, amarker of osteogenic differentiation, was performed asdescribed in literature.25 Briefly, the frozen sections ofFlexBone explants were incubated with 1.5 mM naphthol-As-Mx phosphate disodium salt, 0.1% Fast Red and 2.7%DMF (v/v) in 0.1M Tris acid maleate buffer (pH 8.4) for 30min, and the positive stains (in red) were detected by opti-cal microscopy.

RESULTS

Preparation and compressive behavior ofas-prepared and fully hydrated FlexBone

FlexBone composites with varying mineral con-tents (37–70%) were prepared by crosslinkingHEMA with 2% EGDMA in the presence of eitherporous aggregates of HA nanocrystals (ComHA) orcompact micrometer-sized calcined HA (CalHA) par-ticles (Fig. 1) using ethylene glycol as a solvent. Re-petitive unconfined compressive tests performed onthe as-prepared FlexBone with varying mineral con-tents revealed mineral content-dependent and min-eral microstructure-dependent elastomeric compres-sive behavior. As indicated by the slopes of the com-pressive stress–strain curves shown in [Fig. 2(A)],FlexBone composites are stiffer (steeper slope) thanthe un-mineralized pHEMA hydrogel prepared withthe same degree of crosslinking. In addition, Flex-Bone composites containing higher mineral contentsare stiffer than those containing less HA particles ofthe same type, showing a positive correlationbetween the stiffness and the mineral content of thepolymer-mineral composite. Notable difference incompressive behavior as a function of the type ofHA components incorporated was also observed,with FlexBone containing ComHA much stiffer thanthose containing the same percentages of CalHA.Finally, good overlaps of stress–strain curves wereobserved when 10 consecutive compressive loading/unloading cycles up to �1 MPa (the maximal appli-cable loads of the DMA instrument with the chosensample size) were applied to all as-prepared compo-sites. Such good recovery under compressive strainsup to 40% depending on the composition is expectedto facilitate the press-fitting of these composites intoa defect area. As a reference, the peak contactstresses in natural human joints during light to mod-erate activity typically range from 0.5–6 MPa bymost in vitro measurements,26–29 and up to 18 MPaby some in vivo measurements.30,31 Overall, our datasuggest that as-prepared FlexBone exhibit excellentshape recovery under repetitive, physiologically rele-vant compressive stress despite their high (37–50%)mineral contents.

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The as-prepared composites can undergo solventexchange with water to give fully hydrated Flex-Bone. The residue sulfur-containing radical initiatorstrapped in the as-prepared composites could beremoved during the wash with water as indicatedby the disappearance of the S signal detected fromthe energy dispersive spectroscopy (EDS) performedon the cross-section of the composite upon equilibra-tion with water [Fig. 2(B)]. The compressive behaviorof fully hydrated FlexBone was examined at bodytemperature in water using a DMA equipped with asubmersion compression fixture. As shown in [Fig.2(C)], mineral content-dependent and mineral micro-structure-dependent compressive behavior similar tothose exhibited by as-prepared FlexBone wasobserved with fully hydrated FlexBone. A noticeabledifference, however, is that fully hydrated FlexBonecontaining CalHA failed (with major cracks formed)when >30% of compressive strain was applied. Incontrast, FlexBone containing 37% and 50% ComHAcould withstand repetitive megapascal compressivestress with excellent shape recovery in water. Thedifference observed with the hydrated compositescontaining same percentages of ComHA versusCalHA powders underscores the importance of themicrostructures of the mineral component, and likelytheir differential behavior in interfacing with thepolymer matrix, in determining the bulk mechanicalproperties of the polymer-mineral composites.

