Umbilical-cord-blood-derived mesenchymal stem cells seeded onto fibronectin-immobilized polycaprolactone nanofiber improve cardiac function

Acta Biomaterialia 10 (2014) 3007–3017

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Umbilical-cord-blood-derived mesenchymal stem cells seeded onto fibronectin-immobilized polycaprolactone nanofiber improve cardiac function Byung-Jae Kang a,1, Hwan Kim b,1, Seul Ki Lee c, Joohyun Kim a, Yiming Shen c, Sunyoung Jung d, Kyung-Sun Kang e, Sung Gap Im f, So Yeong Lee c, Mincheol Choi d, Nathaniel S. Hwang b,⇑, Je-Yoel Cho a,⇑ a

Department of Veterinary Biochemistry, BK21 Plus and Research Institute of Veterinary Science, College of Veterinary Medicine, Seoul National University, Seoul, Republic of Korea School of Chemical and Biological Engineering, BioMAX Institute, Seoul National University, Seoul, Republic of Korea c Laboratory of Veterinary Pharmacology, College of Veterinary Medicine, Seoul National University, Seoul, Republic of Korea d Department of Veterinary Medical Imaging, College of Veterinary Medicine, Seoul National University, Seoul, Republic of Korea e Adult Stem Cell Research Center, College of Veterinary Medicine, Seoul National University, Seoul, Republic of Korea f Department of Chemical and Biomolecular Engineering, Korea Advanced Institute of Technology, Daejeon, Republic of Korea b

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Article history: Received 15 October 2013 Received in revised form 8 March 2014 Accepted 10 March 2014 Available online 19 March 2014 Keywords: Heart Umbilical-cord-blood-derived mesenchymal stem cell Polycaprolactone nanofiber Fibronectin Transplantation

a b s t r a c t Stem cells seeded onto biofunctional materials have greater potency for therapeutic applications. We investigated whether umbilical-cord-blood-derived mesenchymal stem cell (UCB-MSC)-seeded fibronectin (FN)-immobilized polycaprolactone (PCL) nanofibers could improve cardiac function and inhibit left ventricle (LV) remodeling in a rat model of myocardial infarction (MI). Aligned nanofibers were uniformly coated with poly(glycidyl methacrylate) by initiated chemical vapor deposition followed by covalent immobilization of FN proteins. The degree of cell elongation and adhesion efficacy were improved by FN immobilization. Furthermore, genes related to angiogenesis and mesenchymal differentiations were up-regulated in the FN-immobilized PCL nanofibers in comparison to control PCL nanofibers in vitro. 4 weeks after the transplantation in the rat MI model, the echocardiogram showed that the UCB-MSCseeded FN-immobilized PCL nanofiber group increased LV ejection fraction and fraction shortening as compared to the non-treated control and acellular FN-immobilized PCL nanofiber groups. Histological analysis indicated that the implantation of UCB-MSCs with FN-immobilized PCL nanofibers induced a decrease in MI size and fibrosis, and an increase in scar thickness. This study indicates that FN-immobilized biofunctional PCL nanofibers could be an effective carrier for UCB-MSC transplantation for the treatment of MI. Ó 2014 Published by Elsevier Ltd. on behalf of Acta Materialia Inc.

1. Introduction Heart failure (HF) is the leading cause of death all over the world [1]. An important cause of HF is myocardial infarction (MI), which is often caused by coronary artery obstruction leading to cardiomyocyte loss due to a lack of oxygen supply. The cardiomyocytes in adult mammalian hearts have limited self-regeneration capacities, and they intrinsically do not repopulate themselves when subjected to MI [2]. This overburdens the surviving myocardium and ultimately leads to impairment of left ventricular (LV) function. Therefore, an alternative therapeutic intervention is required to compensate for the inadequate intrinsic repair mechanism. ⇑ Corresponding authors. Tel.: +82 2 880 1268; fax: +82 2 886 1268 (J.-Y. Cho). Tel.: +82 2 880 1635; fax: +82 2 888 7295 (N.S. Hwang). E-mail addresses: [email protected] (N.S. Hwang), [email protected] (J.-Y. Cho). 1 These authors contributed equally to this work. http://dx.doi.org/10.1016/j.actbio.2014.03.013 1742-7061/Ó 2014 Published by Elsevier Ltd. on behalf of Acta Materialia Inc.

Despite the improvements in medical and surgical treatments for MIs, therapeutic options are limited as many approaches, including reactivation of cardiomyocyte cell cycle activity and reduction of myocardial cell death, showed unsuccessful inhibition of ventricular enlargement [3,4]. In addition, the morbidity/mortality rate remains high, even after these interventions [2,5,6]. Recently, stem-cell-based approaches have received considerable attention for their potential to regenerate the infarcted heart and restore cardiac function after an MI [2,6–9]. In particular, umbilical-cord-blood (UCB)-derived mesenchymal stem cells (MSCs), which have the capacity to differentiate into osteoblasts, chondrocytes and adipocytes, were shown to differentiate into cardiomyocytes, both in vitro and in vivo [2,10–12]. UCB-MSCs are a promising candidate for the treatment of MI due to their lower immunogenic response compared to other stem cells, and this may potentially allow UCB-MSCs for off-the-shelf cell source for allogeneic therapies [13].

