Modulation of Spreading, Proliferation, and Differentiation of Human

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Modulation of Spreading, Proliferation, and Differentiation of Human Mesenchymal Stem Cells on Gelatin-Immobilized Poly(L-lactide-co-E-caprolactone) Substrates Young Min Shin,† Kyung-Soo Kim,‡ Youn Mook Lim,§ Young Chang Nho,§ and Heungsoo Shin*,† Department of Bioengineering and Cardiology Division, Department of Internal Medicine, Hanyang University, 17 Haengdang-dong, Seongdong-gu, Seoul 133-791, Korea, and Radiation Research Center for Industry and Environment, Advanced Radiation Technology Institute, Korea Atomic Energy Research Institute, 1266 Sinjeong-dong, Jeongeup-si, Jeollabuk-do, 580-185, Korea Received December 21, 2007; Revised Manuscript Received April 14, 2008

Controlled adhesion and continuous growth of human mesenchymal stem cells (hMSCs) are essential for scaffoldbased delivery of hMSCs in tissue engineering applications. The main goal of this study is to develop biofunctionalized synthetic substrates to actively control adhesion, spreading, and proliferation of hMSCs. γ-Ray irradiation was employed to graft acrylic acid (AAc) to biodegeradable poly(L-lactide-co--caprolactone) (PLCL) films. Gelatin, a natural polymer, was then immobilized on the AAc grafted PLCL film (AAc-PLCL) to induce biomimetic interactions with the cells. The graft yield of AAc increased as the irradiation dose and AAc concentration increased, and the presence of gelatin (gelatin-AAc-PLCL) following immobilization was confirmed using ESCA. To investigate cell responses, hMSCs isolated from a human mandible were cultured on the various substrates and their adhesion, spreading, and proliferation were examined. After three days of culture, the DNA concentration from the cells cultured on gelatin-AAc-PLCL film was 2.9-fold greater than that on the PLCL film. Immunofluorescent staining of hMSCs cultured on the gelatin-AAc-PLCL films demonstrated homogeneous localization of F-Actin and vinculin in their cytoplasm, while mature adhesive structure was not observed from the cells cultured on other substrates. Furthermore, the ratio of projected area of adherent single cells on gelatinAAc-PLCL films was significantly larger (116.80 ( 12.78%) than that on the PLCL films (30.11 ( 5.07%). Our results suggest that gelatin-immobilized PLCL substrates may be potentially used in tissue engineering, particularly as a stem cell delivery carrier for the regeneration of target tissue.

1. Introduction A biologically functional scaffold is an essential element for the successful regeneration of damaged tissue.1–3 It serves as a physical support for cell attachment and subsequent stimulant of cellular activities such as migration, proliferation, and differentiation.4,5 Additionally, scaffold materials should be degraded upon implantation and possess appropriate mechanical stability to support tissue structure while scaffold/cell composites undergo the tissue regeneration processes. In particular, the mechanical compliance of scaffolds to tissue and the extracellular matrix (ECM) environment is important such that many cell types including mesenchymal stem cells (MSCs) demonstrate extreme sensitivity to tissue level elasticity.6,7 We recently reported mechanical properties of elastic biodegradable poly(Llactide-co--caprolactone) (PLCL), demonstrating that the contractile phenotypes and ECM protein expression of smooth muscle cells cultured on PLCL scaffolds was affected by mechanical stimulation.8,9 Unlike other relatively rigid poly(Rhydroxy ester)s, such as poly(L-lactide) and poly(caprolactone), the elastic nature of PLCL attributed to maintaining mechanical stability under cyclic loading in vitro, which may be potentially used as a scaffold material for engineering elastic soft tissue * To whom correspondence should be addressed. Tel.: +82-2-2220-2346. Fax: +82-2-2298-2346. E-mail: [email protected]. † Department of Bioengineering, Hanyang University. ‡ Department of Internal Medicine, Hanyang University. § Korea Atomic Energy Research Institute.

such as muscle, skin, and blood vessels. However, despite its excellent mechanical properties and biocompatibility, the hydrophobicity of PLCL often lead to poor cell affinity and limited cell proliferation, and the lack of biofunctionality hindered active control of cellular responses.10 Many cell types in tissue engineering are anchorage-dependent and are required to form a tight mechanical binding to scaffolds as an initial step in their cellular-level interactions. Cell adhesion to biomaterial surfaces is mainly mediated by cell membrane integrin receptors and their binding to extracellular matrix (ECM) molecules.11 Therefore, adhesive ECM proteins, such as fibronectin, laminin, and their functional putative peptides, such as Arg-Gly-Asp (RGD) and Tyr-Ile-GlySer-Arg (YIGSR), were immobilized to the surface of scaffolds to control cell adhesion.12–15 Alternative bioactive molecules for surface immobilization have been considered for this function and include natural polymers such as collagen, growth factors, and gelatin: these molecules have demonstrated favorable interactions with adherent cells by modulating cell adhesion and spreading.16–18 Among these natural polymers, gelatin (the denatured form of collagen) has been widely used for the surface modification of scaffolds due to its cytocompatility and relatively low cost of production.19 Because most poly(R-hydroxy ester)s, including PLCL, have limited reactive functional groups, the surface modification is carried out by two steps; the reactive linkers are first added or exposed by chemical reactions, and then desirable biofunctional moieties are introduced. One of these processes includes surface

10.1021/bm701410g CCC: $40.75  2008 American Chemical Society Published on Web 06/18/2008

Human Mesenchymal Stem Cells

etching that involves a base treatment to cleave some of the surface ester bonds to generate carboxyl end groups.10,19 However, this process may not be useful in most tissue engineering applications under consistent mechanical loading due to the weakening of the mechanical properties as a consequence of partial degradation of the main chain. Plasma treatment has been the most widely used to change the surface chemistry by introducing various functional groups or in situ graft polymerization.20–23 As compared to plasma treatment and surface etching, γ-ray irradiation was recently utilized in biomedical fields, which has several advantages; it is a relatively simple and clean process without the requirement of chemical additives or catalysts for graft polymerization on various polymer surfaces.24 In addition, there is greater control over graft yield and a larger penetration depth than plasma treatment. These methods can be easily obtained by changing the irradiation dose.25 The ultimate goal of this study was to develop a biofunctionalized degradable polymer substrate that is able to actively control cell responses, with a particular interest in stem cells. A two-step modification process was used to functionalize PLCL substrates to modulate cell adhesion, proliferation, and spreading; acrylic acid (AAc), used as a spacer, was grafted directly to the surface of the PLCL film by γ-ray irradiation, and gelatin was subsequently immobilized by chemical coupling to the carboxylic acid groups in the grafted AAc. The effect of irradiation dose on the grafted amount of AAc on the PLCL film was investigated to optimize the condition for gelatin conjugation. The morphology of the adhesive structure, the projected area of adherent cells, and cell proliferation on the gelatin-grafted PLCL film were then examined using a model cell type (human mesenchymal stem cells (hMSCs) derived from mandibles).

