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INTRODUCTION

BONE GRAFT SUBSTITUTES have been widely used toheal critical-size long bone defects and maxillofacial

skeleton defects due to trauma, tumor resection, congen-ital deformity, and tissue degeneration.

A vascularized and cancellous autograft shows opti-mal skeletal incorporation, because it is osteoconductive,osteoinductive, and osteogenic, but the harvesting of anautograft is traumatic and restricted by limited availabil-ity, causes morbidity at the donor site, and often bringsabout complications.1

Other bone graft substitutes are allograft, xenograft,and synthetic materials. The use of an allograft is suf-fering from limited sources and the possibility of thetransmission of diseases, such as human immuno-deficiency virus (HIV).2 As synthetic biomaterial, hy-droxyapatite shows superb biocompatibility and osteoconductivity, but low absorption after long-term im-plantation.3 Other synthetic biomaterials used to buildthree-dimensional (3D) scaffolds for bone tissue engi-neering are, for instance, the biodegradable poly(lacticacid) (PLA), poly(glycolic) acid (PGA), and their copoly-mer poly(lactic acid-glycolic acid) (PLAGA), where min-

TISSUE ENGINEERINGVolume 11, Number 5/6, 2005© Mary Ann Liebert, Inc.

Calcified Matrix Production by SAOS-2 Cells Inside aPolyurethane Porous Scaffold, Using a Perfusion Bioreactor

L. FASSINA, M.Sc.,1 L. VISAI, Ph.D.,2 L. ASTI, Ph.D.,3 F. BENAZZO, M.D.,3P. SPEZIALE, Ph.D.,2 M.C. TANZI, Ph.D.,4 and G. MAGENES, Ph.D.1

ABSTRACT

The repair and regeneration of damaged or resected bone are problematic. Bone autografts showoptimal skeletal incorporation, but often bring about complications. Hence, there is increasing in-terest in designing new biomaterials that could potentially be used in the form of scaffolds as bonesubstitutes. In this study we used a hydrophobic cross-linked polyurethane in a typical tissue-engi-neering approach, that is, the seeding and in vitro culturing of cells within a porous scaffold. Thepolyurethane porous scaffold had an average pore diameter of 624 �m. Using a perfusion bioreac-tor, we investigated the effect of shear stress on SAOS-2 human osteoblast proliferation and calci-fied matrix production. The physical, morphological, and compressive properties of the polyurethanefoam were characterized. At a scaffold perfusion rate of 3 mL/min, in comparison with static con-ditions without perfusion, we observed 33% higher cell proliferation; higher secretion of osteopon-tin, osteocalcin, decorin, and type I collagen (9.16-fold, 71.9-fold, 30.6-fold, and 18.12-fold, respec-tively); and 10-fold increased calcium deposition. The design of the bioreactor and the design of thepolyurethane foam aimed at obtaining cell colonization and calcified matrix deposition. This cul-tured biomaterial could be used, in clinical applications, as an osteoinductive implant for bone re-pair.

1Dipartimento di Informatica e Sistemistica, University of Pavia, Pavia, Italy.2Dipartimento di Biochimica, University of Pavia, Pavia, Italy.3Dipartimento SMEC, University of Pavia, IRCCS San Matteo, Pavia, Italy.4Dipartimento di Bioingegneria, Politecnico di Milano, Milan, Italy.

eralization of the extracellular matrix by osteoblasts hasbeen observed.4,5

Gorna and Gogolewski6,7 have drawn attention to theideal features of a cancellous bone graft substitute: itshould be biodegradable, porous with interconnectedpores of adequate size allowing for the ingrowth of cap-illaries and perivascular tissues; it should attract mes-enchymal stem cells from the surrounding bone and pro-mote their differentiation into osteoblasts; and it shouldcalcify in vivo and be elastomeric to avoid shear forcesat the interface between bone and bone graft substitute.The biomaterials for such an ideal bone graft substitutecould be, for instance, biodegradable porous polymer–ce-ramic matrices,8 starch-based polymer matrices,9 andbiodegradable polyurethane foams.6,7 Nevertheless, someproblems may arise when there is not a suitable equilib-rium between the material degradation in vivo and tissueregeneration.10,11

In this study we have used a hydrophobic cross-linkedpolyurethane in order to avoid possible problems due tobiodegradation: the particular chemical structure of ourpolyurethane is proof against biodegradation, and in gen-eral the susceptibility of polyurethane to biodegradationdecreases with increasing material hydrophobia12; there-fore our biomaterial could be defined as “biostable” inthe long term.13

In addition, this polyurethane has been designed13 inorder to follow the typical approach in tissue engineer-ing, an approach that involves the seeding and in vitroculturing of cells within a porous scaffold before im-plantation.14–17

The culture method is of great importance. Static cul-ture environments suffer from limited diffusion and of-ten result in inhomogeneous cell and extracellular matrixdistribution, confining tissue ingrowth to the outer sur-faces of the scaffold.18–20 To overcome the drawbacksassociated with static culture systems, several bioreactorshave been designed.19–29 The ideal bioreactor suppliessuitable levels of oxygen, nutrients, cytokines, growthfactors, and mechanical stimulation. Spinner flask biore-actors, in comparison with static systems, mitigate theexternal diffusional limitations, but not the internal lim-itations.20 To avoid this problem, Botchwey et al. havedeveloped a culture system that makes use of 3D micro-carrier scaffolds and the high-aspect-ratio vessel rotatingbioreactor21–24; Bancroft et al. have developed a flowperfusion bioreactor.26,27

In this study we have followed the path of the flowperfusion bioreactor, because it offers several advantagesfor bone tissue engineering.27 It provides enhanced wasteremoval and enhanced delivery of oxygen, nutrients, cy-tokines, and growth factors throughout the internal struc-ture of the culture scaffold, it stimulates by fluid shearstress the seeded osteoblasts, the cell function of whichis mechanically modulated.29–39

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The aims of this work were to investigate the effect ofshear stress on SAOS-2 human osteoblast proliferation,and on calcified matrix production, using a perfusionbioreactor. Thus we attempted to populate a porouspolyurethane scaffold with bone cells and their extracel-lular matrix.

For bone reconstruction, a possible alternative path tobiodegradation may be the in vivo integration of abiostable scaffold, because biodegradable material mayelicit in vivo a foreign body inflammatory reaction withbone resorption.10 In this way, our scaffold, filled withhuman bone proteins and calcium minerals, could be usedin clinical applications as an osteoinductive agent forbone repair.40

MATERIALS AND METHODS

Polyurethane foam composition, synthesis, andcharacterization

Cross-linked polyurethane foam was synthesized byone-step bulk polymerization, using a polyether–polyolmixture (component A, ElastoCoat 96827/4, EC; Elas-togran, Villanova d’Asti, Italy) with polymeric MDI(component B, B141; BASF, Ludwigshafen, Germany)using 0.001% (w/wA) iron acetylacetonate as catalyst and2% (w/wA) water as expanding agent. Water and ironacetylacetonate were added to the weighed amount ofcomponent A, and mixed with a mechanical stirrer at2000 rpm for 40 s. The appropriate quantity of compo-nent B was added, and the reaction mixture was stirredfor 90 s and poured into a custom-made polymethyl-methacrylate mold in order to allow expansion of thefoam into a fixed volume. The mold was firmly closedand the expansion reaction was allowed to take place atroom temperature by CO2 production from the water–iso-cyanate reaction. The polyurethane foam was extractedfrom the mold after 24 h, and postcured at room tem-perature for 7 days. To remove potentially noxious mol-ecules characterized by low molecular weight, that is, by-products and unreacted matter, the foam was purified byimmersion in absolute ethanol for 48 h at room temper-ature.13 Cylindrical specimens (diameter, 15 mm; height,10 mm) were cut from the foam with a manual die formorphological and mechanical characterization. Densityof the cylindrical specimens was evaluated according toUNI EN ISO 845 standard practice. Porosity, expressedas the percentage of open pores, was evaluated accord-ing to EN ISO 4590 standard practice with a home cus-tom-made picnometer, which measures gas-impenetrablevolume and compared with the total volume of rigid cel-lular plastic. The average pore size was determined ac-cording to ASTM D 3376-94 standard practice. A com-pressive mechanical test was performed on fivecylindrical specimens with an Instron (Canton, MA)