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The functional beneficial effects of implanted stem cells are most likely due to the paracrine factors secreted by the cells [2,6,14]. MSCs have been shown to produce stem cell homing, anti-apoptosis, anti-inflammatory, anti-scarring and angiogenic cytokines that support other cells present in the injured myocardium, and this may be the primary mechanism for the recovery of cardiac function and the reduction of myocardial fibrosis [2,6,10,14]. Stem cells can be delivered into the heart by an intravenous injection into coronary artery or by a direct injection into the myocardium. However, stem cell delivery with these techniques is limited due to low retention and survival [2,14–17]. Typically, more than 80–90% of grafted cells die within 72 h after injection [18,19]. In order to improve survival and engraftments of the implanted cells, combining cells with various nanofiber-based biomaterials can be considered as a promising approach in cardiac tissue repair after MI [3,17,20]. Previous studies showed that these new strategies could improve grafted cell retention and survival [3,17,21,22]. The ideal scaffold for cardiac tissue engineering should display both mechanical and biological properties of native heart tissue, which can improve cell attachment and proliferation, promote cellular morphology and provide the flexibility [23]. Furthermore, scaffolds for cardiac tissue engineering should be vascularized (or at least vascularized soon after implantation) and non-immunogenic, and provide a suitable cellular interaction [17,20]. These properties can alter the expansion and contraction of the heart. Among the various methods utilized for cardiac tissue engineering, electrospinning is of immense interest since it is possible to make nonwoven meshes in nanometer scaffolds that are architecturally similar to the native extracellular matrix (ECM) of the heart [24,25]. Furthermore, electrospun nanofibers provide a high surface-to-volume ratio and versatility to surface engineer for selective biological activity. Electrospun nanofibers based on polycaprolactone (PCL) have been extensively utilized for myocardial regeneration as three-dimensional cardiac tissue regeneration [26,27]. However, these biocompatible materials do not provide appropriate ECM-mediated biological signals for enhanced cellular attachment and migration. Thus, the presentation of appropriate chemical and physical cues onto nanofibers are required for enhanced cellular attachment and bioactivity [28]. Fibronectin (FN) is known as an important ECM molecule for stem cell adhesion, survival and differentiation [29–31]. Furthermore, this protein is expressed in the normal heart. Therefore, an effect of this ECM molecule on stem cell attachment and function can be expected if stem cells are applied once FN has accumulated onto nanofibers. In order to strengthen the cell–ECM interactions, we immobilized FN onto aligned PCL nanofibers functionalized by initiated chemical vapor deposition (iCVD) polymer films in order to modulate UCB-MSC behavior. Specifically, we deposited poly(glycidyl methacrylate) (pGMA), which contains reactive epoxy functional group that readily undergo a ring-opening reaction with peptide amine groups, onto the PCL nanofibers. The objective of the current study was to investigate whether UCB-MSCs could improve cardiac function and inhibit left LV remodeling in a rat model of MI. In this study, we demonstrate that FN immobilization on PCL nanofibers increased the UCB-MSC adhesion. Furthermore, when FN-immobilized PCL nanofibers seeded with UCB-MSCs was implanted on the epicardial surface over infarcted areas, cardiac function was improved and LV remodeling was prevented.

2. Materials and methods 2.1. Isolation and culture of human UCB-MSCs Human UCB-MSCs were isolated as previously described [32,33]. The human UCB-MSC isolation procedure was approved

by the Borame Hospital Institutional Review Board and Seoul National University (IRB No. 0603/001-002-07C1). Briefly, UCB samples from term and preterm deliveries were harvested at the time of birth with the mother’s informed consent (Seoul City Borame Hospital Cord Blood Bank). The mesenchymal stem cells were separated from the UCB using Ficoll-Paque TM PLUS (Amersham Bioscience, Uppsala, Sweden), and suspended in Dulbecco’s modified Eagle medium (DMEM; Gibco, Grand Island, NY, USA), containing 20% fetal bovine serum (FBS; Gibco), 100 IU ml1 penicillin, 100 mg ml1 streptomycin, 2 mM L-glutamine and 1 mM sodium pyruvate. After 24 h, unattached cells and residual nonadherent red blood cells were removed by washing. The medium was changed at 48 h intervals until the cells became confluent. After cells reached 90% confluence, they were subcultured. 2.2. Preparation of aligned-cardiac nanofiber patch PCL with a molecular weight of 43,000–50,000 g mol1 was purchased from Polysciences Inc. (Warington, PA, USA). PCL was dissolved in chloroform at a concentration of 25% (w/v). The electrospinning process was conducted as follows: the polymer solution was transferred into a 5 ml syringe (DAIHAN single-use syringes, with rubber gasket, Korea) with an orthogonally cutended needle (G-22Luer, Korea). A syringe pump (Kd Scientific, Massachusetts, USA) was used to control the solution flow rate at 0.5 ml h1. To produce aligned nanofibers, a voltage of 12 kV was applied between the syringe needle and the rotary drum collector (NanoNC, Korea) with a rotating speed of 1000 rpm. Collected fibers were cut into circular 8 mm diameter shapes and further processed for surface modification. 2.3. Fabrication of functionalized cardiac nanofiber patch The glycidyl methacrylate (GMA) monomer and the tert-butyl peroxide (TBPO) initiator were purchased from Sigma–Aldrich. The pGMA films were deposited onto the nanofibers in a custombuilt reactor (Daeki Hi-Tech Co. Ltd, Korea) via the iCVD process as previously described [34]. For grafting the pGMA films, the GMA monomer was vaporized in a jar, heated to 35 °C and fed into the reactor. An initiator (TBPO) was fed to the reactor through a different port and mixed with the monomer in the reactor. The flow rate of the GMA monomer was kept constant at 1.9 sccm (standard cubic centimeters per minute) while the flow rate of TBPO was maintained at 0.8 sccm. pGMA-coated nanofibers were washed with phosphate buffered saline (PBS) three times, and further immobilized with fluorescein-conjugated chitosan to confirm the reactivity of pGMA. pGMA-coated PCL nanofibers were then immobilized with FN (10 lg ml1) for enhanced cellular adhesions. For the nanofiber characterization, Fourier transform infrared (FTIR, Alpha FT-IR Spectrometer, Bruker, Billerica, MA, USA) spectra were obtained in normal absorbance mode with an average of 64 scans. All spectra were baseline corrected. FT-IR bands were recorded as spectra in the range of 400–4000 cm1. Water contact angles were determined on a contact angle analyzer (Kruss DSA 100, Germany) using the sessile drop method. Measurements were performed at room temperature, 30 s after deposition of the single drops of distilled water on the sample surface. The average of three measurements was used as a result. For in vitro degradation of nanofibers, the nanofiber scaffolds were immersed in 1 M sodium hydroxide (NaOH), and washed three times with PBS. Then all samples were dried and weighed for mass loss determination. The tensile strength of fabricated nanofibers was measured using and Instron 5583 (Instron Corporation, MA, USA) with the use of a 3.35 N load cell under a cross-head speed of 3 mm min1 at ambient conditions. All nanofiber samples were prepared in the form of