2. Materials and Methods 2.1. Materials. L-Lactide was obtained from Purac Biochem (Gorinchem, Netherlands). -Caprolactone and N-hydroxysuccinimide (NHS) were purchased from Aldrich (Milwaukee, WI). 1,6-Hexanediol, calcium hydride, toluene, stannous octoate, ammonium iron(II) sulfate hexahydrate, toluidine blue, N-(3-dimethylamionipropyl)-N′-ethylcarbodiimide hydrochloride (EDC), acrylic acid (AAc), 4-morphilinoethanesulfonic acid (MES), fluorescamine, and gelatin type B (from bovine skin) were purchased from Sigma (St. Louis, MO). Dulbecco’s Modified Eagle Medium (DMEM) with low glucose, penicillinstreptomycin (PS), and phosphate buffered saline (PBS, pH 7.4) were obtained from Gibco (Grand Island, NY). Fetal bovine serum (FBS) was purchased from United Search Partners (Austin, TX). Ultrapure water was used after purification with Milli-Q Plus System (Millipore, MA). All other chemicals and solvents were of analytical grade and used without further purification. 2.2. Synthesis of Poly(L-lactide-co-E-caprolactone) (PLCL) and Preparation of PLCL Film. The synthesis of poly(L-lactide-co-carpolactone) (PLCL) was performed by ring opening polymerization as previously described.8 Briefly, the polymerization of PLCL was carried out in a glass ampule containing L-lactide (LA) and -caprolactone (CL) dissolved in 1,6-hexanediol (LA/CL ) 50:50, molar ratio) at 150 °C for 24 h with mechanical stirring. After the reaction, the polymer was precipitated by dropping into an excess of methanol and the precipitant was dried at 50 °C under vacuum for 72 h. To prepare PLCL films, PLCL was dissolved in chloroform (10 wt %) and poured onto a glass plate walled with aluminum tape and covered with aluminum foil. To evaporate the solvent, the glass plate was placed in an ambient condition for 24 h and thoroughly dried for 72 h in a vacuum oven. The fully dried PLCL film was removed from the glass plate by wetting with 70% ethanol and then redrying it in a dry oven prior to

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surface modification. The fabricated PLCL film was cut into a square (10 mm × 10 mm) and punched into a circle (20 mm diameter) for further experiments. 2.3. AAc Graft onto the PLCL Films Using γ-Ray Irradiation. AAc was grafted onto PLCL films using γ-ray irradiation with cobalt 60.24 The PLCL films were wetted with 70% ethanol, immersed in aqueous AAc solutions (5 mL) in concentrations of 3, 5, and 10 wt%, and then exposed to various radiation doses ranging from 5 to 15 kGy at ambient temperature. After irradiation, unreacted monomers and homopolymers were removed by washing then with an excess amount of distilled water (DW) by stirring for 12 h. The grafting yield (%) was determined by the following equation:

grafting yield(%) )

W1 - W0 × 100 W0

(1)

where W1 and W0 are the weight of the films after and before irradiation, respectively. To examine swelling behavior, the samples (10 mm × 10 mm) were incubated in DW for 24 h, and the change in volume was measured. The swelling ratio was calculated by the following equation:

swelling ratio(%) )

V1 - V0 × 100 V0

(2)

where V1 and V0 are the volume of the films after and before swelling in DW. The surface functional groups of the AAc grafted PLCL films (AAc-PLCL) were examined by attenuated total reflection Fourier transform infrared (ATR-FTIR) spectroscopy (Tensor 37, Bruker optics, MA). To evaluate the irradiation-induced degradation of the PLCL substrates, we measured molecular weights of the PLCL before and after irradiation using a gel permeation chromatography (PL-GPC 220, Polymer Laboratories Ltd., Church Stretton, U.K.) equipped with microstyragel column calibrated with polystyrene. THF was used as a mobile phase at a flow rate of 1.0 mL/min at 40 °C. We also tested the mechanical properties of the grafted PLCL substrates. A universal testing machine (Shimadzu, Japan) was used to determine the tensile strength of the samples (30 mm × 5 mm × 0.06 mm) with a 5 N load cell. The films were mounted on gripping units, and mechanical load was applied at a constant deformation rate of 10 mm/min. 2.4. Quantification of the Grafted AAc. To quantify the amount the grafted carboxylic acid groups, we stained the films with toluidine blue as previously described.25 Briefly, the AAc-PLCL films were incubated in toluidine blue O solution (0.1 M HCl, 20 mg NaCl, and 4 mg toluidine blue O chloride) for 4 h at room temperature. After washing with DW, the films were resolved with 0.1 M NaOH and ethanol solution (1: 4 v/v) until complete decolorization, and then the uptake amounts of toluidine blue O were quantified by measuring the absorbance at 530 nm using a plate reader (Spectra Max M2e, Molecular Devices, CA). The standard calibration curve was obtained by using known concentrations of toluidine blue O solutions.26 2.5. Immobilization of Gelatin onto the AAc-PLCL Films. Gelatin was immobilized onto the AAc-PLCL films (5 wt % AAc, 10 kGy) using common EDC/NHS reaction chemistry. The AAc-PLCL films were prewetted with 70% ethanol, washed with DW, and immersed in 0.1 M 2-(N-morpholino)ethanesulfonic acid (MES) buffer solution (pH 5.07). The films were then reacted with EDC/NHS in MES buffer solution (5 mg/mL) for 1 h at room temperature. Following the reaction, the reaction buffer solution was aspirated, and the films were rewashed to remove any unreacted residues. The gelatin solutions were prepared in the same buffer solution with 0.2, 2, and 4 mg/mL concentrations and reacted with the activated AAc-PLCL films for 12 h at room temperature. Following the reaction, unreacted gelatin was washed away with free MES buffer solution and PBS (pH 7.4). Electron spectroscopy for chemical analysis (ESCALAB, Thermo Scientific, MA) was performed to confirm the presence of gelatin on the surface of AAcPLCL films. 2.6. Quantification of Immobilized Gelatin on the AAc-PLCL Films. After immobilization of gelatin, the amount of gelatin on the AAc-PLCL films were quantified using fluorescamine assay as previ-

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Figure 1. Schematic diagram of the biofuctionalization of PLCL film using γ-ray irradiation.