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model 4200 instrument at a cross-head rate of 1 mm/min,both under dry conditions and after 7 days of treatmentin distilled water (the tests were performed by maintain-ing the specimens in water at 37°C for the entire testtime). Compressive moduli Edry and Ewet, in dry and wetconditions, respectively, were obtained from stress–straincurve elaboration. Cell culture scaffolds (diameter, 12mm; height, 6 mm) were cut from the foam with a man-ual die.

Cells

The human osteosarcoma cell line SAOS-2 was ob-tained from the American Type Culture Collection(HTB85; ATCC, Manassas, VA). The cells were culturedin McCoy’s 5A modified medium with L-glutamine andHEPES (Cambrex Bio Science Baltimore, Baltimore,MD), supplemented with 15% fetal bovine serum, 2% so-dium pyruvate, 1% antibiotics, 10�8 M dexamethasone,and 10 mM �-glycerophosphate (Sigma-Aldrich, Mil-waukee, WI). Ascorbic acid, another osteogenic supple-ment, is a component of McCoy’s 5A modified medium.The cells were cultured at 37°C with 5% CO2, routinelytrypsinized, counted, and seeded onto the porouspolyurethane scaffolds.

Cell seeding

The scaffolds were sterilized by ethylene oxide at 38°Cfor 8 h at 65% relative humidity.41 After a 24-h aerationin order to remove residual ethylene oxide, the scaffoldswere placed inside the two culture systems: the “static”culture system, that is, a standard well plate, and the “dy-namic” culture system, that is, the perfusion bioreactor.A cell suspension of 7 � 105 cells in 300 �L was addedonto the top of each scaffold and, after 2 h, sufficient Mc-Coy’s medium was added to cover the scaffolds. Cellswere allowed to attach overnight, and then the static cul-ture continued in the standard well plate and the perfu-sion bioreactor was turned on.

Perfusion bioreactor

A flow perfusion bioreactor was built (Fig. 1A). It con-sisted of four flow chambers custom-machined in a blockof Plexiglas, that is, polymethylmethacrylate (Fig. 1B).A polyurethane scaffold was contained in each flowchamber.

A multichannel peristaltic roller pump (Cole-Parmer,Vernon Hills, IL) moved the culture medium through thepolyurethane scaffolds: medium flowed from top to bot-tom in order to prevent the trapping of air bubbles be-neath the scaffolds.26,27 The dynamic culture system en-sured perfusion through the scaffolds at a knowncontrollable flow rate. The peristaltic pump moved theculture medium from a common reservoir toward theflow chambers, and the medium was pumped through the

FLOW PERFUSION IN POLYURETHANE: SAOS-2 EFFECTS

scaffolds and returned to a second reservoir (Fig. 1A).Flow chambers and reservoirs were connected by plat-inum-cured silicone tubing (Cole-Parmer), which has ahigh permeability to the metabolic gases O2 and CO2.

All circuit components were sterilized by autoclave ex-cept for the flow chambers, which were sterilized withethylene oxide. The flow perfusion bioreactor was as-sembled, filled with 150 mL of culture medium, andplaced into a standard cell culture incubator with an en-vironment of 37°C and 5% CO2. The scaffolds were cul-tured at a flow rate of 3 mL/min for a total of 16 days.The flow rate was, initially, equal to 0.3 mL/min to al-low better cell attachment and, after a 1-day condition-ing period, it was set at 3 mL/min.

Flow perfusion can improve an osteoblast culturethrough chemotransport enhancement and the action offluid shear forces.26 Goldstein et al.25 proposed a cylin-drical pore model approximation to evaluate the localshear stress that bone cells experience within a scaffold.Taking Goldstein’s model into account and assuming aparabolic flow velocity profile of the culture mediumwith a viscosity equal to 0.01 g/(cm � s), given the per-

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FIG. 1. Scheme of the flow perfusion bioreactor (A) and flowchambers (B).

fusion rate, the scaffold geometry, and the geometry ofthe polyurethane porosity (Table 1), the mean velocityVm through the pores and the wall shear stress �w wereequal to 0.5 cm/s and 0.56 dyn/cm2, respectively. Ban-croft et al.26 demonstrated that osteoblasts are sensitiveto such limited shear stress when mechanical stimulationis continuous for an extended period. The culture mediumwas changed on days 4, 7, 10, and 13, and the condi-tioned culture medium was sampled for analysis on days4, 7, 10, 13, and 16.

Standard well plate culture

Seeded polyurethane scaffolds were simply coveredwith fresh medium on days 4, 7, 10, and 13, and condi-tioned culture medium was sampled for analysis on days4, 7, 10, 13, and 16.

Light microscopy analysis

Cultured polyurethane scaffolds were fixed in forma-lin, dehydrated in a gradient ethanol series up to 100%,and embedded in paraffin. Histological sections (thick-ness, 10 �m) were cut orthogonally to the scaffold axis,and stained with hematoxylin–eosin. Images were takenwith a standard light microscope (Leica Microsystems,Bensheim, Germany) equipped with a digital image cap-ture system (Canon, Tokyo, Japan) at �63 magnifica-tion.

Light microscopy analysis for calcium content detection

Cultured polyurethane scaffolds were fixed in forma-lin, dehydrated in a gradient ethanol series up to 100%,and embedded in paraffin. Histological sections (thick-ness, 10 �m) were cut orthogonally to the scaffold axis,and stained by the von Kossa method (04-170801; Bio-Optica, Milan, Italy). This staining substitutes calcifiedmatrix Ca2� ions with silver atoms, revealing the calciumminerals as black spots (Fig. 2): the ratio between blackspot area and total image area was measured with Im-ageJ (http://rsb.info.nih.gov/ij), and a comparison wasperformed between the dynamic culture and the static cul-ture.

DNA content

Cells were lysed by a freeze–thaw method in steriledeionized distilled water. The released DNA content was

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evaluated with a fluorometric DNA quantification kit(PicoGreen; Molecular Probes, Eugene, OR). A DNAstandard curve, obtained from a known amount of os-teoblasts, was used to express the results as cell numberper scaffold.

Osteopontin measurement

Osteopontin secretion in the culture medium was mea-sured with an immunoenzymatic assay kit (900-27; AssayDesigns, Ann Arbor, MI). The conditioned medium wasstored at �20°C in order to prevent osteopontin degrada-tion and, after thawing, it was diluted to a 1:2 ratio. Thesandwich immunoassay is based on two antibodies to hu-man osteopontin: a polyclonal antibody immobilized on amicrotiter plate and a monoclonal antibody labeled withthe enzyme horseradish peroxidase. Osteopontin secretionsare expressed as femtograms per (cell � scaffold). TheGompertz growth model

y � y0e(b/a)(1�e�ax)

was applied, with MATLAB (The MathWorks, Natick,MA), to the total secretion data measured from the dy-namic culture system, whereas a linear growth model wasapplied to the static culture system.