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rectangular shape with dimensions of 20  10 mm. The thickness of samples was 10 lm. 2.4. Cellular viability, adhesion and morphological analysis UCB-MSCs (2.5  106 cells) were uniformly seeded onto FNimmobilized PCL nanofibers (8 mm in diameter). Initial cellular adhesion rates were measured by PicoGreen AssayÒ. Cells were collected 48 h after seeding and DNA content was quantified using a Quant-iT™ PicoGreenÒdsDNA Assay Kit (Invitrogen™). For viability testing, a live/dead cell viability/cytotoxicity kit (Molecular Probes, L-3224) was used, following the manufacturer’s protocol. For cellular morphological analysis, cells were fixed at 4% paraformaldehyde and stained with 40 6-diamidino-2-phenylindole (DAPI; Sigma–Aldrich). 2.5. Scanning electron microscopy Cell-seeded nanofibers were fixed with a fixative composed of 4.0% formaldehyde, 1.5% glutaraldehyde, 2.5% sucrose, 5 mM calcium chloride and 5 mM magnesium chloride in 0.1 M sodium cacodylate for 20 min and washed with 2.5% sucrose in 0.1 M sodium cacodylate. Fixed samples were washed with cold distilled water and dehydrated with a graded series of cold ethanol for dehydration, and samples were then incubated with 50% ethanol and 50% hexamethyldisilazane (HMDS) for 5 min and exchanged with 100% HMDS for 10 min. Prior to scanning electron microscopy (SEM) imaging, nanofibers were sputtered with platinum for 100 s at 20 mA. Field emission SEM images were obtained with a JEOL 6700 instrument. 2.6. Reverse transcriptase–polymerase chain reaction array Total RNAs were extracted with Trizol, and reverse-transcribed into cDNA using the SuperScript Synthesis System (Invitrogen™). Room temperature reactions were first denatured for 2 min at 95 °C, followed by 35 cycles of 30 s denaturation at 95 °C, annealing and elongation at 72 °C. PCR array analyses were performed using human mesenchymal stem cell PCR arrays (SuperArray Bioscience) according to the manufacturer’s instructions using the Applied Biosystem StepOnePlus™ Real Time PCR System (Applied Biosystems). Internal control was normalized to housekeeping genes, which are B2M, HPRT1, RPL13A, GAPDH, ACTB and HGDC. Gene expression fold was compared with day 1 sample. 2.7. MI model and implantation UCB-MSC-seeded nanofibers were maintained in vitro in keratinocyte-serum free medium (Gibco) supplemented with 10% FBS and antibiotic–antimycotic (Gibco) at 37 °C in a 5% CO2 humidified atmosphere for 7 days prior to implantation. As a control, noncell-seeded FN-immobilized PCL nanofibers were prepared by the same method but without UCB-MSCs. Male Sprague–Dawley rats weighing 260–300 g (Orient Bio, Seongnam, Korea) were used in the MI model. All animal procedures were performed in accordance with the guidelines of the Institutional Animal Care and Use Committee of Seoul National University. MI was induced by ligation of the left coronary artery according to the methods previously reported [35]. 2 weeks after induction of MI, rats were randomized into three treatment groups, which received implantation of UCB-MSCsseeded FN-immobilized PCL scaffold (UCB-MSC/FN-PCL group, n = 10), implantation of non-cell-seeded FN-immobilized PCL scaffold (FN-PCL group, n = 5) or sham operation (control group, n = 4). Cell-seeded or non-seeded FN-immobilized PCL nanofibers were fixed with fibrin sealant (TISSEL, Baxter, Deerfield, IL, USA) over the infarcted region and adjacent normal heart. An antirejection

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strategy was used according to methods previously reported because a rat’s immune system could reject human cell xenograft after implantation of UCB-MSCs [36]. All the rats received cyclosporine (SandimmuneÒ, Novartis, Basel, Switzerland) 10 mg kg1 day1 subcutaneously for 2 weeks starting 2 days prior to grafting, followed by 100 lg ml1 in the drinking water for the next 2 weeks. 2.8. Functional assessment of infarcted myocardium Cardiac function was assessed by transthoracic echocardiography before MI (normal baseline), 1 week after MI (infarct baseline) and 4 weeks after scaffold implantation. Echocardiography was performed using a conventional echocardiographic machine (ProSound Alpha 7; ALOKA, Tokyo, Japan) with a 9 MHz linear transducer by a veterinary radiologist unaware of the treatment condition. For analysis of LV function, LV internal diameter at end-diastole (LVIDd) and LV internal diameter at end-systole (LVIDs) were measured at the anterior wall, from the short-axis view, just below the level of the papillary muscle. LV end-diastolic volume (LVEDV) and LV end-systolic volume (LVESV) were automatically calculated by the single-plane area and length method. The fractional shortening (FS) was calculated as ((LVIDd – LVIDs)/ LVIDd)  100 (%), and ejection fraction (EF) was calculated as ((LVEDV – LVESV)/LVESV)  100 (%). 2.9. Histological and immunohistochemical examinations Hearts were harvested and fixed in 4% paraformaldehyde for 24 h and then cut into four transverse slices through the infarcted area. The slices were then embedded in paraffin and 5 lm histologic sections were stained with Masson’s Trichrome. The infarct size, fibrosis and scar thickness were quantified using ImageJ software (NIH, Bethesda, MD, USA). The infarct size was defined as the percentage of the sum of the infarcted epicardial and endocardial circumferences divided by the sum of the LV epicardial and endocardial circumferences. The percentage area of fibrosis was quantified by the fibrotic area divided by the sum of the fibrotic area and non-fibrotic area. Scar thickness was measured in the center of the infarct, at the border of the infarct and in between, and then these five values were averaged. Immunohistochemical staining was performed to identify the survival of transplanted cells and vessel formation using the primary antibodies: rabbit monoclonal anti-lamin A+C antibody (1:200, abcamÒ, Cambridge, MA, USA) and rabbit polyclonal antivon Willebrand factor (vWF) antibody (1:100, abcamÒ). The sections were incubated with primary antibodies overnight and were subsequently exposed to biotinylated secondary antibody and streptavidin peroxidase complex using Histostain-Plus kit (Invitrogen™). Then, the sections were visualized by reaction with 3,30 diaminobenzidine tetrahydrochloride (liquid DAB substrate kit; Invitrogen™) and were counterstained with hematoxylin. The anti-lamin A+C antibody reacts with human but does not react with mouse and rat. Lamin A+C antibody was used to identify the survival of implanted human UCB-MSCs. The density of blood vessels was determined in the sections stained with anti-vWF antibody. Five high-power fields (HPFs; 200 magnification) within the infarcted region of each section were chosen randomly, and blood vessel density was expressed as the quantity of vessels per mm2. Averages based on five HPFs from each of the two samples per rat were calculated for comparison. 2.10. Statistical analysis Statistical analysis was performed using SPSS version 20.0 (SPSS, Chicago, IL, USA). A Kruskal–Wallis test was used to assess