ously described.27 Briefly, AAc-PLCL substrates (5 wt% - 10 kGy) were activated with EDC/NHS (5 mg/mL), and reacted with 0.2 and 2 mg/mL gelatin solutions for 12 h. For the passive adsorption of gelatin, the AAc-PLCL films were treated with 2 mg/mL gelatin solution for the same reaction time without EDC/NHS activation. All films were immersed in 200 µL of borate buffer (pH 9.2), and 50 µL of fluorescamine solution (4 mg/mL in DMSO) was then added. After rapid vortexing for 60 s, fluorescence intensity was measured using the plate reader (390 nm of excitation and 475 nm of emission wavelength). Standard calibration curve was obtained from the intensity of known concentrations of gelatin (0-1000 µg/mL). 2.7. Human Mesenchymal Stem Cells (hMSCs) Culture. hMSCs, isolated from a mandibles, were provided by the Asan Medical Center (Seoul, Korea). These cells were cultured in low glucose DMEM supplemented with 10% fetal bovine serum and 1% penicillin/ streptomycin in a CO2 incubator at 37 °C with 5% CO2, 95% humidity, and the medium was changed every two days. In this experiment, the cells at the sixth passage were used. 2.8. Cell Adhesion and Proliferation Analysis. To evaluate cell adhesion and proliferation on gelatin-immobilized AAc-grafted PLCL (gelatin-AAc-PLCL) film, DNA from cells cultured on the films were quantified using a Picogreen dsDNA assay kit (Molecular probes, CA). Picogreen assay has been well established in characterizing cell proliferation cultured under 2-D and 3-D culture conditions as previously published.28 Suspended hMSCs were seeded on TCPS (Tissue Culture Plates, Corning, NY; positive control), PLCL, AAc-PLCL, and gelatin-AAc-PLCL films with 6.3 × 103 cells/cm2 and cultured for 8, 24, and 72 h. After the cell culture, the films were gently washed with PBS, and DNA was extracted from the cells using the radio immuno precipitation assay (RIPA) lysis buffer (150 mM NaCl, 1% Triton X-100, 1% sodium deoxycholate, 0.1% SDS, 150 mM Tris, pH 7.2, and protease inhibitors) with vortexing. Following the reaction with the Picogreen reagent, the fluorescent intensity was observed using the plate reader at 480 nm of excitation and 520 nm of emission wavelength. To obtain a standard calibration curve, the intensity of known concentrations of λDNA (30-1000 ng/mL) was measured. 2.9. Immunofluorescent Staining. The cell nucleus, vinculin, and F-Actin were stained to evaluate the morphology of cells on the films. After 24 h of cell culture, the samples were fixed using 3.6% formaldehyde for 15 min at 37 °C and the cells were permeabilized with cytoskeletal buffer solution (10.3 g sucrose, 0.292 g NaCl, 0.06 g MgCl2, 0.476 g HEPES buffer, 0.5 mL Triton X-100, in 100 mL water, pH 7.2) for 5 min at 4 °C. For the staining, the cells were first blocked with 1% BSA for 30 min at 37 °C, then incubated with antivinculin (Upstate, Charlottesville, VA) for 1 h at 37 °C, and subsequently reincubated with Alexa-fluoro 488 rabbit antimouse IgG (Molecular probes, Eugene, OR), Rhodamine-phalloidin (Molecular Probes, Eugene, OR), and Hoechst 33258 (Molecular Probes, Eugene, OR) for 1 h at 37 °C. After the samples were mounted on glass slides, immunofluorescent images were obtained using a fluorescent microscope (TE 2000, Nikon, Japan) and a laser scanning confocal microscope (LSM 510, Zeiss, Germany). The projected cell area from the obtained images was analyzed using Imagepro Plus 4.5 (Media cybernetics, MD). Cell spreading images were randomly chosen from six fields of one slide. The total area and the cell-covered area were

measured, and the number of cell in each image was counted using imaging software. The projected cell area was expressed as ratio of the covered area to the total area. 2.10. Alkaline Phosphatase Activity. The ability of the gelatinAAc-PLCL substrate to control the differentiation of hMSCs was evaluated by measuring alkaline phosphatase (ALP) activity, which is one of the markers related to the osteogenic differentiation of hMSCs. hMSCs were seeded on TCPS, PLCL, AAc-PLCL, and gelatin-AAcPLCL films at 3.1 × 104 cells/cm2 and cultured in growth media for 24 h. Following culture for 24 h, the growth media was replaced with the osteogenic differentiation media supplemented with 2.84 × 10-4 M L-ascorbic acid and 1 × 10-2 M β-glycerol-phosphate and cultured for 7 days. After lysis of hMSCs with RIPA lysis buffer, lysates were reacted with ALP solution for ELISA (Sigma Aldrich, St. Louis, MO) for 30 min at 37 °C, and then the optical density was observed using the plate reader at 405 nm of wavelength. To obtain a standard calibration curve, the intensity of known concentrations of paranitrophenyl phosphate (0-600 µg/mL) was measured. Total DNA contents from each substrate were also quantified as described in cell adhesion assay using Picogreen dsDNA assay kit. 2.11. Statistical Analysis. All data were presented as means ( standard deviation (SD) for n ) 3. A Student’s t-test was used to assess statistical significance of the results (p < 0.05).

3. Results 3.1. AAc Graft onto the PLCL Films Using γ-Ray Irradiation. The initial step of surface modification (Figure 1) involved AAc grafting onto PLCL film in an ambient condition: 5 kGy of the γ-ray from cobalt 60 was irradiated for 1-3 h to the PLCL films immersed within the AAc solutions, at various concentrations (3, 5, and 10 wt %). Before AAc grafting, the average weight and thickness of the PLCL films (10 × 10 mm) were 49.7 ( 8 mg and 0.2 ( 0.03 mm, respectively. The graft yields of the AAc-PLCL films, reacted with 3 and 5 wt % AAc solution, were similar in all irradiation doses, as shown in Figure 2; the range of the graft yield was from 3.29 to 3.66% for 3 wt % AAc and from 6.71 to 8.33% for 5 wt % AAc, respectively. However, when the films were reacted with the 10 wt % AAc solution, the graft yield increased as a function of the irradiation dose. For example, when the irradiation dose was 5, 10, and 15 kGy, the corresponding graft yields were 18.1 ( 2.30, 22.05 ( 2.69, and 29.62 ( 1.58%. The swelling behavior of the AAcPLCL films was observed by the change in film volume after being wetted with DW for 24 h. The average volume of the films before swelling (width × length × thickness, 10.93 ( 0.39 mm × 7.11 ( 0.38 mm × 0.2 ( 0.03 mm) was 15.71 ( 2.34 mm3, as measured by a digital microcaliper. Consistent with the graft yield, swelling ratios were similar for the AAcPLCL films reacted with 3 and 5 wt % AAc, while they significantly increased for the film reacted with 10 wt % AAc (Figure 3). The swelling ratio for AAc-PLCL films reacted with 10 wt % AAc ranged from 25.85 ( 0.35 to 38.40 ( 3.36%. As



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Figure 2. Graft yields of AAc-PLCL films prepared under various irradiation doses and AAc concentrations. The symbol “/” indicates the significant difference (p < 0.01).

