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Acta Biomaterialia 4 (2008) 638–645

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Mechanical and biological propertiesof hydroxyapatite/tricalcium phosphate scaffolds coated

with poly(lactic-co-glycolic acid)

Xigeng Miao *, Dawn Meifang Tan, Jian Li, Yin Xiao, Ross Crawford

Institute of Health and Biomedical Innovation and School of Engineering Systems, Queensland University of Technology,

Corner of Blamey Street and Musk Avenue, Kelvin Grove, QLD 4059, Australia

Received 21 July 2007; received in revised form 10 October 2007; accepted 11 October 2007Available online 25 October 2007

Abstract

Regeneration of bone, cartilage and osteochondral tissues by tissue engineering has attracted intense attention due to its potentialadvantages over the traditional replacement of tissues with synthetic implants. Nevertheless, there is still a dearth of ideal or suitablescaffolds based on porous biomaterials, and the present study was undertaken to develop and evaluate a useful porous composite scaf-fold system. Here, hydroxyapatite (HA)/tricalcium phosphate (TCP) scaffolds (average pore size: 500 lm; porosity: 87%) were pre-pared by a polyurethane foam replica method, followed by modification with infiltration and coating of poly(lactic-co-glycolicacid) (PLGA). The thermal shock resistance of the composite scaffolds was evaluated by measuring the compressive strength beforeand after quenching or freezing treatment. The porous structure (in terms of pore size, porosity and pore interconnectivity) of thecomposite scaffolds was examined. The penetration of the bone marrow stromal stem cells into the scaffolds and the attachment ofthe cells onto the scaffolds were also investigated. It was shown that the PLGA incorporation in the HA/TCP scaffolds significantlyincreased the compressive strength up to 660 kPa and the residual compressive strength after the freezing treatment decreased to160 kPa, which was, however, sufficient for the scaffolds to withstand subsequent cell culture procedures and a freeze–drying process.On the other hand, the PLGA coating on the strut surfaces of the scaffolds was rather thin (<5 lm) and apparently porous, maintain-ing the high open porosity of the HA/TCP scaffolds, resulting in desirable migration and attachment of the bone marrow stromal stemcells, although a thicker PLGA coating would have imparted a higher compressive strength of the PLGA-coated porous HA/TCPcomposite scaffolds.� 2007 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Keywords: Hydroxyapatite; Poly(lactic-co-glycolic acid); Porosity; Quenching; Cell attachment

1. Introduction

Bone defects are conventionally treated by replacementwith bone grafts and synthetic bone filling materials. How-ever, the tissue engineering approach, which stresses tissueregeneration rather than tissue replacement, has becomepopular recently. Porous biomaterials (also called scaffolds)used in tissue engineering allow cells to attach, proliferate,

1742-7061/$ - see front matter � 2007 Acta Materialia Inc. Published by Else

doi:10.1016/j.actbio.2007.10.006

* Corresponding author. Tel.: +61 7 3138 6237.E-mail address: x.miao@qut.edu.au (X. Miao).

differentiate and eventually become specific tissue(s). Whilescaffolds are expected to disappear after implantationin vivo, a certain level of mechanical strength is requiredfor the scaffolds to withstand a certain level of physiologicalloading. The open porosity of the porous scaffolds is alsoimportant for the tissue’s development from cells, where cellculture medium and growth factors can be easily accessedthough the open pores. On the basis of previous studies[1–8] on the preparation and characterization of scaffoldsfor tissue engineering, open porosity, compressive strengthand feasibility for cell migration have been realized to bethe main criteria for good scaffolds.