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TABLE 1. PHYSICAL, MORPHOLOGICAL, AND COMPRESSIVE PROPERTIES OF POLYURETHANE FOAM

Density Average pore diameter Open porosity Edry Ewet

(g/cm3) (�m) (%) (MPa) (MPa)

0.098 � 0.002 624 80 � 2 9.82 � 0.40 4.44 � 0.31

FIG. 2. von Kossa staining of a histological section (A) andcalcium image elaboration with ImageJ: the ratio between blackarea and total area was measured (B).

Osteocalcin measurement

Osteocalcin secretion in the culture medium was mea-sured with an immunoenzymatic assay kit (BT-480; Bio-medical Technologies, Stoughton, MA). The conditionedmedium was stored at �80°C in order to prevent osteo-calcin degradation. The sandwich immunoassay is basedon two monoclonal antibodies. The lowest detectable os-teocalcin content with this assay was 0.5 ng/mL. Osteo-calcin secretion is expressed as femtograms per (cell �scaffold). A linear growth model was applied, with MAT-LAB, to the total secretion data measured from both cul-ture systems.

Anti-decorin polyclonal mouse antibody

Decorin was purified as previously described42 and theprotein content of decorin preparation was determined bythe Bradford method.43 Electrophoretic analysis underdenaturing conditions was done according to Laemmli,44

both before and after chondroitinase ABC digestion.45

Analysis of disaccharides of the glycosaminoglycanchains was performed after digestion with chondroitinaseABC or AC II according to standard methods.46

For anti-decorin polyclonal antibody production,BALB/c mice were injected intraperitoneally with 50 �gof the antigen, which was emulsified with an equal vol-ume of Freund’s complete adjuvant for the first immu-nization, followed by four injections in Freund’s incom-plete adjuvant. The mice were bled and sera were testedfor reactivity against purified decorin by enzyme-linkedimmunosorbent assay (ELISA) and Western blot analy-ses.

Antibodies were purified by affinity chromatographyon protein G–Sepharose columns according to the rec-ommendations of the manufacturer (Amersham Bio-sciences, Uppsala, Sweden) and used in an ELISA.

The anti-decorin polyclonal antibody was tested forpositive cross-reactivity against human decorin.

Decorin measurement

A calibration curve to measure decorin was producedby ELISA. Microtiter wells were coated with increasingconcentrations of decorin, from 1 ng to 1 �g, in coatingbuffer (50 mM Na2CO3, pH 9.5) overnight at 4°C. Someof the wells were coated with bovine serum albumin(BSA) as a negative control. After three washes withPBST (PBS containing 0.1% [v/v] Tween 20), the wellswere blocked by incubation with 200 �L of PBS con-taining 2% (w/v) BSA for 1 h at 22°C. The wells weresubsequently incubated for 1.5 h at 22°C with 100 �L ofanti-decorin polyclonal antibody IgG (1 �g/mL in 1%BSA). After washing, the wells were incubated for 1 hat 22°C with 100 �L of horseradish peroxidase (HRP)-conjugated rabbit anti-mouse IgG (1:1000 dilution in 1%

FLOW PERFUSION IN POLYURETHANE: SAOS-2 EFFECTS

BSA). The wells were finally incubated with 100 �L ofdevelopment solution (phosphate–citrate buffer contain-ing o-phenylenediamine dihydrochloride substrate pre-pared according to the manufacturer’s specifications).The color reaction was stopped with 100 �L of 0.5 MH2SO4 and absorbance values were measured at 490 nmwith a microplate reader (Bio-Rad, Hercules, CA).

Similarly, microtiter wells were coated, overnight at4°C, with 200 �L of conditioned medium produced un-der static and dynamic conditions. After washing,blocked wells were incubated as previously describedwith anti-decorin polyclonal antibody IgG (1 �g/mL in1% BSA) and finally incubated for 1 h at 22°C with HRP-conjugated rabbit anti-mouse IgG (1:1000 dilution in 1%BSA). The color development reaction was performedand absorbance values were measured as previously de-scribed. Decorin secretions are expressed as femtogramsper (cell � scaffold). A linear growth model was applied,with MATLAB, to the total secretion data measured fromboth culture systems.

Type I collagen measurement

Anti-type I collagen polyclonal antibody IgG was agenerous gift from L. Fisher (antiserum LF-67; NationalInstitutes of Health, National Institute of Dental andCraniofacial Research, Craniofacial and Skeletal Dis-eases Branch, Bethesda, MD).47,48 The assay procedurewas similar to the decorin assay procedure. Type I col-lagen secretions are expressed as femtograms per (cell �scaffold). A linear growth model was applied, with MAT-LAB, to the total secretion data measured from both cul-ture systems.

Scanning electron microscopy analysis

Scaffolds were fixed with 2.5% (v/v) glutaraldehydesolution in 0.1 M sodium cacodylate buffer (pH 7.2) for1 h at 4°C, washed with sodium cacodylate buffer, andthen dehydrated at room temperature in a gradient etha-nol series up to 100%. The samples were kept in 100%ethanol for 15 min, and then critical point-dried with CO2.The specimens were mounted on aluminum stubs, sput-ter coated with gold (degree of purity equal to 99%), andthen observed with a Leica Cambridge Stereoscan 440microscope at 8 kV. The unseeded scaffolds were ob-served at �18 and �240 magnifications, whereas the cul-tured scaffolds at �450 and �550 magnifications.

Statistics

Results are expressed as means � standard deviation.To compare the results between static and dynamic sys-tems at each culture medium change, one-way analysisof variance (ANOVA) with post hoc Bonferroni test wasapplied, electing a significance level of 0.05.

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RESULTS

The osteoblasts were seeded onto polyurethane porousscaffolds, and then cultured in a flow perfusion bioreac-tor for 16 days. This culture system permitted the studyof SAOS-2 cells as they proliferated and produced a cal-cified extracellular matrix in a 3D mechanically activeenvironment. We compared the cell–matrix distribution,calcified matrix production, and microscope images be-tween the two culture systems.

Polyurethane foam characterization

Physical, morphological, and compressive propertiesof the polyurethane foam are reported in Table 1. Thehigh porosity value was related to a low density. By scan-ning electron microscopy (SEM) observation the foammorphology appeared rather uniform, with smooth poresurfaces (Fig. 3A and B). The pores appeared intercon-nected. As expected, the polyurethane foam exhibited ahigher compressive modulus under dry compared withwet conditions; under both conditions specimens showedthe typical stress–strain behavior of elastoplastic foams:the initial linear elastic regimen was followed by a plateauof roughly constant stress corresponding to pore collapse.

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Light microscope and SEM analyses

Histological sections and SEM images revealed thatflow perfusion improved cell distribution inside the scaf-folds.

Statically cultured cells were few and essentially or-ganized in a monolayer with a thin discontinuous extra-cellular matrix (Figs. 4A and 5A and B), whereas shearstress induced a 3D modeling of the cell–matrix organi-zation: several cells were organized in a multilayer witha highly developed matrix (Figs. 4B and 5C and D). Sta-tically cultured scaffolds showed the presence of cellsand matrix only on the upper surface; perfusion improvedcell penetration, but the center of the scaffolds remaineddevoid of cells and matrix.