Figure 3. Swelling ratios of AAc-PLCL films as seen prepared under various irradiation doses and AAc concentrations. The symbol “/” indicates the significant difference (p < 0.01).

shown in Figure 4A, the surface characteristics of the AAcPLCL films were analyzed by ATR-FTIR, the carbonyl stretch peak from the ester groups in PLCL and the carboxylic acid groups from the AAc appeared at 1740 cm-1 and 1700 cm-1, respectively. After grafting of AAc, the carbonyl stretch peak from the ester groups in PLCL and acid carbonyl in AAc were overlapped and showed broaden peak shape at 1700 cm-1. When the reacted AAc concentration increased up to 10 wt %, the peaks were broadened due to the introduction of carboxyl acid groups (Figure 4B). Molecular weights of PLCL were determined using GPC before and after γ-ray irradiation. After expose to γ-ray of PLCL, molecular weight was decreased in irradiation dosedependent manner (Table 1). Molecular weight of nonirradiated PLCL was approximately 162000. PLCL exposed to 5, 10, and 15 kGy irradiation doses showed 144000, 118000, and 116000 of molecular weights, correspondingly. Although the γ-ray irradiation appeared to reduce the molecular weight of the polymer, tensile strength, and elongation of the substrates were similar irrespective of the γ-ray irradiation. For example, the tensile strength of AAc grafted PLCL (reacted with 5% AAc) was 2.26 ( 0.72, 2.43 ( 0.79, and 2.53 ( 0.75 by exposure to γ-ray irradiation at the dose of 5, 10, and 15 kGy, respectively. These values were not significantly different from the tensile

Figure 4. Surface chemical properties of the films by ATR-FT-IR. (A) Spectrum of PLCL, AAc, and AAc-PLCL and (B) peak shift of the AAc-PLCL film at 5 kGy. Table 1. Molecular Weights of PLCL Before and After γ-Ray Irradiation sample

Mn

Mw

PDI

nonirradiated PLCL 5 kGy 10 kGy 15 kGy

79993 74651 62420 57385

162073 144806 118217 116607

2.0261 1.9398 1.8939 2.0320

strength of PLCL without irradiation (2.27 ( 0.47). As shown in Table 2, the elongation of the samples showed the similar trend as observed in tensile properties. 3.2. Quantification of the Grafted Carboxylic Acid Groups. To quantify the introduced carboxylic acid groups, a toluidine blue O staining method was used. Toluidine blue O can bind to anionic ions, which has been used in determination of heparin concentration.26 The bound toluidine blue O can be visualized by microscopy or quantitatively analyzed using a UV-visible spectrophotometer. After reaction with toluidine blue O, the PLCL films with no AAc showed no indication of blue staining, while the other samples reacted with AAc were stained various levels of intensity (Figure 5A). The intensity of stained samples became more enhanced in an AAc concentration-dependent manner. The stained toluidine blue O was then completely released in a basic buffer solution, and the absorbance of the resulting solution was measured at 530 nm. Figure 5B shows that when AAc concentration increased, the uptake

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Table 2. Mechanical Properties of PLCL Films Before and After AAc Graft AAc concentration (wt %) 3

irradiation dose (kGy)

5 10 15 5 5 10 15 10 5 10 15 nonirradiated PLCL

tensile strength (MPa)

elongation (%)

2.16 ( 1.06 3.14 ( 0.98 2.61 ( 0.75 2.26 ( 0.72 2.43 ( 0.79 2.53 ( 0.75 2.16 ( 0.45 1.56 ( 0.09 2.70 ( 0.34 2.27 ( 0.47

381.66 ( 125.43 506.33 ( 119.10 420.67 ( 107.94 337.33 ( 119.32 415.00 ( 86.24 293.67 ( 230.17 318.00 ( 136.17 142.00 ( 5.56 402.67 ( 64.26 347.00 ( 32.53

amounts of toluidine blue O increased. The released amount of toluidine blue O from the AAc-PLCL films was approximately 8.68 ( 8.59, 16.83 ( 11.49, and 58.16 ( 15.84 µM/cm2 for the samples reacted with 3, 5, and 10 wt % AAc, respectively. 3.3. Immobilization of Gelatin onto the AAc-Grafted PLCL Films. To introduce biofunctionality to the AAc-PLCL films, gelatin type B, a thermally denatured form of collagen was used. The carboxylic acid groups immobilized to the PLCL film surface were activated using EDC/NHS, and the gelatin was then grafted by constituting amide bonds between the carboxylic acid group of AAc-PLCL film and amine groups of gelatin. When the presence of gelatin on the film was analyzed by ESCA, the new peak was observed only in gelatin-AAc-

Figure 6. Gelatin immobilization on the AAc-PLCL films: (A) ESCA spectrums and (B) the amount of immobilized gelatin of the various substrates.

Figure 5. Quantification of surface carboxylic acid groups: (A) toluidine blue-stained AAc-PLCL films and (B) uptake amount of toluidine blue in the AAc-PLCL films. The “/” indicates the significant difference (p < 0.01).

PLCL film at 400 ev, which was corresponding to the presence of nitrogen (Figure 6A). The amount of immobilized gelatin on the AAc-PLCL films was assessed by the fluorescamine assay. Gelatin-AAc-PLCL films reacted with various gelatin concentration (0.2 and 2 mg/mL), and passively absorbed gelatin-AAc-PLCL film reacted with 2 mg/mL gelatin solution showed different amount of immobilized gelatin. Gelatin-AAcPLCL films reacted with 2 mg/mL gelatin solution showed 2.02 ( 0.83 µg/cm2 of gelatin, which was approximately 7-fold increased relative to that from passive absorption group (0.29 ( 0.04 µg/cm2). In addition, the amount of gelatin from the gelatin-AAc-PLCL films reacted with 0.2 mg/mL gelatin solution was also 0.49 ( 0.20 µg/cm2 (Figure 6B). 3.4. Cell Adhesion and Proliferation on the GelatinAAc-PLCL Films. To assess cellular response, hMSCs were cultured on the substrates for up to 3 days and cell adhesion, spreading, and proliferation were studied. As shown in Figure 7A, the DNA concentrations for all the groups after 8 h of cell adhesion were similar (326.8 ( 76.1 ng/film, 342.72 ( 72.71 ng/film, and 407.2 ( 72.7 ng/film, for PLCL, AAc-PLCL, and



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the Actin stress fibers were clearly observed in the cells, similar to those on the FN-coated glass (Figure 8B). To quantify the projected cell spreading area, images were analyzed (magnification 200×) from six different fields of each sample. The projected surface area of adherent cells was calculated as the total cell cover area was normalized by the number of total cells in the field (the total area of each field was 300000 pixels). The cell-covered areas on each film were 20050.0 ( 1733.8 (FN-glass), 8260.2 ( 1132.2 (PLCL), 15285 ( 2041.1 (AAcPLCL), and 45783.7 ( 11838.4 pixels (gelatin-AAc-PLCL; Figure 9A). The number of cells on the corresponding films were 11.0 ( 3.1 (FN-glass), 11.5 ( 3.4 (PLCL), 11.7 ( 3.0 (AAc-PLCL), and 17.3 ( 5.5 (gelatin-AAc-PLCL; data not shown). Based on these parameters, the relative projected area of adherent single cells on the gelatin-AAc-PLCL film was 116.80 ( 12.78%, which was significantly greater than that on the other types of PLCL substrates. The values on the AAcPLCL and PLCL film were 58.45 ( 16.66 and 30.11 ( 5.07%, respectively. The relative values on each film were calculated to a positive control group (FN-glass), which occupied 100.00 ( 27.58% of the total area (Figure 9B). 3.6. Osteogenic Differentiation of hMSCs on the Gelatin-AAc-PLCL Films. After 7 days of culture of hMSCs on various substrates, the ALP activity was measured to evaluate the effect of gelatin immobilized PLCL films to osteogenic differentiation of hMSCs and normalized with total DNA contents and reaction time (Figure 10). The ALP activity from hMSCs on the gelatin-AAc-PLCL and AAc-PLCL films was 73.83 ( 26.1 and 34.64 ( 12.9 nmol/DNA/30 min, respectively, which was significantly greater than that from PLCL film (10.02 ( 2.31 nmol/DNA/30 min) (/p < 0.05). The order of ALP activity of each sample was gelatin-AAc-PLCL ) TCPS > AAc-PLCL > PLCL.