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X. Miao et al. / Acta Biomaterialia 4 (2008) 638–645 639

For bone tissue engineering, the design of scaffoldsshould mimic the structure and properties of the boneextracellular matrices. Because bone consists of a porouscomposite of interpenetrating phases of hydroxyapatiteand collagen, the scaffolds for bone regeneration shouldbe similarly porous composites with interpenetrating cera-mic and polymer phases. Porous hydroxyapatite/tricalciumphosphate (HA/TCP) composite was thought to be idealfor the ceramic phase, as it is known for its excellent osteo-conductivity and, in some cases, even osteoinductivity [9].On the other hand, poly(lactic-co-glycolic) acid (PLGA)polymer was selected for the polymer phase, as it is anFDA-approved biodegradable polymer with some degreeof ductility and good biocompatibility [10,11]. Thus dualporous HA/TCP/PLGA composite scaffolds should mini-mize the problems confronted either with the porous solePLGA polymer (low compressive strength) or with the por-ous sole HA/TCP ceramics (mechanically weak and brit-tle). In fact, porous calcium phosphate scaffolds [2,3,5]have been toughened by either polymer coating on thestruts or polymer infiltration into the struts (if with openmicropores), or both.

Most studies on scaffolds have dealt with the mechanicalproperties and the cell or tissue ingrowth properties. Thethermal shock resistance of ceramic–polymer compositescaffolds has been largely forgotten. The reason for thiscould be that tissue engineering scaffolds are less likely tobe subjected to a temperature higher than body tempera-ture (37 �C), unlike the case of conventional engineeringceramics and composites, where thermal shock resistanceis an important thermomechanical property. However, itis justifiable to evaluate the thermal shock resistance ofthe HA/TCP/PLGA composite scaffolds. Firstly, the por-ous composite scaffolds are often subjected to a low tem-perature process, with the lowest temperature being thatof liquid nitrogen. For example, when cells are seeded ontothe scaffolds and undergo cell culturing, the engineeredcells/tissues may need to be stored in a freezer or even ina liquid nitrogen tank. Secondly, one active research area,that of bone-cartilage (osteochondral) tissue engineering,requires the preparation of bilayered composite scaffolds,where a porous polymer scaffold will need to be attachedonto a porous ceramic–polymer composite scaffold by alow temperature process, such as the thermally inducedphase separation method [6]. Thus, the resistance of thescaffolds to a low temperature thermal shock has beenidentified as another requirement for tissue engineeringscaffolds.

Successful tissue engineering also requires the uniformseeding of cells in scaffolds, and cell seeding should be fol-lowed by cell attachment, proliferation and differentiation,and secretion of extracellular matrices. No matter whethera dynamic cell culture, such as a perfusion bioreactor or aconventional static cell culture, is used, well-interconnectedpores are prime requirements for the scaffolds. Tissues haveoften been observed to develop preferentially around theperiphery of the scaffolds both in vitro and in vivo, which

could be due to the poor circulation of nutrients in the cen-tral part of the scaffold or to the poor pore interconnectiv-ity across the scaffold. Thus, it is important to evaluate thecell penetration and the cell attachment on the scaffoldsbefore other processes of the cell to tissue developmentare studied, and it is regrettable that such studies havenot previously been attempted with these calcium phos-phate/PLGA composite scaffolds.

2. Materials and methods

2.1. Preparation of composite scaffolds

2.1.1. Coating polyurethane (PU) foams with a ceramicslurry

The ceramic slurry was prepared by mixing 160 g of HA(average particle size 2 lm) with 40 g of b-TCP (averageparticle size 2.5 lm) using a ball mill under a wet conditionfor 2 h. The resulting paste was then dried at 100 �C for24 h followed by heat treatment at 900 �C for 2.5 h, withcooling and heating rates set at 5 �C min�1. The calcinedceramic pieces were milled again for 1 h before 125 ml ofdistilled water was added. The ceramic–water mixturewas further milled for another 24 h. Then 1 ml of 25%ammonium salt of polymethacrylic acid solution (DarvanC, R.T. Vanderbilt) was added to the ceramic paste, fol-lowed by mixing for about 30 min. Finally, 1 ml of2 wt.% polyvinyl alcohol solution was added to producethe final slurry. The PU foams were then dipped into theslurry and gently squeezed several times to allow the slurryto penetrate the foams before the excess slurry wassqueezed out. Compressed air from an airgun was usedto avoid the blockage of pores. The ceramic slurry-coatedPU foams were left to dry at room temperature for at least24 h.