Compressive behavior and microstructures offreeze-dried FlexBone

To better understand how the microstructure ofthe mineral component and the organic–inorganicinterface dictates the macroscopic compressivebehavior of FlexBone, we examined the microstruc-tural response of freeze-dried composites containingComHA versus CalHA under very high compressivestress and strains (>80%). Freeze-drying the fullyhydrated FlexBone composites did not lead to thedelamination of the evenly distributed mineral com-ponents, either ComHA or CalHA, from the polymermatrix that they were embedded in [RepresentativeSEM: Fig. 3(B,D)]. To apply high compressive strainsto the freeze-dried composites, a high capacity MTSwith 100 kN load cell was used to perform uncon-fined compression test. As expected, the freeze-driedcomposites were stiffer than their hydrated counter-parts. Importantly, all tested freeze-dried FlexBonecomposites were able to withstand compressivestress in the order of hundreds of megapascals andcompressive strains of >80% without exhibiting brit-tle fractures despite their high mineral contents [Fig.3(A)]. In contrast, PMMA-based bone/dentalcements or poly(lactic acid)-HA composites reported

in literature typically exhibited brittle fracture at 50–150 MPa compressive loading.32–34

A closer examination of the stress–strain curvesrevealed that freeze-dried composites containingComHA tend to be stiffer than those containing samepercentages of CalHA, as representatively shown in[Fig. 3(A)]. This is consistent with the trend observedwith as-prepared and hydrated FlexBone under lowercompressive stresses. SEM analysis of freeze-driedCalHA-1-50 after compression tests resulting in >80%strains revealed the formation of cracks within thehydrogel phase whereas no distortion or fracture ofthe micrometer-sized compact CalHA particles wasobserved [Fig. 3(B) vs. 3(C)]. These cracks could affectthe slope of the stress–strain curve. In contrast, thehydrogel-infiltrated aggregates of HA nanocrystals infreeze-dried ComHA-1-50 were flattened upon com-pression into plywood-like structures with no disrup-tion of the continuity of the hydrogel matrix [Fig.3(D) vs. 3(E)]. The rearrangement of the nanometer-sized HA crystallites can provide a mechanism forenergy dissipation within the composite under highcompressive stresses.

In vivo osteogenic differentiation of BMSCsupported by FlexBone

To test the cytocompatibility and the in vivo resorp-tion of FlexBone, we seeded hydrated compositesComHA-1-40 with BMSC isolated from rat femur,and implanted them subcutaneously (SC) in 4-weekold male Charles River SD strain rats. The compositeswere retrieved at 14 and 28 days, with a degree of fi-brous tissue encapsulation observed in all cases. Afterremoving the fibrous tissue, the morphology andmineral phase of the retrieved implant were exam-ined by SEM and X-ray powder diffraction (XRD).Little macroscopic change in shape or size of theretrieved FlexBone was observed, reflecting the non-degradable nature of the hydrogel scaffold thatdefines the overall shape of the composite. However,surface roughening was observed with both 14- and28-day explants regardless whether they were pre-seeded with BMSC prior to implantation [Fig.4(A,B)]. This is likely a combined outcome of slowdissolution of the mineral component and the extrac-ellular matrix deposition from cells either pre-seededon or newly attracted to the substrate in vivo. XRDanalyses performed with the explanted composite[Fig. 4(C)] revealed a diffraction pattern matchingwith that of the ComHA powder, suggesting that themajor mineral phase remained unchanged 4 weeksafter the SC implantation.

To determine whether the composite can supportthe osteogenic differentiation of BMSC in vivo, theexplanted composites with preseeded BMSC were



stained histochemically for alkaline phosphatase(ALP) activity, a marker for osteogenic differentia-tion.35 To avoid the harsh paraffin embedding condi-

tions that may compromise ALP enzymatic activity,36

frozen sectioning was performed on the explantsprior to ALP staining. As shown in [Fig. 4(D)], ALP

Figure 3. Microstructures and compressive behavior of freeze-dried FlexBone containing ComHA versus CalHA. (A)Stress–strain curves showing freeze-dried FlexBone containing 50% ComHA is stiffer than the one containing 50% CalHA.Unconfined displacement-controlled (�0.015 mm/s) compression test was performed on a high capacity MTS with a 100-kN load cell. (B) and (C): SEM of the cross-section of freeze-dried CalHA-1-50 before and after being compressed. (D) and(E): SEM of the cross-section of freeze-dried ComHA-1-50 before and after being compressed. The arrows in (C) and (E)indicate the direction of compression.