Figure 7. Proliferation of the hMSCs on the films (A) at 8, 24, and 72 h after seeding, and (B) relative cell adhesion at 24 h on the gelatin-AAc-PLCL films prepared with different gelatin concentrations. The symbol “/” indicates the significant difference (p < 0.05).

gelatin-AAc-PLCL, respectively). However, the cells on each film showed different proliferation rates after 24 h. For example, after 72 h of culture, the DNA concentration of the cells on the gelatin-AAc-PLCL film was 1390.7 ( 159.4 ng, which was significantly greater than that from both PLCL film (478.1 ( 95.7 ng) and AAc-PLCL film (948.91 ( 141.07 ng). To examine the effect of the gelatin concentration on the proliferation of hMSCs, various concentrations of gelatin solution (0.2, 2, 4 mg/ mL) were reacted with the AAc-PLCL films (Figure 7B). After 72 h of culture, the DNA contents on each film were normalized by those on the gelatin-AAc-PLCL film prepared by the reaction with 2 mg/mL gelatin solution. The relative DNA contents from the films reacted with 0.2 and 4 mg/mL gelatin solution were 75.40 ( 9.21 and 86.60 ( 10.81%, respectively. 3.5. Cell Spreading on the Gelatin-AAc-PLCL Films. The morphology of adherent cells on the films after 24 h is shown in Figure 8. The cells on PLCL film maintained round shapes and Actin stress fibers were not observed within the cytoplasm but were only localized in the periphery of the cell membrane (Figure 8C). However, the cells on the AAc-PLCL film were more spread than those on the PLCL film, and the number of adherent cells also increased (Figure 8D). The morphology of the cells on the gelatin-AAc-PLCL film showed a more polygonally elongated shape (Figure 8E) than other groups, and

4. Discussion The delivery of stem cells into tissue defects has attracted many scientists and medical doctors in the field of tissue engineering because of their potential to aid repair or regeneration of damaged tissue.3 Stem cells are capable of differentiating into multiple cell types such as osteoblasts, chondrocytes, and adipocytes when induced under appropriate microenvironments.29 A previous report has demonstrated that stem cells delivered within three-dimensional scaffolds exhibited differentiated properties of several types of target cells.30 However, currently available stem cell delivery systems are not ideal and often lead to critical problems, including the differentiation of the stem cells into undesirable phenotypes, limited integration of engineered tissue, and abnormal recovery of biological function.31 Therefore, it is important to develop appropriate stem cell delivery carriers that can control the adhesion of stem cells, maintain their proliferation and stemness, and subsequently direct desirable cell differentiation for the regeneration of particular types of tissue.1 Herein, we developed biofunctionalized substrates from a biodegradable polymer, PLCL, using the immobilization of gelatin, and demonstrated their potential to control adhesion, spreading, and proliferation of MSCs. To develop biofunctionalized substrates, we employed a γ-ray irradiation method to modify the surface of the biodegradable synthetic polymer, PLCL, which has no functional groups. Many surface modification techniques, including ultraviolet (UV) irradiation, ion-beam irradiation, plasma treatment, and chemical synthesis, involve two-step processes to introduce bioactive functional groups on the surface of polymeric substrates; linkers

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Figure 8. Morphologies of hMSCs cultured on various substrates: (A) phase contrast light microscope image of the adherent hMSCs on the tissue cell culture plate (scale bar ) 300 µm) and immunofluorescent staining of the adherent cells on the films, (B) fibronectin (FN)-coated glass, (C) PLCL, (D) AAc-PLCL, and (E) gelatin-AAc-PLCL. Larger images indicate the population of the cells on each film, obtained using a fluorescent microscope at 100× magnification (scale bar ) 200 µm). Inserted square images show detailed morphology of the adherent cells captured using laser scanning confocal microscope at 1000× magnification (scale bar ) 50 µm).

are immobilized directly to the synthetic polymer substrates and bioactive molecules are then conjugated to the linkers.17,25,31–33 The γ-ray irradiation was successful to immobilize carboxylic acid linkers to the surface of the PLCL film. As compared with other modification methods, γ-ray irradiation has been widely utilized, particularly in sterilization and surface modification of biomaterials due to its versatility in reaction conditions and relative easiness.24 In general, excessive exposure of polymer to γ-ray is known to induce degradation of polymer chains, which can reduce molecular weights and mechanical property of the polymer. As shown in Table 1, we observed the irradiation-induced decrease in the molecular weight of PLCL substrates. This could be due to the cleavage of long polymer chains by γ-ray-irradiation. However, the tensile strength and elongation properties of the substrates seem to be minimally affected by the irradiation (Table 2). PLCL is a hydrophobic polymer, and therefore, it is possible that the PLCL films can be formed by strong hydrophobic interactions of entangled chains. These strong interactions may attribute to the maintenance of mechanical properties of the substrates although the molecular weight of the polymers was slightly reduced.34 However, further detailed experiments would be necessary to

specifically understand the mechanism of irradiation-induced degradation of the polymer chains. Within the present study, we selected the 5 wt % AAc concentration and 10 kGy irradiation dose for the final optimized condition for the preparation of gelatin-AAc-PLCL film, whose mechanical properties may not be greatly altered in comparison to the PLCL film without irradiation. The graft yields of AAc by γ-ray were increased according to the irradiation doses and the AAc concentration, and swelling behavior showed similar patterns to the corresponding graft yields. In addition, we observed the broaden carbonyl peak of PLCL ranging from 1740 to 1700 cm-1, when the amount of grafted AAc increased, and the carbonyl peak shape changed to a more broaden shape. Those observations seem to be related to overlay of carbonyl stretching of PLCL (usually 1750-1730 cm-1) and the carboxylic acid stretching of AAc (usually 1725-1700 cm-1). Moreover, the quantitative analysis of AAc-PLCL films using toluidine blue O confirmed that the AAc-PLCL films showed greater intensity following reaction with toluidine blue O as the stock concentration of AAc increased (no positive staining for PLCL film). Previously, Grondahl and colleagues employed same γ-ray sources and reported the graft of AAc to poly(3-hydroxybu-


Figure 9. Relative area of adherent cells on the films. (A) The total cell-covered area and (B) the relative projected area of adherent single cells. The “/” indicates the significant difference (p < 0.05).