These observations were confirmed by the measure-ment of DNA content after 16 days of culture. In the sta-tic culture the cell number per scaffold grew to 11.2 �106 � 6.4 � 104, whereas in the dynamic culture the cellnumber per scaffold was 17.4 � 106 � 8 � 104, with

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FIG. 4. Histological sections of a static culture (A) and a dy-namic culture (B): Osteoblast layers adhered to the polyurethanescaffolds (arrows). Hematoxylin–eosin staining; original mag-nification, �63.

FIG. 3. SEM observation of an unseeded polyurethane scaf-fold. Original magnification: (A) �18; (B) �240.

p � 0.05: the shear stress and consequent prostaglandinsecretion34–36 stimulated cell proliferation at a ratio ofabout 2:3. Because the DNA may remain entrapped inthe calcified matrix, an underestimation of culture cellu-larity is possible.

Calcified matrix deposition

The relative amount of calcium contained in the scaf-folds was quantified in order to evaluate matrix calcifi-cation. The ratio between the black area and the total im-age area was 0.2 � 0.08% in the static culture, whereasit was 2.0 � 0.4% in the dynamic culture: the fluid flowperfusion increased matrix calcification by about 10-fold,with p � 0.05.

Measurement of osteopontin, osteocalcin, decorin,and type I collagen secretions

Osteopontin, osteocalcin, decorin, and type I collagenwere detected in both culture systems; nevertheless,higher levels were found in the dynamically conditionedmedium (Figs. 6A, 7A, 8A, and 9A): at each mediumchange the difference between the static and the dynamicsystems was significant (p � 0.05).

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Secretion of osteopontin in the static culture wassteady, whereas in the dynamic culture it showed a rise-and-fall temporal pattern with an expression peak be-tween days 7 and 10 (Fig. 6A).

Total osteopontin secretion showed two different tem-poral patterns (Fig. 6B). In the static culture a linearmodel was applied, and a secretion rate of 0.51 fg/(cell �day � scaffold) was measured (R2 � 0.992). The Gom-pertz growth model was applied to the dynamic culture:the highest secretion rate was observed between days 8and 12, and it was equal to 7.29 fg/(cell � day � scaf-fold) (R2 � 0.999).

Osteocalcin secretion was steady (Fig. 7A), and totalsecretion was linear in both culture systems (Fig. 7B):we observed a secretion rate of 0.28 fg/(cell � day �scaffold) (R2 � 0.983) and of 17.30 fg/(cell � day �scaffold) (R2 � 0.998) in the static and dynamic systems,respectively.

Decorin secretion was steady (Fig. 8A), and totaldecorin secretion showed a linear temporal pattern in bothculture systems (Fig. 8B): we observed a secretion rateof 0.45 fg/(cell � day � scaffold) (R2 � 0.996) in thestatic culture, and a secretion rate of 12.63 fg/(cell �day � scaffold) (R2 � 0.998) in the dynamic culture.

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FIG. 5. SEM observations of a static culture (A and B) and a dynamic culture (C and D). Original magnification: (A and C)�450; (B and D) �550.

Similar remarks can be made for type I collagen (Fig.9A and B): we observed a secretion rate of 10.76fg/(cell � day � scaffold) (R2 � 0.951) and of 185.70fg/(cell � day � scaffold) (R2 � 0.995) in the static anddynamic cultures, respectively.

An overview of the enhancements caused by the per-fusion bioreactor is reported in Table 2. These observa-tions were in accordance with the microscope images ofhighly developed matrix (Figs. 4B and 5C and D).

DISCUSSION

We reported in the introduction the “golden rules” ofGorna and Gogolewski about the design of a bone graft

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substitute.6,7 Their biodegradable cross-linked poly-urethanes are suitable to bone repair in terms of pore size,porosity, and mechanical properties; the mechanicalproperties of such substitutes are of minor importance, asthe mechanical stability of the bone–scaffold system maybe ensured by metallic internal fixation devices.7 Ourpolyurethane was similar in terms of pore size, porosity,and mechanical properties, nevertheless, it was non-biodegradable owing to its chemical structure and hy-drophobia. This design decision was justified. Previouspapers have shown the ability of microporous and macro-porous biodegradable polymeric scaffolds to support invitro cell adhesion and growth.5,49,50 The long-term invivo performance of these scaffolds is dependent mainlyon the rate of biomaterial degradation and on the ability

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FIG. 6. Osteopontin secretion at each medium change (A) and total secretion (B): At each medium change perfusion increasedosteopontin synthesis (p � 0.05).

to form new bone. Two problems must be considered:first, the rate of degradation is affected by the biomater-ial chemical structure and morphology, specifically theamorphous/crystalline content, and, second, the degrada-tion products, although physiologically well accepted,cause a local inflammatory response that decreases therate of tissue regeneration. Therefore our polyurethanewas made nonbiodegradable. “Nonbiodegradable” doesnot mean “nondegradable,” because all commercial andnoncommercial nonbiodegradable polyurethanes are sub-ject more or less to oxidation, hydrolysis, crack forma-tion, and calcification in the long term (the susceptibil-ity of polyurethanes to calcification, which is notdesirable when these biomaterials are used in cardiovas-cular implants, may become a required property in bonetissue engineering). For that reason our biocompatible

FLOW PERFUSION IN POLYURETHANE: SAOS-2 EFFECTS

polyurethane may be defined as relatively “biostable” inthe long term, in other words “biointegrable,” in whichthe biointegrability is a consequence of the biocompati-bility and biostability.51 Although degradation is a com-mon requirement for scaffolds, this work was aimed atthe alternative approach of scaffold biointegration, a con-cept derived from nonpolymeric scaffolds for osseointe-gration. The utilization of this polyurethane foam wouldovercome the problems associated with biodegradablepolymers, such as degradation kinetics asynchronouswith tissue regeneration and adverse reactions to degra-dation products.51,52

The choice of SAOS-2 osteoblasts was also justified.An ability to induce the formation of new bone at a spe-cific site would represent a significant advance in bonerepair and tissue engineering. This property seems to be-

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FIG. 7. Osteocalcin secretion at each medium change (A) and total secretion (B): At each medium change perfusion increasedosteocalcin synthesis (p � 0.05).

long to SAOS-2 cells. These osteoblasts uniquely con-tain an osteoinductive activity, whereas other human os-teosarcoma cells, such as U-2 OS, cannot replicate thatbone-inducing ability.40 Devitalized SAOS-2 cells, ex-tracts, and secretions induced the formation of new bonewhen implanted subcutaneously in nu/nu mice.53–55 Thebone-inducing ability has also been found in secretionsof SAOS-2 cells in conditioned culture medium.56 Theseosteoblasts can be grown, virtually indefinitely, to pro-duce large quantities of osteoinductive factors, such asbone sialoprotein and osteonectin, at a fraction of the costof the individual recombinant proteins.40

To facilitate SAOS-2 culture, we decided to utilize aperfusion bioreactor, because it is efficient in terms ofmetabolic gas exchange, supply of culture medium with

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growth factors and osteogenic agents, and supply andmaintenance of a known, controllable, and mechanicallyactive environment. The osteoblasts, during mineralizedtissue formation, show strong metabolic demands, so ef-ficient nutrient delivery is fundamental to bone tissue en-gineering. Botchwey et al.21–24 developed a microcarriermatrix system for osteoblast culture inside a rotatingbioreactor, and elaborated a theoretical model of nutrienttransport through the scaffold. The Botchwey modelshowed that passive glucose diffusion was unable to pro-vide the osteoblasts with the minimum glucose concen-tration beyond a few hundred microns within the scaf-fold, whereas the Botchwey rotating bioreactor and theGoldstein perfusion bioreactor25 appeared extremely ef-ficient in nutrient transport. In comparison with Gold-

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FIG. 8. Decorin secretion at each medium change (A) and total secretion (B): At each medium change perfusion increaseddecorin synthesis (p � 0.05).

stein’s design parameters, in our perfusion system the av-erage pore size was larger, the porosity was higher, thescaffold thickness was exactly the same, the cell densitywas lower, and the mean velocity Vm through the poreswas higher. Therefore, considering the Botchwey model,our perfusion bioreactor was efficient in nutrient deliv-ery.