Figure 4. In vivo resorption and osteogenic differentiation of bone marrow cells supported by FlexBone ComHA-1-40.(A) SEM micrograph of a composite (pre-seeded with 20,000-cells/cm2 BMSC) retrieved 28 days after SC implantation inrat; (B) SEM micrograph of a composite (without pre-seeded BMSC) retrieved 14 days after SC implantation in rat; (C)XRD of the explanted sample shown in (A), with diffraction patterns matching with that of the commercial HA powder;(D) ALP staining (red) of a 12-lm frozen section of an explanted composite (pre-seeded with 5,000-cells/cm2 BMSC) onday 14. Magnification: 4003. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com.]

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activity (indicated by red stains) was detected 14days postimplantation on the periphery of theComHA-1-40 preseeded with 5000-cells/cm2 BMSC.More extensive ALP activity was also detected 28days after the implantation on FlexBone preseededwith 20,000-cells/ cm2 BMSC. These data suggest thatFlexBone was able to support the attachment andin vivo osteoblastic differentiation of osteoblast pre-cursor cells.

DISCUSSION

We reported the preparation of a class of elasto-meric pHEMA-HA composites, FlexBone, consistingof high percentages of osteoconductive HA using astraightforward protocol. The high viscosity of ethyl-ene glycol, the solvent used during the fabrication ofFlexBone, facilitated the easy dispersion of 50 wt %HA particles within the hydrogel formulation,thereby preventing the HA particles from settling bygravity during solidification. The intrinsic affinity ofthe hydroxyl side chains of the crosslinked pHEMAmatrix to the surface of calcium apatite crystals ledto the formation of strong interfaces between the or-ganic and inorganic components. The good surfacebonding of HA particles to the pHEMA matrix wasmaintained upon freeze-drying and contributed tothe ability of the freeze-dried composites to with-stand hundreds of megapascal compressive stressand >80% compressive strains without exhibitingbrittle fractures.

Side-by-side comparisons of the compressivestress–strain curves obtained with FlexBone compo-sites in as-prepared [Fig. 2(A)], hydrated [Fig. 2(C)]and freeze-dried [Fig. 3(A)] states revealed convinc-ing correlations between the content/microstructuresof the mineral component and the macroscopic com-pressive behavior of the composite. We have shownthat the stiffness of FlexBone positively correlateswith the content of a given microstructure of HA,with the slope of stress–strain curves of ComHA-1-50, for instance, steeper than that of ComHA-1-37regardless of their solvent environment (ethyleneglycol or water). The same trend was also observedwith FlexBone containing CalHA. In natural bone,the bending, compression and tensile moduli ofcompact bone have been shown to exhibit a strongpositive correlation with its mineral content.10,37,38

Our data have also demonstrated significantimpact of the size and microscopic scale aggregation(structure) of HA minerals on the bulk compressivebehavior of FlexBone. Whether in as-prepared, fullyhydrated or freeze-dried state, FlexBone containingporous aggregates of HA nanocrystals (ComHA) arealways significantly stiffer and stronger (‘‘stronger’’

in terms of their resistance to crack formation andpropagation under compression) than those contain-ing the same percentage of compact micrometer-sized CalHA. The process of solvent exchange withwater did not compromise the ability of as-preparedFlexBone containing ComHA to withstand repetitivephysiological compressive stress and moderate(>10%) compressive strains, a feature highly desira-ble for the surgical insertion and use of FlexBone assynthetic bone grafts. In contrast, hydration signifi-cantly weakened the compressive strength of Flex-Bone containing CalHA (e.g. ultimate strength <0.6MPa in water for CalHA-1-37 and CalHA-1-50),making them less suitable for moderate weight-bear-ing applications in vivo. Poor structural integrationof polymer matrices with mineral components arealso known to contribute to rapid and significantdegradation of the mechanical integrity of other syn-thetic high mineral-content composites (e.g. PLA/HA composites) in aqueous environment.39