Figure 10. ALP assay after 7 days of culture in osteogenic differentiation media (/p < 0.05).

tyrate-co-3-hydroxyvalerate) (PHBV), demonstrating that the intensities of toluidine blue O-stained PHBV films increased as the graft yields increased.25 Additionally, Ying et al. modified


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poly(ethylene terephtalate) (PET) film with AAc under UV irradiation, and the presence of the surface carboxylic acid groups was indirectly analyzed by employing the same staining method.35 The AAc concentration range (from 0.2 to 10 wt %) for the reactions from previous studies was similar with that from our experimental design (from 3 to 10 wt %). The surface concentration of the carboxylic acid groups on the PET film reacted with 10wt % AAc in the previous report was 0.56 µM/ cm2, which was one-tenth comparing to our result (58.16 µM/ cm2). Although direct comparison may not be appropriate, our results suggest that γ-ray irradiation can be as effective (or even better) as other conventional surface modification strategies in the same experimental environment. To prepare biofunctionalized substrates that can favorably interact with cells, gelatin was immobilized to the AAc-PLCL. Natural polymers such as collagen, fibronectin, gelatin, and their cell interactive peptide sequence such as RGD and YIGSR have been immobilized on the scaffold and showed enhancement in cellular responses with respect to adhesion, spreading, proliferation, and differentiation.19,31,33,36,37 For example, the RGD immobilized-PCL film enhances the expression of adhesion receptors in human osteoblasts, and provides an interactive substrate with cells to generate more focal adhesions.32 Increased chondrocyte adhesion on gelatin immobilized-PLLA film was also reported and gelatin immobilized-PLLA film shows superior cell viability over that of normal PLLA film.19 Additionally, nanostructured collagen substrates demonstrate the ability to control the growth and the differentiation of mesenchymal stem cells.38 The amount of immobilized gelatin seemed to be affected by concentration of gelatin solution and grafted AAc. The quantification of amine groups after conjugation of gelatin confirmed that the immobilized amount of gelatin was increased as AAc-PLCL film reacted with higher concentration of gelatin solution. The passive absorption of gelatin was not effective for the immobilization, as shown in Figure 6B. As compared to the concentration of carboxylic acid groups, the presence of relatively lower amount of immobilized gelatin on the surface of AAc-PLCL film may be due to the size of gelatin molecule. Although saturated amounts of activated functional groups are available for allowing gelatin conjugation, the molecule conjugated to one functional group may mask other conjugation sites. Lee and colleagues reported that AAc-grafted silicone rubber substrate also showed limited protein immobilization, which is consistent with our results.39 In this particular study, it is hard to know the exact concentration of gelatin that is able to fully cover the surface of the substrate. Our experimental results demonstrated that cellular adhesion was affected in a concentration dependent manner up to 2 mg/mL of gelatin concentration and showed similar level of total DNA contents in 2 mg/mL and 4 mg/mL of gelatin solution, indicating that 2 mg/mL working solution may be the threshold level for our reaction condition. Therefore, immobilized gelatin was maximized as reacted with 2 mg/mL gelatin solution. In this study, cell responses, including adhesion, spreading, proliferation, and differentiation, were evaluated on gelatin-AAcPLCL film using human mesenchymal stem cells (hMSCs) isolated from the human mandible. hMSCs are known to have an ability to differentiate into multilineage cell types, including bone, cartilage, muscle, and tendon, when residing in an appropriate environment.40 Therefore, many researchers have focused on the development of novel scaffolds, which provide a defined artificial environment to control stem cell function in regenerative medicine.41–45 In particular, the initial adhesion and subsequent proliferation of hMSCs on synthetic substrates have

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been areas of active research because hMSCs are anchoragedependent cells, and the progression of adhesion-related events is highly implicated in differentiation and survival of hMSCs.46 The proliferation of hMSCs cultured on polymer surfaces was affected by the surface chemical properties; the initial adhesion was similar, however, the proliferation rate of hMSCs at 72 h was enhanced only in the gelatin-AAc-PLCL. Because PLCL film is highly hydrophobic, the cells may have been loosely attached, which may have attributed to relatively low cell proliferation. The graft of AAc rendered the PLCL film more hydrophilic, which may have contributed to enhance hMSCs proliferation. Previous studies have demonstrated that the hydrophilicity of the material surface plays a critical role in cell adhesion and proliferation.31 In addition, electrostatic interactions between charges in cell membranes and positively/ negatively charged surface functional groups are important parameters in determining cell adhesion and proliferation.47 Our results are in good agreement with other findings and, furthermore, indicate that the signals for the proliferation of hMSCs cultured to AAc-PLCL film can be enhanced by the addition of gelatin. Cui et al. reported that gelatin-immobilized PLLA films can enhance the proliferation of chondrocytes.19 The interaction of gelatin and hMSCs may have synergistically influenced the proliferation of the cell on the hydrophilic or electrostatic control by the substrate. The effect of gelatin on the proliferation of hMSCs depended on the reacting gelatin concentration of up to 2 mg/mL, which seemed to be the saturated concentration for the immobilization of gelatin to the AAc-PLCL film. Anchorage-dependent cells need to strongly adhere to substrates and, thereby, can survive, differentiate, and transfer signals into the cytosol to maintain cell homeostasis. Therefore, many reports highlighted the importance of cell spreading on engineered substrates. For example, Cheng et al. reported that when human myoblasts are cultured on PCL film, adherent cells cannot spread and most of the cells are washed out from the normal PCL film. On the contrary, the cells cultured on the PCL-collagen film spread well in all directions and their proliferation rate also increases, suggesting an important relationship between proliferation and adhesion of the cells on the collagen-modified PCL substrate.31 In our results, the adherent cells on the PLCL film were poorly spread, and they maintained a round shape without clear observation of Actin stress fibers. Although the cells on the AAc-PLCL film were spread more widely than cells on the PLCL film, the formation of the Actin stress fibers was immature. However, the formation of the Actin stress fibers in the cells cultured on gelatin-AAcPLCL was clearly observed and the morphology and relative spreading area of a single adherent cell were similar to the cells on the FN-glass. These results suggest that the initial binding and formation of the highly spread morphology of hMSCs can be an important consideration when evaluating cell-substrate interaction. To examine the effect of gelatin-AAc-PLCL on the differentiation of hMSCs, we investigated the osteogenic differentiation of hMSCs cultured on the functional substrates. Our results showed that the alkaline phosphatase (a common marker of osteogenic differentiation of hMSCs) activity of hMSCs on the gelatin-AAc-PLCL films was increased 7-fold greater than that on the pristine PLCL films, which was in good agreement of our cell adhesion and proliferation results. The study about the control of hMSC differentiation into multiple lineages using the gelatin-AAc-PLCL film should be the target of our future investigation. In comparison with tissue culture plate, gelatin-AAc-PLCL showed lower cell adhesion and

Shin et al.

proliferation, however, the differentiating hMSCs produced similar levels of ALP, indicating that different signaling may be implicated in each cell behavior. Therefore, it is important to consider various characteristics of the substrates to fully modulate desirable cell function. Collectively, the biofunctionalized substrates immobilized with gelatin can offer a favorable environment for hMSC survival, and gelatin plays a critical role as an active mediator to connect the substrate and cells.