In this study we have demonstrated that polyurethanescaffolds can be suitable for calcified matrix productionunder perfusion conditions, making the biomaterial use-ful for biointegration. We have reported that, under dy-namic conditions, secretions of osteopontin, osteocalcin,decorin, and type I collagen were greatly enhanced.

In comparison with static culture, perfusion increasedtotal osteopontin secretion about 9-fold and the maximum

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secretion rate about 14-fold. Osteopontin is an extracel-lular glycosylated bone phosphoprotein secreted duringthe early stages of osteogenesis, before the onset of min-eralization. It binds calcium, it is probably involved inthe regulation of hydroxyapatite crystal growth,57,58 and,through specific interaction with the vitronectin receptor,it promotes the attachment of cells to the matrix.59 Thisprocess is necessary for matrix mineralization. We ob-served the maximum secretion rate between days 8 and12. A previous study58 has shown the temporal sequenceof cytoplasmic maturation and secretion of the matrixproteins: after 4 days of culture, osteocalcin, bone sialo-protein, osteopontin, and osteonectin accumulated insidethe osteoblasts, and after 8 days of culture, when the os-teoblasts reached confluency, the authors observed the

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FIG. 9. Type I collagen secretion at each medium change (A) and total secretion (B): At each medium change perfusion in-creased type I collagen synthesis (p � 0.05).

secretion of bone sialoprotein and osteopontin, their or-ganization into fibers, and the epitactic nucleation andgrowth of hydroxyapatite crystals. Therefore the mea-surement of total osteopontin secretion on various daysduring the whole culture period and its interpretation ac-cording to the Gompertz growth model allowed us to hy-pothesize that, inside the polyurethane scaffolds underperfusion conditions, SAOS-2 cells reached confluencyat day 8 and developed a calcifying multilayer in the fol-lowing days. The choice of the Gompertz growth modelwas justified, because this model can easily describe ageneral biological phenomenon in which three temporalpatterns are recognized: the first with a rapidly increas-ing rate, the second with a constant and maximum rate,and the third with a decreasing rate.60,61

The production of a mineralized matrix is peculiar tothe osteoblastic phenotype; the relative amount of cal-cium contained in the scaffolds was quantified in orderto evaluate matrix calcification. In comparison with sta-tic culture, perfusion increased the matrix calcificationabout 10-fold, and this result was in agreement with the�9-fold increase in osteopontin secretion.

FASSINA ET AL.

Bancroft et al.26 showed that marrow stromal os-teoblasts, under perfusion conditions, began to secrete os-teocalcin after 13 days of culture. In our study SAOS-2cells, which are mature osteoblasts, secreted osteocalcinwith a linear temporal pattern during the whole cultureperiod: in comparison with static culture, perfusion in-creased total osteocalcin secretion about 72-fold and in-creased the secretion rate about 60-fold. Osteocalcin, themajor noncollagenous and vitamin K-dependent bonematrix protein, is a marker for differentiated and matureosteoblasts. Osteocalcin is small (5800 D), and it is char-acterized by �-carboxyglutamic acid residues: the ex-pression of osteocalcin is induced after the onset of min-eralization; therefore, in the presence of calcium, theseresidues allow a specific conformational change in theprotein, which in turn promotes osteocalcin binding tobone minerals and consequent accumulation in the ma-trix.62–64

We have reported the secretion of decorin under dy-namic conditions. The temporal pattern of decorin se-cretion was linear in both culture systems: in comparisonwith static culture, perfusion increased total decorin se-cretion about 31-fold and increased the secretion rateabout 28-fold. Decorin is a member of the family of ex-tracellular matrix proteoglycans characterized by a pro-tein core containing 10 tandem leucine-rich repeats, eachof about 24 amino acids, flanked by cysteine clus-ters.65–67 Leucine-rich repeats are involved in protein–protein interactions. Determination of the decorin levelis important because this proteoglycan is considered akey regulator for the assembly and function of many ex-tracellular matrix proteins. Decorin plays a major role inthe lateral growth of collagen fibrils, delaying lateral as-sembly on the surface of the fibrils.68,69 This might con-trol fibril dimensions, uniformity of the fibril diameter,and regular spacing of the fibrils. Evidence showed thatthere are interactions between decorin and transforminggrowth factor �, and between decorin and epidermalgrowth factor receptors.70,71

Luan et al.58 found a linear temporal pattern in totalalkaline phosphatase activity and in total type I collagenaccumulation in a static culture of osteoblasts. Perfusionincreased the total secretion of type I collagen about 18-fold and increased the secretion rate about 17-fold. TypeI collagen is the most important and abundant structuralprotein of bone matrix. Therefore we could appreciatethe upregulation, the mechanical modulation of SAOS-2phenotype intensity inside the perfusion bioreactor, thatis, inside a mechanically active environment.

Remarks about the biomaterial–protein–cell system arenecessary. Interactions between the extracellular matrixproteins and cells can directly control cell behavior suchas attachment, migration, proliferation, differentiation,and apoptosis,72 and there has been active research onthe effects of material surface properties on cell behav-

696

TABLE 2. ENHANCEMENTS IN EXTRACELLULAR MATRIX

SYNTHESIS CAUSED BY PERFUSION BIOREACTOR

Total secretion after 16 daysof culture [fg/(cell � scaffold)]

Static culture Dynamic culture

Osteopontin 7.33 67.18Osteocalcin 3.82 274.67Decorin 6.42 196.47Type I collagen 147.49 2673.75

Static culture Dynamic culture

Osteopontin 0.51 7.29 (maximum rate)Osteocalcin 0.28 17.30Decorin 0.45 12.63Type I collagen 10.76 185.70

Total secretionafter 16 days

of culture Secretion rate

Osteopontin 9.16-fold 14.29-fold (maximum rate)Osteocalcin 71.90-fold 60.21-foldDecorin 30.60-fold 27.80-foldType I collagen 18.12-fold 17.25-fold

Enhancements caused by the perfusion bioreactor

Secretion rate [fg/(cell � day � scaffold)]

ior.73 It is well known that a rapid adsorption of proteinsis the first event occurring when a biomaterial is im-planted or comes into contact with culture mediumserum.74,75 Osteoblast attachment is mediated throughadsorbed proteins such as fibronectin74,76 and vit-ronectin,77,78 and protein adsorption is enhanced by thesurface hydrophobia.76,79

We could suppose that osteoblasts adhered to the hy-drophobic polyurethane through serum-adsorbed pro-teins, and that they then adhered to secreted matrix pro-teins, especially type I collagen.80 Perfusion made moreadhesion proteins available, in other words it permittedmore efficient development of the biomaterial–matrix–cell system.

CONCLUSIONS

In this study a flow perfusion system was applied to abiostable polyurethane foam instead of a metallic fibermesh.26,28 The perfusion bioreactor accelerated the me-tabolism of SAOS-2 osteoblasts in two ways: in com-parison with the static system and other less efficientbioreactors, it enhanced chemotransport and providedsuitable levels of shear stress.