2.1.2. Sintering of the ceramic-coated PU foams

The ceramic slurry-coated PU foams were fired in anelectric furnace (Modutemp Furnace) using a four-stageschedule, comprising (1) heating from room temperatureto 600 �C at a rate of 1 �C min�1 to burn out the PU foam;(2) raising the temperature from 600 to 1200 �C at a rate of5 �C min�1; (3) holding the temperature at 1200 �C for 4 hto sinter the ceramic; and (4) cooling the furnace down toroom temperature at a rate of 5 �C min�1. The HA/TCPscaffolds were removed from the furnace after it had cooleddown. Each sample was weighed and kept in a desiccator.

2.1.3. Coating the sintered HA/TCP scaffolds with PLGA

PLGA pellets (Sigma–Aldrich; PLA:PGA = 75:25; mol.wt = 90,000–126,000) were dissolved in dichloromethane(CH2Cl2) solvent, such that every 4 g of PLGA was dis-solved in 10 ml of dichloromethane. The sintered HA/TCP scaffolds with small dimensions of about10 · 10 · 10 mm3 were then immersed into the PLGA solu-tion for more than 30 s each to allow for complete infiltra-tion. The soaked scaffolds were then placed in a centrifuge

640 X. Miao et al. / Acta Biomaterialia 4 (2008) 638–645

running at 350 rpm for 1 min to remove the excess PLGAsolution from the scaffolds. The scaffolds were then takenout of the centrifuge tubes and left to dry in a fume hoodovernight, and were weighed individually before beingstored in a desiccator.

2.1.4. Quenching the scaffolds for evaluating the thermal

shock resistance

The HA/TCP scaffolds with and without PLGA coatingwere quenched by placing the scaffolds in a �80 �C freezerfor 30 min and then placing them at room temperature for30 min. This was done three times. Compression testingwas then performed on the quenched scaffolds, as describedin Section 2.2.3.

2.2. Sample characterization

2.2.1. Total porosity of the HA/TCP scaffolds without the

PLGA coating

When the HA/TCP scaffolds were not infiltrated and/orcoated with any PLGA, the pores responsible for the totalporosity were the macropores between the struts and themicropores within the struts. The total porosity of the sin-tered HA/TCP scaffolds was determined by using the fol-lowing equations: bulk density (qB) = weight of thesample divided by volume of the sample; theoretical densityof the HA/TCP composite (q0) = 3.16 g cm�2; relative den-sity (R.D.) = (qB/q0) · 100%; and total porosity = 100% �R.D. The dimensions and the weight of each sample weremeasured and recorded using a vernier calliper and an elec-tronic balance, respectively. Three identical specimens wereused to determine the total porosity.

2.2.2. Structural observation

The scaffolds were sectioned with a knife edge along thesagittal and transverse planes to give the best overview ofthe porous structure to confirm the pore interconnectivityand homogeneity. The scaffold samples were mounted ontoaluminium stubs with carbon tape and coated with goldfilm on a sputter coater (BioRad SC500). The porous struc-ture of the scaffolds was then examined using a scanningelectron microscope (FEI QUANTA 200) under an accel-eration voltage of 10 kV.

2.2.3. Compressive testing for scaffolds with and without

quenchingThe compressive strength and the compressive modulus

of each of the scaffolds were measured using a Hounsfieldtesting machine (Model: H10K/M527). The dimensionsof each sample were about 10 · 10 · 10 mm3. The scaffoldswere reasonably homogeneous and thus the orientation ofthe cut surfaces was not specially recorded. Rubber padswere placed on the top and bottom surfaces of each sampleto ensure an evenly distributed load on the sample. Thecross-head loading speed was set at 0.5 mm min�1. Sevento nine identical specimens for each sample group wereused for the compressive testing.