5. Conclusions To prepare biofunctionalized substrates, AAc was grafted to the PLCL film by γ-ray irradiation, and gelatin was immobilized to the activated AAc. The graft yield was dependent on the irradiation dose and AAc concentrations; the grafted amount of AAc increased with higher irradiation dose and AAc concentration. The evaluation of the biological performance of the gelatin-AAc-PLCL films was conducted using hMSCs as a model cell type. The proliferation of hMSCs at 72 h was dependent on the surface properties, and the DNA concentrations extracted from the cells cultured on the gelatin-AAc-PLCL film were 2.9- and 1.5-fold greater than those on the PLCL and the AAc-PLCL films, respectively. Fluorescent microscopic analysis revealed that cells on gelatin-AAc-PLCL film showed a mature form of Actin stress fibers and a polygonally elongated structure in all directions, while limited spreading of hMSCs was observed on other substrates. In addition, the spreading areas of adherent cells on the gelatin-AAc-PLCL film were 3.8- and 2.0-fold greater than those on the PLCL and the AAc-PLCL films, respectively. In osteogenic differentiation, gelatin-AAc-PLCL film facilitated differentiation of hMSCs. Collectively, our results suggest that the gelatin-modified synthetic substrates may hold great promise as substrates for the delivery of hMSCs in many tissue engineering applications. Acknowledgment. This work was supported by the Seoul R&BD program (CR070027) funded by Seoul Development institute and the Korea Science and Engineering Foundation (KOSEF) grant funded by the Korea government (MOST: No. 2008-01224; to H.S.).

References and Notes (1) Shin, H.; Jo, S.; Mikos, A. G. Biomimetic materials for tissue engineering. Biomaterials 2003, 24 (24), 4353–64. (2) Hench, L. L.; Polak, J. M. Third-generation biomedical materials. Science 2002, 295 (5557), 1014–7. (3) Langer, R.; Tirrell, D. A. Designing materials for biology and medicine. Nature 2004, 428 (6982), 487–92. (4) Lee, S. H.; Shin, H. Matrices and scaffolds for delivery of bioactive molecules in bone and cartilage tissue engineering. AdV. Drug DeliVery ReV. 2007, 59 (4-5), 339–59. (5) Weigel, T.; Schinkel, G.; Lendlein, A. Design and preparation of polymeric scaffolds for tissue engineering. Expert ReV. Med. DeVices 2006, 3 (6), 835–51. (6) Discher, D. E.; Janmey, P.; Wang, Y. L. Tissue cells feel and respond to the stiffness of their substrate. Science 2005, 310 (5751), 1139–43. (7) Engler, A. J.; Sen, S.; Sweeney, H. L.; Discher, D. E. Matrix elasticity directs stem cell lineage specification. Cell 2006, 126 (4), 677–89. (8) Jeong, S. I.; Kim, S. H.; Kim, Y. H.; Jung, Y.; Kwon, J. H.; Kim, B. S.; Lee, Y. M. Manufacture of elastic biodegradable PLCL scaffolds for mechano-active vascular tissue engineering. J. Biomater. Sci., Polym. Ed. 2004, 15 (5), 645–60. (9) Jeong, S. I.; Kwon, J. H.; Lim, J. I.; Cho, S. W.; Jung, Y.; Sung, W. J.; Kim, S. H.; Kim, Y. H.; Lee, Y. M.; Kim, B. S.; Choi, C. Y.; Kim, S. J. Mechano-active tissue engineering of vascular smooth muscle using pulsatile perfusion bioreactors and elastic PLCL scaffolds. Biomaterials 2005, 26 (12), 1405–11. (10) Garkhal, K.; Verma, S.; Tikoo, K.; Kumar, N. Surface modified poly(Llactide-co-epsilon-caprolactone) microspheres as scaffold for tissue engineering. J. Biomed. Mater. Res. A 2007, 82 (3), 747–56.

Human Mesenchymal Stem Cells (11) Shin, H. Fabrication methods of an engineered microenvironment for analysis of cell-biomaterial interactions. Biomaterials 2007, 28 (2), 126–33. (12) Pompe, T.; Keller, K.; Mothes, G.; Nitschke, M.; Teese, M.; Zimmermann, R.; Werner, C. Surface modification of poly(hydroxybutyrate) films to control cell-matrix adhesion. Biomaterials 2007, 28 (1), 28–37. (13) Aframian, D. J.; Cukierman, E.; Nikolovski, J.; Mooney, D. J.; Yamada, K. M.; Baum, B. J. The growth and morphological behavior of salivary epithelial cells on matrix protein-coated biodegradable substrata. Tissue Eng. 2000, 6 (3), 209–16. (14) Ho, M. H.; Lee, J. J.; Fan, S. C.; Wang, D. M.; Hou, L. T.; Hsieh, H. J.; Lai, J. Y. Efficient modification on PLLA by ozone treatment for biomedical applications. Macromol. Biosci. 2007, 7 (4), 467–74. (15) Carlisle, E. S.; Mariappan, M. R.; Nelson, K. D.; Thomes, B. E.; Timmons, R. B.; Constantinescu, A.; Eberhart, R. C.; Bankey, P. E. Enhancing hepatocyte adhesion by pulsed plasma deposition and polyethylene glycol coupling. Tissue Eng. 2000, 6 (1), 45–52. (16) Suh, H.; Hwang, Y. S.; Lee, J. E.; Han, C. D.; Park, J. C. Behavior of osteoblasts on a type I atelocollagen grafted ozone oxidized poly L-lactic acid membrane. Biomaterials 2001, 22 (3), 219–30. (17) Ma, Z.; Gao, C.; Gong, Y.; Shen, J. Cartilage tissue engineering PLLA scaffold with surface immobilized collagen and basic fibroblast growth factor. Biomaterials 2005, 26 (11), 1253–9. (18) Zhu, Y.; Gao, C.; Liu, X.; He, T.; Shen, J. Immobilization of biomacromolecules onto aminolyzed poly(L-lactic acid) toward acceleration of endothelium regeneration. Tissue Eng. 2004, 10 (1-2), 53–61. (19) Cui, Y. L.; Hou, X.; Qi, A. D.; Wang, X. H.; Wang, H.; Cai, K. Y.; Ji Yin, Y.; De Yao, K. Biomimetic surface modification of poly(Llactic acid) with gelatin and its effects on articular chondrocytes in vitro. J. Biomed. Mater. Res. A 2003, 66 (4), 770–8. (20) Park, K.; Ju, Y. M.; Son, J. S.; Ahn, K. D.; Han, D. K. Surface modification of biodegradable electrospun nanofiber scaffolds and their interaction with fibroblasts. J. Biomater. Sci., Polym. Ed. 2007, 18 (4), 369–82. (21) Teraoka, F.; Nakagawa, M.; Hara, M. Surface modification of poly(Llactide) by atmospheric pressure plasma treatment and cell response. Dent. Mater. J. 2006, 25 (3), 560–5. (22) Yamaguchi, M.; Shinbo, T.; Kanamori, T.; Wang, P. C.; Niwa, M.; Kawakami, H.; Nagaoka, S.; Hirakawa, K.; Kamiya, M. Surface modification of poly(L-lactic acid) affects initial cell attachment, cell morphology, and cell growth. J. Artif. Organs. 2004, 7 (4), 187–93. (23) Yang, J.; Wan, Y.; Yang, J.; Bei, J.; Wang, S. Plasma-treated, collagenanchored polylactone: Its cell affinity evaluation under shear or shearfree conditions. J. Biomed. Mater. Res. A 2003, 67 (4), 1139–47. (24) Nho, Y.; Lim, Y.; Lee, Y. Preparation, properties and biological application of pH-sensitivie poly(ethylene oxide) (PEO) hydrogels grafted with acrylic (AAc) using gamma-ray irradiation. Radiat. Phys. Chem. 2004, (71), 237–240. (25) Grondahl, L.; Chandler-Temple, A.; Trau, M. Polymeric grafting of acrylic acid onto poly(3-hydroxybutyrate-co-3-hydroxyvalerate): Surface functionalization for tissue engineering applications. Biomacromolecules 2005, 6 (4), 2197–203. (26) Hinrichs, W. L. J.; Hoopen, H. W. M. T.; Wissink, M. J. B.; Engbers, G. H. M.; Feijen, J. Design of a new type of coating for the controlled release of heparin. J. Controlled Release 1997, 45 (2), 163–176. (27) Choi, J. S.; Yoo, H. S. Electrospun nanofibers surface-modified with fluorescent proteins. J. Bioact. Compat. Polym. 2007, 22 (5), 508– 524. (28) Ng, K. W.; Leong, D. T.; Hutmacher, D. W. The challenge to measure cell proliferation in two and three dimensions. Tissue Eng. 2005, 11 (1-2), 182–91.