Statically cultured SAOS-2 osteoblasts showed normalproliferation, but their rate of matrix synthesis was low.In the perfusion bioreactor, that is, in a mechanically ac-tive environment, the osteoblasts not only proliferated,but also showed intensely their characteristic phenotype.

Calcified matrix filled scaffolds, and dynamically con-ditioned media abounding in osteoinductive factors couldbe used as products for bone repair and other tissue-en-gineering applications40,53–56; nevertheless, a better result could be obtained with bone marrow stromal os-teoblasts instead of SAOS-2 cells for total immunocom-patibility with the patient.

Another fundamental property of conditioned mediumcontaining bone sialoprotein is angiogenesis induction81;therefore it could be interesting to use SAOS-2 dynami-cally conditioned medium in a human umbilical vein en-dothelial cell culture with fibronectin coating.82

In conclusion, biointegrable polyurethane foams couldcontain osteoconductive bioabsorbable calcium phos-phates, hence improving the mechanically stimulating en-vironment.

ACKNOWLEDGMENTS

The authors thank Prof. L.W. Fisher, Prof. A. Icaro-Cornaglia, Prof. A. Calligaro, Prof. F. Auricchio, Dr. S.Farè, Dr. P. Petrini, Dr. C. Giudici, Mr. A. Mortara, andMr. F. Barzon for support. This work was supported byFAR grants (2003) from the University of Pavia to Prof.F. Benazzo and to Prof. G. Magenes.

FLOW PERFUSION IN POLYURETHANE: SAOS-2 EFFECTS

REFERENCES

1. Bauer, T.W., and Muschler, G.F. Bone graft materials: Anoverview of the basic science. Clin. Orthop. 371, 10, 2000.

2. Hernigou, P., Gras, G., Marinello, G., and Dormont, D. In-activation of HIV by application of heat and radiation: Im-plication in bone banking with irradiated allograft bone.Acta Orthop. Scand. 71, 508, 2000.

3. Yuasa, T., Miyamoto, Y., Ishikawa, K., Takechi, M., Na-gayama, M., and Suzuki, K. In vitro resorption of three ap-atite cements with osteoclasts. J. Biomed. Mater. Res. 54,344, 2001.

4. Ishaug-Riley, S.L., Crane, G.M., Gurlek, A., Miller, M.J.,Yasko, A.W., Yaszemski, M.J., and Mikos, A.G. Ectopicbone formation by marrow stromal osteoblast transplanta-tion using poly(DL-lactic-co-glycolic acid) foams implantedinto the rat mesentery. J. Biomed. Mater. Res. 36, 1, 1997.

5. Ishaug-Riley, S.L., Crane-Kruger, G.M., Yaszemski, M.J.,and Mikos, A.G. Three-dimensional culture of rat calvar-ial osteoblasts in porous biodegradable polymers. Bioma-terials 19, 1405, 1998.

6. Gorna, K., and Gogolewski, S. Biodegradablepolyurethanes for implants. II. In vitro degradation and cal-cification of materials from poly(�-caprolactone)-poly(eth-ylene oxide) diols and various chain extenders. J. Biomed.Mater. Res. 60, 592, 2002.

7. Gorna, K., and Gogolewski, S. Preparation, degradation,and calcification of biodegradable polyurethane foams forbone graft substitutes. J. Biomed. Mater. Res. 67A, 813,2003.

8. Devin, J.E., Attawia, M.A., and Laurencin, C.T. Three-di-mensional degradable porous polymer–ceramic matricesfor use in bone repair. J. Biomater. Sci. Polym. Ed. 7, 661,1996.

9. Gomes, M.E., Sikavitsas, V.I., Behravesh, E., Reis, R.L.,and Mikos, A.G. Effect of flow perfusion on the osteogenicdifferentiation of bone marrow stromal cells cultured onstarch-based three-dimensional scaffolds. J. Biomed.Mater. Res. 67A, 87, 2003.

10. Suganuma, J., and Alexander, H. Biological response of in-tramedullary bone to poly-L-lactic acid. J. Appl. Biomater.Res. 22, 1071, 1988.

11. Vert, M., Mauduit, J., and Li, S. Biodegradation ofPLA/GA polymers: Increasing complexity. Biomaterials15, 1209, 1994.

12. Mathur, A.B., Collier, T.O., Kao, W.J., Wiggins, M., Schu-bert, M.A., Hiltner, A., and Anderson, J.M. In vivo bio-compatibility and biostability of modified polyurethanes. J.Biomed. Mater. Res. 36, 246, 1997.

13. Tanzi, M.C., Farè, S., Petrini, P., Tanini, A., Piscitelli, E.,Zecchi-Orlandini, S., and Brandi, M.L. Cytocompatibilityof polyurethane foams as biointegrable matrices for thepreparation of scaffolds for bone reconstruction. J. Appl.Biomater. Biomech. 1, 58, 2003.

14. Laurencin, C., Ambrosio, A., Borden, M., and Cooper, J.Tissue engineering: Orthopedic applications. Annu. Rev.Biomed. Eng. 1, 19, 1999.

15. Stock, U.A., and Vacanti, J.P. Tissue engineering: Currentstate and prospects. Annu. Rev. Med. 52, 443, 2001.

697

16. Ratner, B.D., and Bryant, S.J. Biomaterials: Where we havebeen and where we are going. Annu. Rev. Biomed. Eng.6, 41, 2004.

17. Sikavitsas, V.I., van den Dolder, J., Bancroft, G.N., Jansen,J.A., and Mikos, A.G. Influence of the in vitro culture pe-riod on the in vivo performance of cell/titanium bone tis-sue-engineered constructs using a rat cranial critical sizedefect model. J. Biomed. Mater. Res. 67A, 944, 2003.

18. Ishaug, S.L., Crane, G.M., Miller, M.J., Yasko, A.W.,Yaszemski, M.J., and Mikos, A.G. Bone formation by three-dimensional stromal osteoblast culture in biodegradable poly-mer scaffolds. J. Biomed. Mater. Res. 36, 17, 1997.

19. Freed, L.E., and Vunjak-Novakovic, G. Tissue engineeringbioreactors. In: Lanza, R.P., Langer, R., and Vacanti, J.P.,eds. Principles of Tissue Engineering. San Diego, CA: Aca-demic Press, 2000, pp. 143–156.

20. Sikavitsas, V.I., Bancroft, G.N., and Mikos, A.G. Forma-tion of three-dimensional cell/polymer constructs for bonetissue engineering in a spinner flask and a rotating wall ves-sel bioreactor. J. Biomed. Mater. Res. 62, 136, 2002.

21. Botchwey, E.A., Pollack, S.R., Levine, E.M., and Lau-rencin, C.T. Bone tissue engineering in a rotating bioreac-tor using a microcarrier matrix system. J. Biomed. Mater.Res. 55, 242, 2001.

22. Botchwey, E.A., Dupree, M.A., Pollack, S.R., Levine,E.M., and Laurencin, C.T. Tissue engineered bone: Mea-surement of nutrient transport in three-dimensional matri-ces. J. Biomed. Mater. Res. 67A, 357, 2003.

23. Botchwey, E.A., Pollack, S.R., El-Amin, S., Levine, E.M.,Tuan, R.S., and Laurencin, C.T. Human osteoblast-likecells in three-dimensional culture with fluid flow. Biorhe-ology 40, 299, 2003.

24. Yu, X., Botchwey, E.A., Levine, E.M., Pollack, S.R., andLaurencin, C.T. Bioreactor-based bone tissue engineering:The influence of dynamic flow on osteoblast phenotypicexpression and matrix mineralization. Proc. Natl. Acad.Sci. U.S.A. 101, 11203, 2004.