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differences among the groups. A post hoc test was performed along with a Mann–Whitney U test. A P value less than 0.05 was considered to be statistically significant. 3. Results 3.1. Fabrication of aligned-cardiac nanofiber patch In order to immobilize FN on the PCL nanofiber, we introduced pGMA film, resulting in 50 nm thick pGMA. Fig. 1 demonstrates schematically the procedure for producing pGMA-coated and FN immobilized pGMA-PCL nanofibers. First, pGMA films were deposited on the aligned PCL nanofibers via iCVD. Surface coating of pGMA functionalized nanofibers with epoxy groups. Consequently, pGMA-coated PCL nanofibers were immobilized with FN (10 lg ml1) for enhanced cellular adhesions (Fig. 1A). Immobilization of FN formed new ether bonds after epoxy rings were broken. In order to assess the in vivo functionality of FN immobilized nanofiber, we seeded UCB-MSCs onto fabricated nanofibers, then transplanted as a patch in myocardial infarction model (Fig. 1B). 3.2. Characteristics of fabricated PCL nanofibers SEM micrograph of electrospun fibrous scaffolds revealed uniform aligned fibers. Furthermore, SEM images of nanofibers showed that post modification of pGMA coating by iCVD did not affect the diameter or the alignment of the original fibers (Fig. 2A). The existence of epoxy groups after pGMA coating by iCVD was confirmed by FT-IR analysis (Fig. 2B). Compared to PCL nanofibers, a small peak appeared at 910 nm1 for pGMA-coated PCL nanofibers (iCVD), which represent the epoxy group. However, the peak was reduced after FN immobilization because new bonds were formed (1064 nm1) and epoxy rings were broken (Fig. 2B, as indicated by arrow). Binding of amine-containing and fluorescein-conjugated chitosan (Chitosan-FITC) to the pGMA-coated PCL was visualized by fluorescent microscopic analysis (Fig. 2C). Conformal conjugation of chitosan-FITC to the pGMA-coated nanofibers confirmed the reactivity of the epoxy group of pGMA on nanofibers. With a short incubation time, pGMA-PCL gave two times higher fluorescence intensity than the control nanofibers (Fig. 2D).

3.3. PCL nanofiber in vitro degradation Polycaprolactone belongs to the class of aliphatic polyesters and is degraded by hydrolytic degradation [37]. PCL is a semi-crystalline polymer. PCL is reported to degrade slowly over a period of 2–4 years depending on its properties, such as molecular weight and crystallinity. To evaluate the degradation properties of PCL nanofiber, we used accelerated conditions using alkaline medium (1 N NaOH) to promote hydrolysis of the polyesters. As shown in Fig. 2E, our results reveal linear degradation profiles among the PCL, pGMA-coated PCL nanofibers (iCVD), and FN-immobilized PCL nanofiber scaffolds (FN). The PCL nanofiber showed a higher degradation rate (rate constant K = 10.438) compared to the pGMA-coated PCL nanofiber (K = 4.8458) and FN-immobilized PCL nanofiber (K = 4.2962). The pGMA-coated PCL nanofiber and the FN-immobilized PCL nanofiber showed only a 26.62% and 30.76% mass loss under accelerated conditions in 6 h, compared with a 65.3% mass loss for the control PCL nanofiber. These results may indicate that the iCVD treatment to the nanofibers delayed the degradation of the material compared to the non-coated nanofibers. We concluded that the degradation kinetics of iCVD-treated materials were considerably slower than non-iCVD treated nanofibers due to its greater hydrophobicity and crystallinity by pGMA under alkaline medium conditions. Also iCVD coating of pGMA may have prevented the NaOH-dependent PCL degradation. 3.4. Characteristics of UCB-MSC-seeded FN-immobilized PCL nanofiber The influences of topography and artificial surface modulations were investigated by plotting the apparent contact angles. The contact angle of FN-immobilized PCL nanofiber turned out to be 106.0°, which can be interpreted as being more hydrophilic compared to those of PCL- and iCVD-coated PCL nanofibers, which were 133.5° and 131.7°, respectively. We further investigated cellular response on FN-immobilized PCL nanofibers. UCB-MSCs were seeded onto FN-immobilized PCL nanofiber at 2  105 cells per scaffold. Even though the FN-immobilized PCL nanofiber is more hydrophilic than others, infiltration of cells into the interior of the nanofiber was limited due to hydrophobic surface and narrow pore size (data not shown). Changes in cellular morphology usually result

Fig. 1. Schematic illustration of fabrication of functionalized nanofibers for MI. (A) PCL nanofibers were prepared and coated with pGMA by iCVD. FNs were immobilized onto the pGMA-PCL nanofibers. (B) UCB-MSCs were seeded onto the FN-immobilized PCL nanofibers and transplanted onto a MI model rat.

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Fig. 2. Characterization of functionalized nanofibers. (A) SEM images of control nanofibers and pGMA-coated nanofibers. Scale bar = 50 lm. (B) FTIR spectra of control PCL, pGMA-coated nanofiber by iCVD (iCVD), and FN-immobilized nanofiber (FN). (C) Fluorescence images of control PCL nanofibers and pGMA nanofibers immobilized with chitosan-FITC. (D) Fluorescence intensity measurement of control PCL nanofibers and pGMA nanofibers immobilized with chitosan-FITC, n = 10. ⁄P < 0.05 vs. pGMA group. (E) In vitro degradation rate of control PCL, pGMA-coated nanofibers by iCVD (iCVD) and FN-immobilized nanofibers (FN).

from varying cell–substrate interactions leading to cytoskeletal reorganizations. SEM examination showed that the morphology of UCB-MSCs adhering to the PCL scaffolds was altered after the FN coating (Fig. 3B). The cells on the PCL nanofiber showed round cellular morphology with abundant cellular pods. High hydrophobicity of PCL fibers inhibits the cellular protrusion and microfibrillar adhesion to the nanofiber. However, FN-immobilized nanofibers provided the optimal microenvironment for UCB-MSCs attachment. Actin cytoskeletal images of UCB-MSCs cultured on FN-immobilized PCL nanofiber indicated that the UCB-MSCs adhered most to the FN-immobilized PCL nanofiber with fibroblastic morphology (Fig. 3C).

3.5. Cell adhesion, proliferation and viability in scaffold For the number of cell attachments, 2  105 cells were seeded in each nanofiber construct and cultured for 24 h prior to counting cell numbers. The total number of cells attached per area (0.1 mm2) was doubled in FN-immobilized nanofibers compared to PCL nanofibers (Fig. 4A). For the proliferation study, 5  105 cells were seeded in each nanofiber and cultured for 72 h prior to performing a Quant-iT™ PicoGreenÒdsDNA Assay. All nanofiber constructs were digested in papainase solution to extract the DNA out of the cell. The proliferation rate, assessed by DNA content, was significantly increased in FN-immobilized nanofibers, which was 0.01 lg, as compared with non-coated control nanofibers, 0.004 and 0.006 lg, respectively (Fig. 4B). The cytotoxicity of the scaffolds was examined by a viability assay. Cell viability, when assessed by live/dead assay, was not statistically different among all the tested scaffolds over 7 days of cultivation and showed that only 10% of cultured cells were dead (Fig. 4C). These results indicate that FN-immobilized PCL nanofibers can support the cell attachment and survival.