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(29) Guillot, P. V.; Cui, W.; Fisk, N. M.; Polak, D. J. Stem cell differentiation and expansion for clinical applications of tissue engineering. J. Cell. Mol. Med. 2007, 11 (5), 935–44. (30) Hwang, N. S.; Varghese, S.; Elisseeff, J. Controlled differentiation of stem cells. AdV. Drug DeliVery ReV. 2007, 60 (2), 199–214. (31) Cheng, Z.; Teoh, S. H. Surface modification of ultra thin poly(caprolactone) films using acrylic acid and collagen. Biomaterials 2004, 25 (11), 1991–2001. (32) Amato, I.; Ciapetti, G.; Pagani, S.; Marletta, G.; Satriano, C.; Baldini, N.; Granchi, D. Expression of cell adhesion receptors in human osteoblasts cultured on biofunctionalized poly-(-caprolactone) surfaces. Biomaterials 2007, 28 (25), 3668–78. (33) Kim, T. G.; Park, T. G. Biomimicking extracellular matrix: cell adhesive RGD peptide modified electrospun poly(D,L-lactic-co-glycolic acid) nanofiber mesh. Tissue Eng. 2006, 12 (2), 221–33. (34) Rutnakornpituk, M.; Ngamdee, P. Surface and mechanical properties of microporous membranes of poly(ethylene glycol)-polydimethylsiloxane copolymer/chitosan. Polymer 2006, 47, 7909–7917. (35) Ying, L.; Yin, C.; Zhuo, R. X.; Leong, K. W.; Mao, H. Q.; Kang, E. T.; Neoh, K. G. Immobilization of galactose ligands on acrylic acid graft-copolymerized poly(ethylene terephthalate) film and its application to hepatocyte culture. Biomacromolecules 2003, 4 (1), 157– 65. (36) Klee, D.; Ademovic, Z.; Bosserhoff, A.; Hoecker, H.; Maziolis, G.; Erli, H. J. Surface modification of poly(vinylidenefluoride) to improve the osteoblast adhesion. Biomaterials 2003, 24 (21), 3663–70. (37) Santiago, L. Y.; Nowak, R. W.; Peter Rubin, J.; Marra, K. G. Peptidesurface modification of poly(caprolactone) with laminin-derived sequences for adipose-derived stem cell applications. Biomaterials 2006, 27 (15), 2962–9. (38) Shih, Y. R.; Chen, C. N.; Tsai, S. W.; Wang, Y. J.; Lee, O. K. Growth of mesenchymal stem cells on electrospun type I collagen nanofibers. Stem Cells 2006, 24 (11), 2391–7. (39) Lee, S. D.; Hsiue, G. H.; Chang, P. C.; Kao, C. Y. Plasma-induced grafted polymerization of acrylic acid and subsequent grafting of collagen onto polymer film as biomaterials. Biomaterials 1996, 17 (16), 1599–608. (40) Li, W. J.; Tuli, R.; Huang, X.; Laquerriere, P.; Tuan, R. S. Multilineage differentiation of human mesenchymal stem cells in a threedimensional nanofibrous scaffold. Biomaterials 2005, 26 (25), 5158– 66. (41) Betz, V. M.; Betz, O. B.; Harris, M. B.; Vrahas, M. S.; Evans, C. H. Bone tissue engineering and repair by gene therapy. Front. Biosci. 2008, 13, 833–41. (42) Fan, H.; Liu, H.; Toh, S. L.; Goh, J. C. Enhanced differentiation of mesenchymal stem cells co-cultured with ligament fibroblasts on gelatin/silk fibroin hybrid scaffold. Biomaterials 2007, 29 (8), 1017– 27. (43) Jones, E.; McGonagle, D. Human bone marrow mesenchymal stem cells in vivo. Rheumatology (Oxford) 2008, 47 (2), 126–31. (44) Vashi, A. V.; Keramidaris, E.; Abberton, K. M.; Morrison, W. A.; Wilson, J. L.; O’Connor, A. J.; Cooper-White, J. J.; Thompson, E. W. Adipose differentiation of bone marrow-derived mesenchymal stem cells using Pluronic F-127 hydrogel in vitro. Biomaterials 2008, 29 (5), 573–79. (45) Yen, A. H.; Sharpe, P. T. Stem cells and tooth tissue engineering. Cell. Tissue Res. 2007, 331 (1), 359–72. (46) Kenar, H.; Kocabas, A.; Aydinli, A.; Hasirci, V. Chemical and topographical modification of PHBV surface to promote osteoblast alignment and confinement. J. Biomed. Mater. Res. A 2007, 85 (4), 1001–10. (47) Yang, J. M.; Huang, M. J.; Yeh, T. S. Preparation of poly(acrylic acid) modified polyurethane membrane for biomaterial by UV radiation without degassing. J. Biomed. Mater. Res. 1999, 45 (2), 133–9.

BM701410G

Modulation of Spreading, Proliferation, and Differentiation of Human

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