2.2.4. Fracture surface analysis

Various types of scaffolds subjected to the compressiontesting were collected and the fracture surfaces were exam-ined by scanning electron microscopy (SEM), as describedin Section 2.2.2.

2.3. In vitro evaluation by cell culture

2.3.1. Sample sterilization

The scaffolds with dimensions of 5 · 5 · 5 mm3 weredecontaminated by soaking them three times in 70% etha-nol for 15 min each, then rinsed three times with potassiumphosphate buffer solution (PBS) for 15 min before beingleft to dry overnight in a sterilized hood. The sterilized scaf-folds were sealed in a 24-well plate and kept in a refrigera-tor at 4 �C for later use.

2.3.2. Cell seeding and culture

Bone marrow stromal stem cells (BMSCs) maintained inour laboratories were used for the study. In brief, BMSCpellets expanded by subculture were resuspended in agrowth medium of the Dulbecco’s modified Eagle’s med-ium supplemented with 10% fetal bovine serum and 1%antibiotic–antimycotic. The cell suspension had a final con-centration of 2.4 · 106 cells ml�1. Then 100 ll of the cellsuspension was applied to each of the scaffolds. After 1 hof incubation, the wells were filled with the culture mediumand incubated at 37 �C for 1 day to observe the initial cellattachment.

2.3.3. Observation of cell penetration and attachment

The growth medium in each well was pipetted out,and immediately replaced with PBS. The rinsing wasrepeated three times for each sample, and then the scaf-folds were fixed with a 3% glutaraldehyde solution. Thescaffolds were then processed by soaking them in anosmium tetroxide solution for 1 h, then dehydratedthrough a series of ethanol solutions with graded concen-trations, followed by two changes of 100% amyl acetatefor 15 min each. The scaffolds were then dried using asupercritical point dryer (Denton Vacuum critical pointdryer) before observation using SEM, as shown in Sec-tion 2.2.2.

3. Results and discussion

3.1. Porous structures of HA/TCP scaffolds with and without

the PLGA coating

The HA/TCP composite porous scaffolds were fabri-cated by replicating the porous structure of the PU foams.The scaffolds, like the PU foams, had a highly intercon-nected structure, with open macropores (Fig. 1B and C)ranging from 300 to 700 lm in diameter. The average poresize of the macropores of the scaffolds was about 500 lm,estimated by the macropore sizes taken from the scanningelectron micrographs. The total porosity of the HA/TCP

Fig. 1. SEM micrographs of the polyurethane foam (A), the cut surface of the HA/TCP scaffold (B) and the polished cross-section of the HA/TCP scaffoldshowing the crack-like defects (arrowed) (C).

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scaffolds was determined to be approximately 87%. Themacroporosity of the scaffolds after infiltration and coatingwith the PLGA was slightly decreased due to the observedthin polymer coating present on the strut surfaces; the mac-ropores were not made significantly smaller by the polymercoating. In a similar study [12], Chen et al. had shown anabout 2.2% reduction in macroporosity as a result of thepolymer coating on the ceramic scaffolds. However, themicropores in the struts were infiltrated with the polymer;the weight gain after the PLGA infiltration and coatingcompared with the bare HA/TCP scaffolds indicated thatthe PLGA corresponded to �23 wt.% of the PLGA-coatedHA/TCP scaffolds, which was a significant amount andwould have a significant effect on the strength and tough-ness of the composite struts.