We hypothesize that the sub-micrometer scaleaggregation of HA nanoparticles in the ComHAacted as ‘‘sponges’’, absorbing the prepolymer hydro-gel formulation and yielding larger surface contactareas between the hydrogel matrix and the ComHAcrystals. The better structural integration of the or-ganic and inorganic components has translated intoa significantly reduced tendency for crack formationand propagation within the resulting compositesunder high compressive stress. SEM studies furtherelucidated that the hydrogel-infiltrated sphericalaggregates of HA nanocrystals flattened into ply-wood-like structures upon compression, providingan important energy-dissipation mechanism for Flex-Bone under compressive stress. The enhancement ofstrength and ductility of ceramics and ceramic nano-composites as a function of decreasing grain sizes isa subject of intense recent investigation anddebate.40,41 No simple extrapolation of earlier find-ings in either ceramic, metallic, or intermetallic sys-tems can predict the behavior of FlexBone since thecombination of the soft hydrogel with the hard apa-tite crystals is quite unique. However, the micro-structure-compressive behavior correlation revealedin our system is reminiscent of that observed withthe analogous composite in nature—bone. It is well-known that the quality of the structural integrationof the hard apatite crystals with the soft collagennetwork on nanoscopic and microscopic levelsdirectly impact the mechanical properties of bone.9

In fact, in aging bone, poorer structural integrationof bone mineral with the collagen matrix is just asimportant as the decreasing mineral content in con-tributing to their weaker and more brittle mechanicalproperties. In the case of FlexBone, the impact ofmineral microstructures on compressive behaviorseems to have out-weighted that of the mineral con-



tent among the samples we examined. For instance,ComHA-1-37 is significantly stiffer than CalHA-1-50and less likely to crack under megapascal-compres-sive stress in water [Fig. 2(A,C)].

Taken together, FlexBone containing ComHAexhibited tunable reversible compressive behavior inphysiologically relevant environment (e.g. in water,at body temperature, and under megapascal com-pressive stress), making them appealing syntheticbone graft candidates. Subcutaneous implantation ofComHA-1-40 preseeded with BMSC in rats showedthat the osteoconductive composite provided a cyto-compatible environment to support the attachment,penetration, and osteogenic differentiation of BMSCin vivo. An ideal synthetic bone graft is designed tofill an area of defect to provide immediate structuralstabilization and to expedite the healing and repairof the skeletal lesion. Ideally, the synthetic grafts canbe eventually remodeled and replaced by newly syn-thesized bone. From this perspective, biodegradabil-ity and osteoinductivity of the synthetic bone graftsare just as important as their osteoconductivity, me-chanical strength, and material handling characteris-tics (e.g. elasticity facilitating surgical insertion).Future improvements include engineering the biode-gradability of the organic matrix, enhancing thein vivo dissolution rate of the osteoconductive min-eral component (e.g. by introducing the more solubleb-tricalcium phosphate, b-TCP,42 to the mineralphase), and locally retaining and releasing osteoin-ductive growth factors and cytokines on and fromthe synthetic scaffold.

CONCLUSIONS

In summary, lightweight FlexBone compositescontaining high percentages of HA were preparedby crosslinking HEMA in the presence of HA usingethylene glycol as a solvent. Despite their high min-eral contents (37–50%), the as-prepared compositesexhibited elastomeric properties and reversible com-pressive behavior under moderate (megapascals)compressive stress. Owing to the excellent structuralintegration between the apatite mineral and thepHEMA network, freeze-dried FlexBone could with-stand hundreds-of-megapascals compressive stressand >80% compressive strain without exhibitingbrittle fractures [Fig. 3(A)]. We further showed thatthe incorporation of porous aggregates of HA nano-crystals, rather than compact micrometer-sized cal-cined HA, into the hydrogel matrix could effectivelyimprove the overall stiffness of FlexBone and its re-sistance to crack formation and propagation undercompression. Upon equilibration with water, thesecomposites retained good structural integration, and

were able to support the attachment and osteoblasticdifferentiation of BMSC in vivo. Combined with theelasticity that facilitates the easy and stable surgicalinsertion of FlexBone into an area of bony defect andenables better accommodation to the micro move-ment of bone, these osteoconductive composites canfind important orthopedic applications.