3.6. Gene expression profile of the cells seeded onto scaffolds To determine the biological effects of immobilized FN on stem cells, we compared gene expression profiles of the human UCBMSCs from in vitro culture depending on the FN coating using PCR array analysis. UCB-MSCs were seeded on uncoated PCL nanofibers, pGMA-coated PCL nanofibers and FN-immobilized PCL nanofibers. Our result showed that the expression levels of some cellular function-related genes were increased in the UCB-MSCs cultured on FN-immobilized PCL nanofibers in comparison with uncoated PCL scaffolds (Fig. 5). Some of these elevated genes are known to be involved in regulating cellular functions of stem cells, including stem cell homing, anti-apoptosis, anti-inflammation, anti-scarring and angiogenesis. Specifically, the gene expression levels of the markers for stem cell homing (FGF-2), anti-apoptosis (IGF-1 and LIF), anti-inflammation (IL-6), anti-fibrosis (MMP-2) and angiogenesis (VEGF-A, KDR, JAG1, ANPEP, TGFb-1 and vWF) were higher in the FN-immobilized PCL nanofibers than those in uncoated PCL scaffolds. These results indicate that the putative paracrine effects of UCB-MSCs cultured on the FN-immobilized PCL nanofibers may be superior to that in uncoated PCL scaffolds for the treatment of MI. All the results in vitro taken together indicate that FN appears to be a favorable coating agent for MSC transplantation. Therefore, FN-immobilized PCL nanofibers were used for further in vivo experiments.

3.7. Assessment of cardiac functions by echocardiography To evaluate the effect of the combined cell and scaffold therapy on LV remodeling and cardiac function, echocardiography was performed. LVIDd, LVIDs, EF and FS were quantified at the baseline before the left anterior descending coronary artery (LAD) ligation, 1 week after LAD ligation before implantation and 4 weeks after

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Fig. 3. Contact angle and morphology of the cell on nanofibers. (A) Contact angle measurements of PCL nanofibers, pGMA-coated nanofibers by iCVD (iCVD) and FNimmobilized pGMA-coated nanofibers (FN). (B) SEM images of UCB-MSC morphology after 24 h in vitro on nanofibers, control PCL, pGMA-coated nanofibers by iCVD (iCVD) and FN-immobilized nanofibers (FN). Circle signifies single-cell morphology and arrow indicates linear alignment of the fiber. Scale bar = 20 lm. (C) Immunofluorescent analysis of UCB-MSC morphology on each type of nanofiber shown with autofluorescent (green) and nuclei with DAPI (blue). Scale bar = 50 lm.

Fig. 4. Cell adhesion and viability in scaffold. (A) Initial cellular adhesion rate of UCB-MSCs onto nanofibers were measured by counting cell numbers per area (n = 3) after 24 h of post seeding. (B) Proliferation rate of UCB-MSCs onto nanofibers were measured by DNA quantification assays (n = 3) after 72 h of post seeding. (C) Viability of the each group was measured by live/dead viability/cytotoxicity kit (n = 3) after 48 h of post seeding. ⁄P < 0.05 vs. iCVD group. #P < 0.05 vs. FN group. All values represent mean ± SEM.

implantation. EF and FS values were within the normal range in all rats before MI induction, but the heart function was significantly compromised in all rats 1 week after infarction (Fig. 6A). To assess the functional effect of cardiac patches with human UCB-MSCs on FN-immobilized PCL nanofibers, LVIDd, LVIDs, EF and FS were calculated and compared 4 weeks after the treatment. The UCB-MSC/FN-PCL group (human UCB-MSCs on FN-immobilized PCL nanofibers), showed attenuation of LV dilation and dysfunction compared to the control and FN-PCL groups (Fig. 6B). LVIDd values were similar in all experimental groups. However, the LVIDs value in the UCB-MSC/FN-PCL group (6.47 ± 0.95 mm) was significantly smaller than that in the control group (7.93 ± 0.33 mm), although the non-cell seeded FN-PCL group (8.02 ± 1.42 mm) was not significantly different than the control group. EF in the UCB-MSC/FN-PCL group (60.2 ± 7.87%) was significantly higher than those in the control (41.82 ± 8.23%) and FN-PCL (38.54 ± 7.43%) groups. Similarly, FS in the UCB-MSC/FN-PCL group

(28.9 ± 5.2%) was also significantly higher than those in tehe control (18.26 ± 4.43%) and FN-PCL (16.56 ± 3.68%) groups.

3.8. Histological and histomorphometric analysis Histological and morphometric analyses were performed to identify LV remodeling after Masson’s trichrome staining of heart sections at 4 weeks. Masson’s trichrome staining shows that infarcted myocardium with fibrosis appears blue, while viable myocardium appears red (Fig. 7A). The MI size was significantly reduced in the UCB-MSC/FN-PCL group as compared to the control and FN-PCL groups (Fig. 7B). In addition, the percentage of ventricular fibrosis was also significantly reduced in the UCB-MSC/FN-PCL group as compared to the control and FN-PCL groups (Fig. 7B). The scar thickness of the UCB-MSC/FN-PCL group was significantly thicker than those of the control and FN-PCL groups (Fig. 7B).

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4. Discussion

Fig. 5. RT2 profiler PCR array data analysis clustergram of the expression of 84 key genes involved in human MSCs. cDNA expression profiling in human MSCs (RT2 Profiler). Clustergram generated from average cDNA Ct values of individual groups of samples. For each group, average Ct values were normalized to a panel of different housekeeping genes. The range of 2DCt values is given below the color scale.

3.9. Survival of transplanted cells We have shown in vitro that our FN-immobilized PCL nanofibers support UCB-MSC survival. Thus, we tested whether the cells survived in vivo for 4 weeks by immunohistochemical staining with human cell marker lamin A+C antibody. No lamin A+C-positive cells were detected in the non-cell-seeded control and FNPCL groups. In the UCB-MSC/FN-PCL group implanted with UCBMSCs, lamin A+C-positive cells were detected in the scaffolds or the boundary between the scaffold and the underlying myocardium (Fig. 8). However, these lamin A+C-positive cells remained isolated from each other and did not penetrate into the adjacent myocardium or infarction area. This result may indirectly indicate that the stem cells prevented cardiac damage, possibly by secreting paracrine factors rather than the differentiation of the implanted cells. 3.10. Blood vessel density analysis The infarcted myocardium was immunohistochemically stained for vWF and the blood vessels in the infarcted region were quantified. Vessel density of the infarcted area in the UCB-MSC/FN-PCL group was significantly higher than those in the control and FNPCL groups (Fig. 9). This result suggests that the UCB-MSCs in the FN-immobilized PCL nanofiber promote angiogenesis by secreting or stimulating angiogenic factors such as VEGF, as demonstrated in the in vitro PCR array.