It was observed that the HA/TCP scaffolds had morecrack-like defects on and within the ceramic strutsthan those coated with PLGA. The crack-like defects

(Fig. 1C) of the HA/TCP scaffolds resulted from theburn-off of the PU struts that were initially coated withthe HA/TCP ceramic particles. However, the crack-likedefects in the HA/TCP struts could be filled with the poly-mer after the polymer solution dipping plus centrifuga-tion. Our previous studies [2,5] had demonstrated theremoval of the crack-like defects due to the PLGA poly-mer phase filling into the open voids. The struts of theHA/TCP scaffolds showed a smooth and sintered surfacethat revealed ceramic grains with sizes between 0.5 and2 lm. On the other hand, the strut surfaces of thePLGA-coated HA/TCP scaffolds had a thin PLGA layerwith a crater-like structure, which was left behind fromthe burst bubbles or caused by the evaporation of thechloroform solvent used to dissolve the PLGA. The thinand porous PLGA coating was advantageous as the bio-active HA/TCP struts were not fully covered or shieldedby the non-bioactive PLGA.

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3.2. Compressive strengths of HA/TCP scaffolds with and

without the PLGA coating

Compression testing was performed on both the HA/TCP scaffolds and the HA/TCP scaffolds coated with thePLGA polymer. It was found that the HA/TCP scaffoldswere able to withstand a maximum compressive stressbetween 0.05 and 0.07 MPa (Fig. 2A), while those coatedwith PLGA showed a compressive strength in the rangeof 0.62–0.79 MPa (Fig. 2B). Obviously the compressivestrength of the scaffolds coated with PLGA was substan-tially higher (by about 10-fold) than those without thePLGA polymer coating. As a general trend, the compres-sive strength of a scaffold depends on its macroporosity,macropore size and macropore geometry, and the strengthof the struts. In the present study, the �2% macroporosityreduction due to the polymer coating did not obviouslyinfluence the macroporous structural parameters. How-ever, the micropores of the HA/TCP scaffolds were filledwith PLGA, resulting in a significant amount of PLGA(i.e. 23 wt.%) in the whole PLGA-coated HA/TCP scaf-folds. In other words, the open micropores and/or crack-like defects in the struts of the scaffolds were infiltratedwith the PLGA phase, which made the original weak andbrittle ceramic struts into stronger and tougher ceramic–polymer composite struts. The formation of the compositestruts could explain the fact that the seemingly thin and

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Fig. 2. Three randomly selected stress–strain curves of the HA/TCPscaffolds (A) and the HA/TCP scaffolds coated with PLGA.

porous polymer coating could improve the compressivestrength of the ceramic scaffolds by more than 10-fold.Similar results were also observed by Chen et al. in a studyon the PDLLA-coated bioactive glass scaffolds [12]. Inaddition, the compressive modulus of the HA/TCP scaf-folds ranged from 2.21 to 2.99 MPa and those coated withPLGA were in the range of 14.08–19.29 MPa.

Polymer infiltration into and coating onto the struts ofthe HA/TCP scaffolds could reduce the defect sizes of theceramic struts and result in ceramic–polymer compositestruts, leading to the increased compressive strengths ofthe scaffolds. In general, ceramics can be strengthenedand toughened by the incorporation of a ductile polymeror a metallic phase. In the present study, the infiltrationof the polymer into the open porous ceramic struts led tothe formation of an interpenetrating ceramic–polymercomposite. The strengthening and toughening mechanismin an interpenetrating composite was studied by Pezzottiet al. [13]; they proposed a micron-scale crack-bridgingmechanism evident by the polymer ligaments that werestretched upon crack opening along the crack wake. Yetthe compressive strength and the compressive modulus ofthe HA/TCP scaffolds coated with the PLGA were stilllower than those of human cancellous bone. This wasdue to the high porosities and large pore sizes of the scaf-folds. However, it is believed that the compressive strengthof the scaffolds can be improved by applying a thickerPLGA coating, which could be realized by using repeatedpolymer solution dipping and centrifugation. A vacuuminfiltration process as shown in Peroglio et al.’s study [14]could be additionally used to facilitate the filling of thecrack-like defects of the HA/TCP ceramic struts. However,a thick PLGA coating has the problem of lack of bioactiv-ity, which was why bioactive glass particles were loadedinto the coating in our previous studies [2,5]. An optimizedcentrifugation process may also be needed to maintain thehigh degree of open macroporosity for the subsequent cellpenetration in the macropores.