More broadly, the strong organic/inorganic inter-face achieved with FlexBone demonstrates that non-covalent binding between apatite crystals and ahighly hydroxylated organic matrix can be exploitedin the rational design of bone-like composites. Inaddition, the mineral content-dependent and mineralmicrostructure-dependent compressive behaviorexhibited by FlexBone underlines the importance oftaking into account the combined effect of these pa-rameters in the rational design of functional struc-tural composites.

References

1. Rice DP, Praemer A, Furner S. Musculoskeletal Conditions inthe United States. Chicago, IL: American Academy of Ortho-pedics; 1999.

2. Bone Health and Osteoporosis: A Report of the Surgeon Gen-eral. Rockville, MD: US Department of Health and HumanServices, Public Health Service, Office of the Surgeon Gen-eral, 2004.

3. Favus M. Primer on the Metabolic Bone Diseases and Disor-ders of Mineral Metabolism, 5th ed. Washington, DC: Ameri-can Society for Bone and Mineral Research; 2003.

4. Cummings SR, Melton LJ. Epidemiology and outcomes ofosteoporotic fractures. Lancet 2002;359:1761–1767.

5. http://www.allcountries.org/uscensus/201_organ_transplants_and_grafts.html United States Census Bureau, US. Departmentof Commerce, US Cencus 2000.

6. Bostrom MPG, Seigerman DA. The clinical use of allografts.Demineralized bone matrices, synthetic bone graft substitutesand osteoinductive growth factors: A survey study. HSS J2005;1:9–18.

7. Goldberg VM, Stevenson S. Bone-graft options-Fact andfancy. Orthopedics 1994;17:809.

8. Goldberg VM, Shaffer JW, Field G, Davy DT. Biology of vas-cularized bone-grafts. Orthop Clin North Am 1987;18:197–205.

9. Weiner S, Wagner HD. The material bone: Structure-mechani-cal function relations. Annu Rev Mater Sci 1998;28:271–298.

10. Follet H, Boivin G, Rumelhart C, Meunier PJ. The degree ofmineralization is a determinant of bone strength: A study onhuman calcanei. Bone 2004;34:783–789.

11. Tong W, Glimcher MJ, Katz JL, Kuhn L, Eppell SJ. Size andshape of mineralites in young bovine bone measured byatomic force microscopy. Calcif Tissue Int 2003;72:592–598.

12. El-Ghannam A. Bone reconstruction: From bioceramics to tis-sue engineering. Expert Rev Med Devices 2005;2:87–101.

13. Kidane A, Szabocsik JM, Park K. Accelerated study on lyso-zyme deposition on poly(HEMA) contact lenses. Biomaterials1998;19:2051–2055.

14. Dalton PD, Flynn L, Shoichet MS. Manufacture of poly(2-hydroxyethyl methacrylate-co-methylmethacrylate) hydrogeltubes for use as nerve guidance channels. Biomaterials 2002;23:3843–3851.

1106 SONG ET AL.


15. Lu S, Anseth KS. Photopolymerization of multilaminated poly(HEMA) hydrogels for controlled release. J Control Release1999;57:291–300.

16. Sefton MV, May MH, Lahooti S, Babensee JE. Making micro-encapsulation work: Conformal coating, immobilization gelsand in vivo performance. J Control Release 2000;65:173–186.

17. Yang JM, You JW, Chen HL, Shih CH. Calorimetric character-ization of the formation of acrylic type bone cements.J Biomed Mater Res 1996;33:83–88.

18. Prati C, Mongiorgib R, Valdreb G, Montanaria G. Hydroxy-ethyl-methacrylate dentin bonding agents: Shear bondstrength, marginal microleakage and SEM analysis. ClinMater 1991;8:137–143.

19. Song J, Saiz E, Bertozzi CR. A new approach to mineraliza-tion of biocompatible hydrogel scaffolds: An efficient processtowards 3-dimensional bonelike composites. J Am Chem Soc2003;125:1236–1243.

20. Song J, Saiz E, Bertozzi CR. Preparation of pHEMA-CP com-posites with high interfacial adhesion via template-drivenmineralization. J Eur Ceram Soc 2003;23:2905–2919.

21. Song J, Malathong V, Bertozzi CR. Mineralization of syn-thetic polymer scaffolds: A bottom-up approach for the de-velopment of artificial bone. J Am Chem Soc 2005;127:3366–3372.