In the present study, we investigated the effects of UCB-MSCseeded FN-immobilized PCL nanofibers on improvements in cardiac function and attenuation of LV fibrotic remodeling in a rat MI model. Human UCB-MSCs were seeded onto FN-immobilized PCL nanofibers and implanted onto the epicardial surface over the infarcted area and adjacent infarction border zones. 4 weeks after transplantation, the cell-seeded UCB-MSC/FN-PCL group had significantly improved LV function, attenuated LV remodeling and increased neovascularization compared to the control and FN-PCL groups. These results can be explained as the outcome of the cardio-protective effects produced by the combined stem cell and scaffold therapy. The FN-immobilized PCL scaffold alone was not able to prevent cardiac dysfunction and LV remodeling after MI. This suggests that the FN-immobilized PCL nanofiber acts as a vehicle for cell delivery to improve the efficacy of stem cell therapy for MI. In this study, UCB was used to obtain MSCs, and we identified the cardio-protective potential of these cells. These cells, when seeded on the FN-immobilized PCL nanofibers, expressed survival and angiogenic factors (Fig. 5). In the present study, the FN-immobilized PCL nanofiber alone did not lead to angiogenesis and the improvement of MI, but UCB-MSC implantation with our FNimmobilized PCL nanofiber induced favorable results. It can be concluded from these findings that the observed improvement after implantation appear by the applied UCB-MSCs not by the FN-immobilized PCL nanofiber itself. The main advantages of UCB-MSCs are that these cells can be harvested painlessly without the risk of patient morbidity. Previous studies have confirmed that the UCB-MSCs are immune-privileged cells with surface characteristics capable of overcoming rejection [13]. Thus, these UCB-MSCs have the potential to be utilized as off-the-shelf allogeneic cells for regenerative medicine. Our findings suggest that UCB-MSCs can be stem cell candidates for cardiac therapy, and can be used in place of BM-MSCs because these cells have the cardio-protective potential and many advantages for clinical application. The application of biomaterials with a variety of cell types has shown an improvement of cardiac function in animal model [38]. Among various polymeric biomaterials utilized in cardiac tissue engineering applications, PCL nanofibers used in this study are especially appropriate for use in the treatment of MI since they exhibit elasticity. In this study, the maximum strength of PCL nanofibers on the entire scaffold was 33.48 MPa with breaking stain at 7.9%. Furthermore, iCVD treatment did not significantly alter the mechanical properties of our fibers (data not shown). When considering cardiac tissue engineering, the scaffold implanted onto the cardiac surface needs to be elastic to resist injury from continuous contraction. In addition, biomaterials for therapeutic cardiac tissue engineering should allow appropriate cellular interaction. ECM protein can serve as a scaffold for cellular interaction including cell attachment, migration and proliferation [39]. In the present study, we immobilized FN as an ECM protein on PCL scaffolds to increase cell transfer efficiency and attachment. FN could be an excellent candidate as it is normally present in the heart and plays a pivotal role in stem cell behavior, including adhesion and proliferation. In a previous study, stem cells attached best to FN in comparison with other ECM molecules [30]. Similarly, in the present study, FN stimulated in vitro UCB-MSC attachment. Electrospinning is a versatile method to create a biomimicking microenvironment similar to that of natural ECM for enhanced cell adhesion and tissue growth. The proposed scaffolds with aligned nanofiber topography are desirable for cardiac tissue engineering due to its ability to induce cellular phenotype found in native cardiac tissue. Previously established methods to surface functionalize nanofibers have been limited to

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Fig. 6. Evaluation of cardiac function by echocardiography. (A) Echocardiography measurements performed before MI and 7 days after MI show that MI induces an increase in LV remodeling and LV chamber dimensions, leading to a decrease in EF and fractional FS. (B) In echocardiographic examination at 4 weeks after treatment, control and FN-PCL groups show LV dilation and significant LV wall thinning. The UCB-MSC/FN-PCL group shows limited LV remodeling and preserved contractility. EF and FS as LV systolic function are significantly higher in the UCB-MSC/FN-PCL group compared to the control and FN-PCL groups. ⁄P < 0.05 vs. control group. #P < 0.05 vs. FN-PCL group. All values represent mean ± SEM.

physical ECM coating. However, we hypothesized that permanent immobilization of natural proteins on the nanofiber surface is required for uninterrupted cell recognition that is essential for stem cell function and differentiation. In the present study, some of the engrafted UCB-MSCs were identified within the seeded scaffolds and the boundary between the scaffold and the underlying myocardium by immunohistochemical examination. However, any migration or penetration of these cells toward the host cardiac tissue was not detected. Nevertheless, MSC-seeded FN-immobilized PCL nanofibers efficiently prevented LV remodeling and dysfunction. These findings suggest that the positive results seen in the UCB-MSC/FN-PCL group compared with the control and FN-PCL groups might be related to paracrine activities. Likewise, previous accumulating evidence has suggested that the main beneficial effects derived from implanted

cells are through the paracrine effects of cytokines secreted by the implanted cells and even a few cells could activate different regeneration pathways such as cell survival, stem cell homing, angiogenesis and matrix remodeling if they produce crucial bioactive mediators [2,6,14,40]. To identify potential paracrine effects of UCB-MSCs seeded in FN-immobilized PCL nanofibers, we evaluated gene expressions of multiple paracrine factors using a PCR gene array study. The expressions of some angiogenic factors such as VEGF-A, KDR, JAG1, ANPEP, TGFb-1 and vWF were significantly increased in the UCB-MSCs on the FN-immobilized PCL nanofibers compared with them on uncoated PCL scaffolds. These findings suggest that MSCs seeded in the FN-immobilized PCL nanofibers could modulate the expression of multiple angiogenic cytokines with the potential to promote neovascularization. Neovascularization contributed to the prolonged cell survival in the FN-immobilized

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Fig. 7. Evaluation of MI size, LV fibrosis and scar thickness. (A) Representative images of heart sections stained with Masson’s trichrome show fibrosis and wall thinning in the infarcted area. Fibrotic areas are colored in blue and viable myocardium in red. (B) Infarct size, percentage of fibrosis and wall thickness are statistically compared between different groups. ⁄P < 0.05 vs. control group. #P < 0.05 vs. FN-PCL group. All values represent mean ± SEM.