3.3. Compressive strengths of the quenched HA/TCP

scaffolds with and without the PLGA coating

The resistance of the HA/TCP composite scaffolds withand without the PLGA coating to the quenching operationor process was also evaluated. It was found that the com-pressive strength and the compressive modulus of theHA/TCP scaffolds with and without the polymer coatingwere all badly affected by the quenching process. Specifi-cally, the compressive strength of the HA/TCP scaffoldsafter quenching was in the range of 0.02–0.03 MPa(Fig. 3A), and the compressive strength for the HA/TCPscaffolds coated with PLGA was in the range of 0.10–0.21 MPa (Fig. 3B). These compressive strength resultswere much lower than those without the quenching(Fig. 2). In other words, the compressive strength of theHA/TCP scaffolds had halved after quenching, while thoseof the coated scaffolds had decreased by about four times

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Fig. 3. Three randomly selected stress–strain curves of the HA/TCPscaffolds after thermal shock treatment (i.e. quenching) (A) and the HA/TCP scaffolds coated with PLGA after quenching.

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(Table 1). Fortunately, the porous HA/TCP scaffolds infil-trated and partially coated with the PLGA had sufficientmechanical strength for scaffold shaping and handling, inspite of the high porosities of the scaffolds and the harshtreatment of quenching the scaffolds.

3.4. Fracture surface morphologies with and without

quenching

Different fracture surface morphologies were observedfor the PLGA-coated HA/TCP scaffolds with and withoutquenching. For the case without quenching, the PLGA coat-ing was observed to be stretched slightly before breaking off,and the PLGA coating was featured with a ductile fracture

Table 1Effects of the PLGA coating and the thermal shock treatment on thecompressive properties of the HA/TCP scaffolds

Samples Thermalshocktreatment

Meancompressivestrength (kPa)

Meancompressivemodulus (kPa)

HA/TCP scaffolds No 60 ± 7 258 ± 24Yes 30 ± 4 108 ± 12

HA/TCP scaffoldscoated withPLGA

No 66 ± 7 168 ± 15Yes 16 ± 2 665 ± 50

surface due to some degree of ductility of PLGA. On theother hand, a flat fracture surface was observed for the cera-mic component, indicating the brittle nature of the ceramicmaterial. Most importantly, the remaining PLGA coatingwas still intact after fracture and attached closely to the cera-mic struts (Fig. 4A), which indicated the good interfacialbonding of the PLGA coating. For the case with quenching,the PLGA coating on the scaffolds was observed to detachfrom the ceramic struts (Fig. 4B). The PLGA coating wasalso observed to have been stretched before breaking offand the ceramic struts were still rather brittle.

The interfacial debonding was believed to be caused bythe different thermal shrinkage of the PLGA coating andthe ceramic struts during quenching. PLGA should havea much higher thermal expansion coefficient than that ofthe HA/TCP composite in the low temperature range (fromroom temperature down to liquid nitrogen temperature).Thus there must exist a high level of thermal stresses to

Fig. 4. SEM micrographs of the fracture surfaces of the PLGA-coatedHA/TCP scaffolds: (A) without quenching showing a good interfacialbonding and (B) after quenching showing the debonding of an interface(double arrowed).

Fig. 6. SEM micrographs showing the attachment of the stem cells onboth the exposed ceramic strut surface (A) and on the PLGA surface

644 X. Miao et al. / Acta Biomaterialia 4 (2008) 638–645

cause the interfacial separation. However, accurate analysisof the debonding is not possible due to the unknown ther-mal expansion coefficients and the complex structure of thePLGA-coated HA/TCP scaffolds. The debonding of thePLGA coating could also cause the decrease in compressivestrength of the PLGA-coated HA/TCP composite scaf-folds, as it is well known that interfacial bonding playsan important role in the strengthening and toughening ofceramic–polymer composite systems [15,16].