22. An YH. Mechanical properties of bone. In: An YH, DraughnRA, editors. Mechanical Testing of Bone and the Bone-Implant Interface. Boca Raton: CRC Press; 2000. p 41–63.

23. Phelps JB, Hubbard GB, Wang X, Agrawal CM. Microstruc-tural heterogeneity and the fracture toughness of bone.J Biomed Mater Res 2000;51:735–741.

24. Miline M, Kang M-I, Quail JM, Baran DT. Thyroid hormoneexcess increases insulin-like growth factor I transcripts inbone marrow cell cultures: Divergent effects on vertebraland femoral cell cultures. Endocrinology 1998;139:2527–2534.

25. Drissi H, Hushka D, Aslam F, Nguyen Q, Buffone E, Koff A,van Wijnen AJ, Lian JB, Stein JL, Stein GS. The cell cycle reg-ulator p27(kip1) contributes to growth and differentiation ofosteoblasts. Cancer Res 1999;59:3705–3711.

26. Ahmed AM, Burke DL. In vitro measurement of static pres-sure distribution in synovial joints-part I: Tibial surface of theknee. J Biomech Eng 1983;105:216–225.

27. Brown TD, Shaw DT. In vitro contact stress distributions inthe natural human hip. J Biomech 1983;16:373–384.

28. Whalen RT, Carter DR. Influence of physical activity on theregulation of bone density. J Biomech 1988;21:825–837.

29. Brand RA. Joint contact stress: A reasonable surrogate for bi-ological processes? Iowa Orthop J 2005;25:82–94.

30. Hodge WA, Fijan RS, Carlson KL, Burgess RG, Harris WH,Mann RW. Contact pressures in the human hip joint meas-ured in vivo. Proc Natl Acad Sci USA 1986;83:2879–2883.

31. HodgeWA, Carlson KL, Fijan RS, Burgess RG, Riley PO, HarrisWH, Mann RW. Contact pressures from an instrumented hipendoprosthesis. J Bone Joint Surg Am 1989;71:1378–1386.

32. Saha S, Pal S. Mechanical properties of bone cement: Areview. J Biomed Mater Res 1984;18:435–462.

33. Shikinami Y, Okuno M. Bioresorbable devices made of forgedcomposites of hydroxyapatite (HA) particles and poly-L-lac-tide (PLLA): Part I. Basic characteristics. Biomaterials 1999;20:859–877.

34. Kuhn K-D. Bone Cements. New York: Springer; 2000.35. Vanhoof VO, Debroe ME. Interpretation and clinical-signifi-

cance of alkaline-phosphatase isoenzyme patterns. Crit RevClin Lab Sci 1994;31:197–293.

36. Farley JR, Hall SL, Herring S, Libanati C, Wergedal JE. Refer-ence-standards for quantification of skeletal alkaline-phos-phatase activity in serum by heat inactivation and lectin pre-cipitation. Clin Chem 1993;39:1878–1884.

37. Currey JD. The effect of porosity and mineral content on theYoung’s modulus of elasticity of compact bone. J Biomech1988;21:131–139.

38. Currey JD. Physical characteristics affecting the tensile failureproperties of compact bone. J Biomech 1990;23:837–844.

39. Russias J, Saiz E, Nalla RK, Gryn K, Ritchie RO, Tomsia AP.Fabrication and mechanical properties of PLA/HA compo-sites: A study of in vitro degradation. Mater Sci Eng C2006;26:1289–1295.

40. Sternitzke M. Review: Structural ceramic nanocomposites.J Eur Ceram Soc 1997;17:1061–1082.

41. Dollar M, Dollar A. On the strength and ductility of nano-crystalline materials. J Mater Proc Tech 2004;157/158:491–495.

42. Kwon S-H, Jun Y-K, Hong S-H, Lee I-S, Kim H-E. Calciumphosphate bioceramics with various porosities and dissolu-tion rates. J Am Ceram Soc 2002;85:3129–3131.



elastomeric high-mineral content hydrogel-hydroxyapatite composites for orthopedic applications

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