25. Goldstein, A.S., Juarez, T.M., Helmke, C.D., Gustin, M.C.,and Mikos, A.G. Effect of convection on osteoblastic cellgrowth and function in biodegradable polymer foam scaf-folds. Biomaterials 22, 1279, 2001.

26. Bancroft, G.N., Sikavitsas, V.I., van den Dolder, J.,Sheffield, T.L., Ambrose, C.G., Jansen, J.A., and Mikos,A.G. Fluid flow increases mineralized matrix deposition in3D perfusion culture of marrow stromal osteoblasts in adose-dependent manner. Proc. Natl. Acad. Sci. U.S.A. 99,12600, 2002.

27. Bancroft, G.N., Sikavitsas, V.I., and Mikos, A.G. Designof a flow perfusion bioreactor system for bone tissue-en-gineering applications. Tissue Eng. 9, 549, 2003.

28. van den Dolder, J., Bancroft, G.N., Sikavitsas, V.I.,Spauwen, P.H., Ambrose, C.G., Jansen, J.A., and Mikos,A.G. Flow perfusion culture of marrow stromal osteoblastsin titanium fiber mesh. J. Biomed. Mater. Res. 64A, 235,2003.

29. Sikavitsas, V.I., Bancroft, G.N., Holtorf, H.L., Jansen, J.A.,and Mikos, A.G. Mineralized matrix deposition by marrowstromal osteoblasts in 3D perfusion culture increases withincreasing fluid shear forces. Proc. Natl. Acad. Sci. U.S.A.100, 14683, 2003.

FASSINA ET AL.

30. Ingber, D.E. Tensegrity: The architectural basis of cellularmechanotransduction. Annu. Rev. Physiol. 59, 575, 1997.

31. Burger, E.H., and Klein-Nulend, J. Mechanotransductionin bone: Role of the lacuno-canalicular network. FASEBJ. 13(Suppl.), S101, 1999.

32. Weinbaum, S., Cowin, S.C., and Zeng, Y. A model for theexcitation of osteocytes by mechanical loading-inducedbone fluid shear stresses. J. Biomech. 27, 339, 1994.

33. Klein-Nulend, J., van der Plas, A., Semeins, C., Ajubi,N.E., Frangos, J.A., Nijweide, P.J., and Burger, E.H. Sen-sitivity of osteocytes to biomechanical stress in vitro.FASEB J. 9, 441, 1995.

34. Smalt, R., Mitchell, F.T., Howard, R.L., and Chambers,T.J. Induction of NO and prostaglandin E2 in osteoblastsby wall-shear stress but not mechanical strain. Am. J. Phys-iol. 273, E751, 1997.

35. Ajubi, N.E., Klein-Nulend, J., Alblas, M.J., Burger, E.H.,and Nijweide, P.J. Signal transduction pathways in-volved in fluid flow-induced PGE2 production by cul-tured osteocytes. Am. J. Physiol. Endocrinol. Metab. 276,E171, 1999.

36. Bakker, A.D., Soejima, K., Klein-Nulend, J., and Burger,E.H. The production of nitric oxide and prostaglandin E2

by primary bone cells is shear stress dependent. J. Biomech.34, 671, 2001.

37. Bidwell, J.P., Alvarez, M., Feister, H., Onyia, J., and Hock,J. Nuclear matrix proteins and osteoblast gene expression.J. Bone Miner. Res. 13, 155, 1998.

38. Pavalko, F.M., Norvell, S.M., Burr, D.B., Turner, C.H.,Duncan, R.L., and Bidwell, J.P. A model for mechano-transduction in bone cells: The load-bearing mechano-somes. J. Cell. Biochem. 88, 104, 2003.

39. Chen, N.X., Ryder, K.D., Pavalko, F.M., Turner, C.H.,Burr, D.B., Qiu, J., and Randall, L.D. Ca2� regulates fluidshear-induced cytoskeletal reorganization and gene ex-pression in osteoblasts. Am. J. Physiol. Cell Physiol. 278,C989, 2000.

40. Anderson, H.C., Reynolds, P.R., Hsu, H.H., Missana, L.,Masuhara, K., Moylan, P.E., and Roach, H.I. Selective syn-thesis of bone morphogenetic proteins-1, -3, -4 and bonesialoprotein may be important for osteoinduction by Saos-2 cells. J. Bone Miner. Metab. 20, 73, 2002.

41. Belanger, M.C., Marois, Y., Roy, R., Mehri, Y., Wagner,E., Zhang, Z., King, M.W., Yang, M., Hahn, C., andGuidoin, R. Selection of a polyurethane membrane for themanufacture of ventricles for a totally implantable artifi-cial heart: Blood compatibility and biocompatibility stud-ies. Artif. Organs 24, 879, 2000.

42. Vogel, K.G., and Evanko, S.P. Proteoglycans of fetalbovine tendon. J. Biol. Chem. 262, 13607, 1987.

43. Bradford, M.M. A rapid and sensitive method for the quan-titation of microgram quantities of protein utilizing theprinciple of protein-dye binding. Anal. Biochem. 72, 248,1976.

44. Laemmli, U.K. Cleavage of structural proteins during theassembly of the head of bacteriophage T4. Nature 227, 680,1970.

45. Yamagata, T., Saito, H., Habuchi, O., and Suzuki, S. Pu-rification and properties of bacterial chondroitinases andchondrosulfatases. J. Biol. Chem. 243, 1523, 1968.

698

46. Yoshida, K., Miyauchi, S., Kikuchi, H., Tawada, A., andTokuyasu, K. Analysis of unsaturated disaccharides fromglycosaminoglycuronan by high-performance liquid chro-matography. Anal. Biochem. 177, 327, 1989.

47. Fisher, L.W., Stubbs, J.T., III, and Young, M.F. Antiseraand cDNA probes to human and certain animal model bonematrix noncollagenous proteins. Acta Orthop. Scand.266(Suppl.), 61, 1995.

48. Bernstein, E.F., Chen, Y.Q., Kopp, J.B., Fisher, L.W.,Brown, D.B., Hahn, P.J., Robey, F.A., Lakkakorpi, J., andUitto, J. Long-term sun exposure alters the collagen of thepapillary dermis: Comparison of sun-protected and pho-toaged skin by Northern analysis immunohistochemicalstaining, and confocal laser scanning microscopy, J. Am.Acad. Dermatol. 34, 209, 1996.

49. Saad, B., Ciardelli, G., Matter, S., Welti, M., Uhlschmid,G.K., Neuenschwander, P., and Suter, U.W. Degradableand highly porous polyesterurethane foam as biomaterial:Effects and phagocytosis of degradation products in os-teoblasts. J. Biomed. Mater. Res. 39, 594, 1998.

50. Ma, P.X., Zhang, R., Xiao, G., and Franceschi, R. Engi-neering new bone tissue in vitro on highly porous poly(�-hydroxyl acids)/hydroxyapatite composite scaffolds. J.Biomed. Mater. Res. 54, 284, 2001.

51. Farè, S., Petrini, P., Tanzi, M.C., Bigi, A., and Roveri, N.Biointegrable 3D polyurethane/�-TCP composites for bonereconstruction. In: Mantovani, D., ed. Advanced Materialsfor Biomedical Application. Montreal, Canada: CanadianInstitute of Mining, Metallurgy and Petroleum, 2002, pp.17–26.