3.5. Initial cell penetration into and attachment onto the

PLGA-coated HA/TCP scaffolds

The bone marrow stromal stem cells were seeded into thescaffolds by adding drops of the cell suspension. As a result,the cells were not homogeneously seeded across the surfaceof the scaffolds. The penetration of the cells into the scaf-folds was evaluated using the cross-sections of the scaffolds.SEM observation (Fig. 5) revealed that the cells could pen-etrate to a depth of 4.5 mm, yet the cell density along thedepth direction was not homogeneous; the cells were mostplentiful on the surface and the cell density graduallydecreased with depth. The significant cell penetration depthwas due to the highly interconnected pores of the HA/TCPscaffolds coated with PLGA. The cell seeding homogeneitycan be improved by using a dynamic cell seeding method,such as the spinner flask stirring culture method.

The bone marrow stromal cells were observed to attachwell onto the strut surfaces of the PLGA-coated (actuallypartially coated) HA/TCP scaffolds; the cells could attachwell to the exposed ceramic HA/TCP surface and thePLGA coating surface (Fig. 6), indicating the biocompati-bility of both the ceramic phases and the PLGA phase.While the PLGA-coated HA/TCP scaffolds showedadequate cell penetration and attachment, the stem cells’proliferation and differentiation was not examined. It is

Fig. 5. SEM micrograph showing the attachment of the stem cells ontothe strut located at a depth of approximately 4.5 mm from the top surfaceof a PLGA-coated HA/TCP scaffold.

coated on the ceramic strut (B) of the PLGA-coated HA/TCP scaffold.

believed that the cells’ proliferation and differentiation iscontrolled by the struts’ surface chemistry, the surface mor-phology and potentially the presence of a growth factor.Thus, the PLGA-coated HA/TCP scaffolds can be furthermodified to increase not only the mechanical propertiesbut also the biological ones.

4. Conclusions

In the present study, HA/TCP composite scaffolds weremodified with PLGA. The low temperature thermal shockresistance and the initial cell penetration and the attach-ment were evaluated. The present study led to the followingconclusions:

(1) The pore size (�500 lm) and the pore interconnectiv-ity of the fabricated HA/TCP scaffolds closely mim-icked those of the initial polyurethane foams used.

X. Miao et al. / Acta Biomaterialia 4 (2008) 638–645 645

The average total porosity of the scaffolds was about87%. PLGA solution dipping followed by centrifuga-tion resulted in a thin layer of porous PLGA coating,which did not impair the open porous structure butimparted the coated scaffolds with mechanicalintegrity.

(2) The HA/TCP scaffolds modified with the PLGAcoating had a compressive strength (0.66 MPa) anda compressive modulus (16.85 MPa) which wereremarkably higher that those of the bare HA/TCPscaffolds (0.06 and 2.58 MPa, respectively). Afterquenching testing, the compressive strength and thecompressive modulus of the HA/TCP scaffolds mod-ified with PLGA were decreased to 0.16 and6.65 MPa, respectively, but the quenched scaffoldscould still tolerate the handling actions for the cellculture procedure.

(3) The highly porous and well interconnected scaffoldsenabled the bone marrow stromal stem cells to pene-trate to a depth of 5 mm. The seeded bone marrowstromal stem cells showed good initial attachmentonto both the HA/TCP biphasic ceramic and thePLGA polymer of the scaffolds.

The PLGA-coated HA/TCP scaffolds may be useful intheir own right for bone tissue engineering involving noor low load-bearing applications. The monolithic compos-ite scaffolds could also be useful for integrating with a solepolymer scaffold layer, so that the formed bilayered scaf-folds could be useful for osteochondral tissue engineering.Work is being done to modify the bilayered scaffolds anddirect the bidifferentiation of bone marrow stromal stemcells for osteochondral tissue engineering.

Acknowledgement

The research resulting in this paper was supported bythe 2006 IHBI research seeding grant.

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