Fig. 8. Representative immunostaining images from three transplant groups after 4 weeks using lamin A+C antibody to visualize the transplanted cells. Positive staining for lamin A+C is shown in only the UCB-MSC/FN-PCL group, but not in the control and FN-PCL groups. Scale bar = 100 lm.

PCL nanofibers, and this may induce increased collateral flow and contractility, and consequentially improvement of cardiac function. In addition, stem cell homing (FGF-2), anti-apoptosis (IGF-1 and LIF), anti-inflammation (IL-6) and anti-fibrosis (MMP-2) genes showed significantly higher expression levels in the

FN-immobilized PCL nanofiber group than in the uncoated scaffold groups. The aforementioned results indirectly indicate that UCBMSCs seeded in the FN-immobilized PCL nanofiber can stimulate a profound neovascularization and attenuation of LV remodeling through their producing cytokines. The ECM provides cells with

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Fig. 9. Vessel density in the MI sites. Representative images of myocardial sections from control (A), FN-PCL (B) and UCB-MSC/FN-PCL (C) groups show immunohistochemical vWF staining of the infarcted myocardium. The vascular endothelial cells in the blood vessels are colored in brown (arrows). (D) Blood vessel densities were higher in the UCB-MSC/FN-PCL group than in the control and FN-PCL groups. Scale bar = 200 lm. ⁄P < 0.05 vs. control group. #P < 0.05 vs. FN-PCL group. All values represent mean ± SEM.

not only structural support but also biochemical and physical cues that regulate cell phenotype [41]. FN used in this study as the ECM component also plays a role in regulating the cell phenotype, and interacts via integrin receptors, which regulates the cellular cytoskeleton and cell behavior [42]. Therefore, the interaction between integrin and FN may lead to improving the putative paracrine effect of stem cells, although we did not yet verify and understand the mechanisms of FN-mediated cellular signaling.

5. Conclusion The FN-immobilized PCL nanofiber used in this study enhanced cellular adhesion, which makes it an attractive carrier for cell transplantation in the treatment of MI. Furthermore, the transplantation of UCB-MSCs with FN-immobilized PCL nanofiber improved cardiac function, prevented LV remodeling and stimulated angiogenesis in the rat MI model. Therefore, the combination of FNimmobilized PCL nanofibers and UCB-MSCs may be a promising strategy for future cardiac clinical application.

Acknowledgements This work was supported by the National Research Foundation (NRF) funded by the Ministry of Science, ICT & Future Planning, Republic of Korea (2012M3A9C6049716 and 20110019355).

Appendix A. Figures with essential color discrimination Certain figures in this article, particularly Figs. 1–3 and 5–9, are difficult to interpret in black and white. The full color images can be found in the on-line version, at http://dx.doi.org/10.1016/ j.actbio.2014.03.013.

References [1] Lloyd-Jones D, Adams RJ, Brown TM, Carnethon M, Dai S, De Simone G, et al. Heart disease and stroke statistics—2010 update: a report from the American Heart Association. Circulation 2010;121:e46–e215. [2] Segers VF, Lee RT. Stem-cell therapy for cardiac disease. Nature 2008;451:937–42. [3] Karam J-P, Muscari C, Montero-Menei CN. Combining adult stem cells and polymeric devices for tissue engineering in infarcted myocardium. Biomaterials 2012;33:5683–95. [4] Müller-Ehmsen J, Peterson KL, Kedes L, Whittaker P, Dow JS, Long TI, et al. Rebuilding a damaged heart long-term survival of transplanted neonatal rat cardiomyocytes after myocardial infarction and effect on cardiac function. Circulation 2002;105:1720–6. [5] Savoye C, Equine O, Tricot O, Nugue O, Segrestin B, Sautière K, et al. Left ventricular remodeling after anterior wall acute myocardial infarction in modern clinical practice (from the REmodelage VEntriculaire [REVE] study group). Am J Cardiol 2006;98:1144–9. [6] Lu F, Zhao X, Wu J, Cui Y, Mao Y, Chen K, et al. MSCs transfected with hepatocyte growth factor or vascular endothelial growth factor improve cardiac function in the infarcted porcine heart by increasing angiogenesis and reducing fibrosis. Int J Cardiol 2012. doi: 10.1016/j.ijcard.2012.06.052. Available from: http://www.ncbi.nlm.nih.gov/pubmed/22981278. [7] Kim B-O, Tian H, Prasongsukarn K, Wu J, Angoulvant D, Wnendt S, et al. Cell transplantation improves ventricular function after a myocardial infarction: a preclinical study of human unrestricted somatic stem cells in a porcine model. Circulation 2005;112:I-96–I-104. [8] Yoon SJ, Fang YH, Lim CH, Kim BS, Son HS, Park Y, et al. Regeneration of ischemic heart using hyaluronic acid-based injectable hydrogel. J Biomed Mater Res Appl Biomater 2009;91:163–71. [9] Orlic D, Kajstura J, Chimenti S, Jakoniuk I, Anderson SM, Li B, et al. Bone marrow cells regenerate infarcted myocardium. Nature 2001;410:701–5. [10] Caplan AI, Dennis JE. Mesenchymal stem cells as trophic mediators. J Cell Biochem 2006;98:1076–84. [11] Wollert KC, Drexler H. Clinical applications of stem cells for the heart. Circ Res 2005;96:151–63. [12] Miyahara Y, Nagaya N, Kataoka M, Yanagawa B, Tanaka K, Hao H, et al. Monolayered mesenchymal stem cells repair scarred myocardium after myocardial infarction. Nat Med 2006;12:459–65. [13] Lee OK, Kuo TK, Chen W-M, Lee K-D, Hsieh S-L, Chen T-H. Isolation of multipotent mesenchymal stem cells from umbilical cord blood. Blood 2004;103:1669–75. [14] Mathieu E, Lamirault G, Toquet C, Lhommet P, Rederstorff E, Sourice S, et al. Intramyocardial delivery of mesenchymal stem cell-seeded hydrogel preserves cardiac function and attenuates ventricular remodeling after myocardial infarction. PLoS ONE 2012;7:e51991.