52. Farè, S., Petrini, P., and Tanzi, M.C. 3D polyurethane/�-TCP composite scaffolds for bone tissue engineering. Ab-stract presented at the IEEE-EMBS Special Topic Confer-ence on Cellular and Tissue Engineering, Genoa, Italy,2002. Abstract no. 40.

53. Anderson, H.C., Sugamoto, K., Morris, D.C., Hsu, H.H.,and Hunt, T. Bone-inducing agent (BIA) from cultured hu-man SAOS-2 osteosarcoma cells. Bone Miner. 16, 49,1992.

54. Anderson, H.C., Hsu, H.H., Raval, P., Hunt, T.R., Schwap-pach, J.R., Morris, D.C., and Schneider, D.J. The mecha-nism of bone induction and bone healing by human os-teosarcoma cell extracts. Clin. Orthop. 313, 129, 1995.

55. Anderson, H.C., Gurley, D.J., Hsu, H.H., Aguilera, X.M.,Davis, L.S., and Moylan, P.E. Secretion of a bone-induc-ing agent (BIA) by cultured Saos-2 human osteosarcomacells. J. Musculoskel. Res. 3, 39, 1999.

56. Laitinen, M., Jortikka, L., Halttunen, T., Bohling, T., Mart-tinen, A., and Lindholm, T.S. Soluble factors from humanSaos-2 osteosarcoma cells induce ectopic bone formationand osteoblastic differentiation of cultured mesenchymalcells. J. Musculoskel. Res. 1, 21, 1997.

57. Kasugai, S., Nagata, T., and Sodek, J. Temporal studies onthe tissue compartmentalization of bone sialoprotein (BSP),osteopontin (OPN), and SPARC protein during bone for-mation in vitro. J. Cell. Physiol. 152, 467, 1992.

58. Luan, Y.J., Praul, C.A., and Gay, C.V. Confocal imagingand timing of secretion of matrix proteins by osteoblastsderived from avian long bone. Comp. Biochem. Physiol. A126, 213, 2000.

FLOW PERFUSION IN POLYURETHANE: SAOS-2 EFFECTS

59. Butler, W.T. Structural and functional domain of osteo-pontin. Ann. N.Y. Acad. Sci. 760, 6, 1995.

60. Gompertz, B. On the nature of the function expressive ofthe law of human mortality and on a new method of de-termining the value of life contingencies. Philos. Trans. R.Soc. 513, 1825.

61. Emmans, G.C. A method to predict the food intake of do-mestic animals from birth to maturity as a function of time.J. Theor. Biol. 186, 189, 1997.

62. Hauschka, P.V., Lian, J.B., Cole, D.E.C., and Gundberg,C.M. Osteocalcin and matrix Gla protein: Vitamin K-de-pendent proteins in bone. Physiol. Rev. 69, 990, 1989.

63. Garnero, P., Grimaux, M., Seguin, P., and Delmas, P.D.Characterization of immunoreactive forms of human os-teocalcin in vivo and in vitro. J. Bone Miner. Res. 9, 255,1994.

64. Calvo, M.S., Eyre, D.R., and Gundberg, C.M. Molecularbasis and clinical application of biological markers of boneturnover. Endocr. Rev. 17, 333, 1996.

65. Tenni, R., Viola, M., Welser, F., Sini, P., Giudici, C., Rossi,A., and Tira, M.E. Interaction of decorin with CNBr pep-tides from collagens I and II. Evidence for multiple bind-ing sites and essential lysyl residues in collagen. Eur. J.Biochem. 269, 1428, 2002.

66. Hocking, A.M., Shinomura, T., and McQuillan, D.J.Leucine-rich repeat glycoproteins of the extracellular ma-trix. Matrix Biol. 17, 1, 1998.

67. Iozzo, R.V. Matrix proteoglycans: From molecular designto cellular function. Annu. Rev. Biochem. 67, 609, 1998.

68. Vogel, K.G., Paulsson, M., and Heinegard, D. Specific in-hibition of type I and type II collagen fibrillogenesis by thesmall proteoglycan of tendon. Biochem. J. 223, 587, 1984.

69. Sini, P., Denti, A., Tira, M.E., and Balduini, C. Role ofdecorin on in vitro fibrillogenesis of type I collagen. Gly-coconj. J. 14, 871, 1997.

70. Yamaguchi, Y., Mann, D.M., and Ruoslahti, E. Negativeregulation of transforming growth factor-� by the proteo-glycan decorin. Nature 346, 281, 1990.

71. Csordas, G., Santra, M., Reed, C.C., Eichstetter, I., Mc-Quillan, D.J., Gross, D., Nugent, M.A., Hajnoczky, G., andIozzo, R.V. Sustained down-regulation of the epidermalgrowth factor receptor by decorin: A mechanism for con-trolling tumor growth in vivo. J. Biol. Chem. 275, 32879,2000.

72. Martins-Gren, M. Dynamics of cell–ECM interaction. In:Lanza, R.P., Langer, R., and Vacanti, J.P., eds. Principlesof Tissue Engineering. San Diego, CA: Academic Press,2000, pp. 33–55.

73. Saltzman, W.M. Cell interactions with polymers. In: Lanza,R.P., Langer, R., and Vacanti, J.P., eds. Principles of Tis-sue Engineering. San Diego, CA: Academic Press, 2000,pp. 221–235.

74. El-Ghannam, A., Ducheyne, P., and Shapiro, I.M. Effectof serum proteins on osteoblast adhesion to surface-modi-fied bioactive glass and hydroxyapatite. J. Orthop. Res. 17,340, 1999.

75. Kilpadi, K.L., Chang, P.L., and Bellis, S.L. Hydroxylap-atite binds more serum proteins, purified integrins, and os-teoblast precursor cells than titanium or steel. J. Biomed.Mater. Res. 57, 258, 2001.

699

76. Yang, Y., Cavin, R., and Ong, J.L. Protein adsorption ontitanium surfaces and their effect on osteoblast attachment.J. Biomed. Mater. Res. 67A, 344, 2003.

77. Webster, T.J., Ergun, C., Doremus, R.H., Siegel, R.W., andBizios, R. Specific proteins mediate enhanced osteoblastadhesion on nanophase ceramics. J. Biomed. Mater. Res.51, 475, 2000.

78. Webster, T.J., Schadler, L.S., Siegel, R.W., and Bizios, R.Mechanisms of enhanced osteoblast adhesion on nanophasealumina involve vitronectin. Tissue Eng. 7, 291, 2001.

79. MacDonald, D.E., Deo, N., Markovic, B., Stranick, M., andSomasundaran, P. Adsorption and dissolution behavior ofhuman plasma fibronectin on thermally and chemicallymodified titanium dioxide particles. Biomaterials 23, 1269,2002.

80. Geissler, U., Hempel, U., Wolf, C., Scharnweber, D.,Worch, H., and Wenzel, K. Collagen type I-coating ofTi6Al4V promotes adhesion of osteoblasts. J. Biomed.Mater. Res. 51, 752, 2000.

FASSINA ET AL.

81. Bellahcène, A., Bonjean, K., Fohr, B., Fedarko, N.S.,Robey, F.A., Young, M.F., Fisher, L.W., and Castronovo,V. Bone sialoprotein mediates human endothelial cell at-tachment and migration and promotes angiogenesis. Circ.Res. 86, 885, 2000.

82. Ingber, D.E., and Folkman, J. Mechanochemical switchingbetween growth and differentiation during fibroblastgrowth factor-stimulated angiogenesis in vitro: Role of ex-tracellular matrix. J. Cell Biol. 109, 317, 1989.

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