B.-J. Kang et al. / Acta Biomaterialia 10 (2014) 3007–3017 [15] Marsano A, Maidhof R, Luo J, Fujikara K, Konofagou EE, Banfi A, et al. The effect of controlled expression of VEGF by transduced myoblasts in a cardiac patch on vascularization in a mouse model of myocardial infarction. Biomaterials 2013;34:393–401. [16] Siepe M, Giraud M-N, Pavlovic M, Receputo C, Beyersdorf F, Menasché P, et al. Myoblast-seeded biodegradable scaffolds to prevent post-myocardial infarction evolution toward heart failure. J Thorac Cardiovasc Surg 2006;132:124–31. [17] Piao H, Kwon J-S, Piao S, Sohn J-H, Lee Y-S, Bae J-W, et al. Effects of cardiac patches engineered with bone marrow-derived mononuclear cells and PGCL scaffolds in a rat myocardial infarction model. Biomaterials 2007;28:641–9. [18] Laflamme MA, Murry CE. Regenerating the heart. Nat Biotechnol 2005;23:845–56. [19] Toma C, Pittenger MF, Cahill KS, Byrne BJ, Kessler PD. Human mesenchymal stem cells differentiate to a cardiomyocyte phenotype in the adult murine heart. Circulation 2002;105:93–8. [20] Venugopal JR, Prabhakaran MP, Mukherjee S, Ravichandran R, Dan K, Ramakrishna S. Biomaterial strategies for alleviation of myocardial infarction. J R Soc Interface 2012;9:1–19. [21] Araña M, Peña E, Abizanda G, Cilla M, Ochoa I, Gavira JJ, et al. Preparation and characterization of collagen based ADSC-carrier sheets for cardiovascular application. Acta Biomater 2013;9:6075–83. [22] Liu Z, Wang H, Wang Y, Lin Q, Yao A, Cao F, et al. The influence of chitosan hydrogel on stem cell engraftment, survival and homing in the ischemic myocardial microenvironment. Biomaterials 2012;33:3093–106. [23] Zimmermann WH, Melnychenko I, Eschenhagen T. Engineered heart tissue for regeneration of diseased hearts. Biomaterials 2004;25:1639–47. [24] Lin Y-D, Luo C-Y, Hu Y-N, Yeh M-L, Hsueh Y-C, Chang M-Y, et al. Instructive nanofiber scaffolds with VEGF create a microenvironment for arteriogenesis and cardiac repair. Sci Transl Med 2012;4:146ra09. [25] Prabhakaran MP, Kai D, Ghasemi-Mobarakeh L, Ramakrishna S. Electrospun biocomposite nanofibrous patch for cardiac tissue engineering. Biomed Mater 2011;6:055001. [26] Ishii O, Shin M, Sueda T, Vacanti JP. In vitro tissue engineering of a cardiac graft using a degradable scaffold with an extracellular matrix-like topography. J Thorac Cardiovasc Surg 2005;130:1358–63. [27] Karikkineth BC, Zimmermann WH. Myocardial tissue engineering and heart muscle repair. Curr Pharm Biotechnol 2013;14:4–11. [28] Yang F, Cho S-W, Son SM, Hudson SP, Bogatyrev S, Keung L, et al. Combinatorial extracellular matrices for human embryonic stem cell differentiation in 3D. Biomacromolecules 2010;11:1909–14. [29] Tate MC, Shear DA, Hoffman SW, Stein DG, Archer DR, LaPlaca MC. Fibronectin promotes survival and migration of primary neural stem cells transplanted into the traumatically injured mouse brain. Cell Transplant 2002;11:283–95.

3017

[30] Van Dijk A, Niessen H, Ursem W, Twisk J, Visser F, van Milligen F. Accumulation of fibronectin in the heart after myocardial infarction: a putative stimulator of adhesion and proliferation of adipose-derived stem cells. Cell Tissue Res 2008;332:289–98. [31] Wijelath ES, Rahman S, Murray J, Patel Y, Savidge G, Sobel M. Fibronectin promotes VEGF-induced CD34 cell differentiation into endothelial cells. J Vasc Surg 2004;39:655–60. [32] Seo Y, Yang S-R, Jee MK, Joo EK, Roh K-H, Seo M-S, et al. Human umbilical cord blood-derived mesenchymal stem cells protect against neuronal cell death and ameliorate motor deficits in Niemann Pick type C1 mice. Cell Transplant 2011;20:1033–47. [33] Park S-B, Yu K-R, Jung J-W, Lee S-R, Roh K-H, Seo M-S, et al. BFGF enhances the IGFs-mediated pluripotent and differentiation potentials in multipotent stem cells. Growth Factors 2009;27:425–37. [34] Kim M-J, Lee B, Yang K, Park J, Jeon S, Um SH, et al. BMP-2 peptidefunctionalized nanopatterned substrates for enhanced osteogenic differentiation of human mesenchymal stem cells. Biomaterials 2013;34:7236–46. [35] Han TH, Lee K, Park JB, Ahn D, Park J-H, Kim D-Y, et al. Reduction in synaptic GABA release contributes to target-selective elevation of PVN neuronal activity in rats with myocardial infarction. Am J Physiol Regul Integr Comp Physiol 2010;299:R129–39. [36] Matthew B. Injected versus oral cyclosporine for human neural progenitor grafting in rats. J Stem Cell Res Ther 2012. doi: 10.4172/2157-7633.S10-003. Available from: http://www.omicsonline.org/2157-7633/2157-7633-S10003.pdf. [37] Lam CX, Savalani MM, Teoh SH, Hutmacher DW. Dynamics of in vitro polymer degradation of polycaprolactone-based scaffolds: accelerated versus simulated physiological conditions. Biomed Mater 2008;3:034108. [38] Christman KL, Lee RJ. Biomaterials for the treatment of myocardial infarction. J Am Coll Cardiol 2006;48:907–13. [39] Wei H-J, Chen C-H, Lee W-Y, Chiu I, Hwang S-M, Lin W-W, et al. Bioengineered cardiac patch constructed from multilayered mesenchymal stem cells for myocardial repair. Biomaterials 2008;29:3547–56. [40] Fedak PW. Paracrine effects of cell transplantation: modifying ventricular remodeling in the failing heart. Semin Thorac Cardiovasc Surg 2008;20:87–93 [Summer]. [41] Philp D, Chen SS, Fitzgerald W, Orenstein J, Margolis L, Kleinman HK. Complex extracellular matrices promote tissue-specific stem cell differentiation. Stem Cells 2005;23:288–96. [42] Pimton P, Sarkar S, Sheth N, Perets A, Marcinkiewicz C, Lazarovici P, et al. Fibronectin-mediated upregulation of a5b1 integrin and cell adhesion during differentiation of mouse embryonic stem cells. Cell Adh Migr 2011;5:73–82.