BMEN 5325 UONT Health & Medical Stem Cell Fate Using Nano Fibrous

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Electrospinning Nanofibers for Neural Tissue Engineering Term paper for BMEN 5325 by name Department of Biomedical Engineering University of North Texas April 23rd, 2021 Introduction: 8.5/10 Research Summary: 10/10 Critical Review: 45/50 Recommendation: 9.5/10 Conclusion: 5/5 Structure: 9/10 Format: 3/5 Total: 90/100 1. Introduction Neural tissue is primarily composed of neurons and supporting glial cells. The complexity and specialization of the nervous system makes developing neural scaffolds for neural tissue trauma and diseases a challenging task for scientists and clinicians. Peripheral nerves have a limited capacity for regeneration following physical damage and current treatment options such as nerve grafting has limitations [1]. Autologous grafts involve the harvesting and implantation of a patient-derived donor nerve. This is mostly reserved for large nerve defects and lead to functional recovery [1]. However, this procedure requires multiple surgeries, leaves the donor nerve site nonfunctional, and there is a limited availability of donor nerves. Allografts are nerves harvested from other humans, or animals, but are also limited due to possible immune system rejection and disease related complications. Thus, the need for a nano engineered biomimetic neural structure is imperative to overcome the limitations of nerve grafting [1]. The application of nanotechnology in neural tissue engineering has great potential to overcoming the limitations seen with nerve grafting. By fabricating an implantable and biodegradable neural scaffold seeded with a variety of cellular or protein therapies, the possibility of complete nerve regeneration may be feasible in the future. A promising avenue of research in neural tissue engineering is the process of electrospinning polymer nanofibers. Electrospinning nanofibers is a relatively simple and versatile procedure and has been applied successfully to a variety of tissue engineering studies. Figure 1 shows a typical apparatus for electrospinning which consist of three main components: a high voltage power supply, a syringe pump and an electroconductive collecting surface. A wide range of natural or synthetic polymers can be used to fabricate the nanofibers. The electrospinning of nanofibers has tunable properties, and the morphology of an individual nanofiber is determined by the setting of certain variables. The syringe pump is the main control unit of the electrospinning process, and sets the parameters for the flow rate of the polymer solution/melts. The high voltage power supply is then applied to syringe needle, which produces the whipping jet motion of the pumped polymer solution and accumulates on the grounded collector. The small diameter of the electrospun nanofibers is a result of the whipping motion that exerts a strong axial force [1]. The polymer solution of choice must have the optimum viscoelastic properties to maintain its morphology during this whipping process. This continuous acceleration and stretching of the polymer solution results in the electrospun nanofibers, which are generally as thin as tens of nanometers in diameter [1]. The geometry and kinetics of the grounded collector also plays a crucial role in determining the overall orientation of the produced nanofibers. A rapidly rotating collector results in more aligned nanofibers, while a stationary collector produces randomly oriented nanofibers. Figure 1. Schematic view of electrospinning technique Nanofibers aligned into uniaxial arrays provide effective cues to direct and enhance neurite outgrowth, which is more advantageous than other materials such as hydrogels or nonaligned nanofibers. By manipulating the alignment, morphology, and stacking, the polymer nanofibers can be fabricated into a nerve guidance conduit (NGC) which can be fabricated to mimic the native extracellular matrix (ECM) of neural tissues. NGCs can also be seeded with a variety of bioactive molecules or growth factors to facilitate nerve regeneration, which make electrospun nanofibers a prime target of research in neural tissue engineering. A derivative of electrospinning is known as coaxial electrospraying. Coaxial electrospraying produces multilayer nanoparticles by introducing coaxial electrified jets [2]. Polymeric nanoparticles can be used to encapsulate, deliver, and release various therapeutic agents such as proteins, drugs, and gene therapies. Advantages of this process include high encapsulation efficiency, protection from bioactivity, and uniform size distribution [2]. This paper will discuss recent research in neural tissue engineering and summarize the ideal properties of a neural scaffold. It will also elaborate different strategies researchers have taken to fabricate nanofibers for neural tissue engineering applications. In addition, recommendations for future research will be discussed for the regenerating injured neural tissue using electrospun nanofibers. 2. The Ideal Neural Scaffold The ideal neural scaffold should be biocompatible and optimally improve cell adhesion, proliferation, migration, and axonal extension [3]. The scaffold should also provide the mechanical and chemical cues to promote new neural tissue formation. It should also be biodegradable in vivo, so there is no need for an additional removal surgery. Neural scaffolds can also be seeded with bioactive proteins and growth factors, but have ongoing limitations such as short-term retention, rapid half-life in circulation and rapid loss of biological activity in vivo [3]. Figure 2 shows the important properties for an ideal neural scaffold. Figure 2. The ideal properties of a tissue engineered neural construct [3]. 3. Aligned nanocomposite scaffold for neural regeneration To overcome the limitations seen in previous reports utilizing bioactive factor seeded neural scaffolds, Zhu et al, developed a sustained biodegradable core-shell nanospheres made of poly lactic-co-glycolic acid (PLGA), encapsulated with bovine serum albumin (BSA). This was done via a coaxial electrospraying technique. Coaxial electrospraying allows the fabrication of a controllable core-shell nanosphere with bioactive factors within the core and outer shell [3]. BSA is a large globular protein and was used as a nutrient to improve neural cell behavior [3]. Following the fabrication of the core-shell nanosphere, Zhu et al, electrospun polycaprolactone (PCL) microfibers to create the nanocomposite 3D scaffold that directed neural cell growth. In vitro analysis of the cell-scaffold interaction was performed using PC-12 cells. Figure 3 shows that PC-12 grew well on all the scaffolds used in this study, but cell proliferation was significantly higher in the PCL with BSA embedded nanospheres than PCL controls, and PCL scaffolds sprayed directly with BSA after 4 and 6 days. Figure 3. PC-12 cell proliferation in nanocomposite scaffold at 2, 4, and 6 days [3]. Confocal micrographs of PC-12 cells cultured with nerve growth factor (NGF) on aligned and random scaffolds with and without nanospheres are shown below in Figure 4. Outgrowth and extension of neurites were seen on both aligned and random scaffolds. Neuronal markers TuJ1 and MAP2 were stained to indicate neural differentiation of PC-12 cells. After seven days, all the scaffolds demonstrated differentiation. The orientation of the differentiated neurites extended along the axis of the aligned fibers, parallel with neighboring cells [3]. Axons on the randomly aligned fibers extended radially with no specific direction. This study showed that aligned fibrous scaffolds with topographical cues show a superior ability to direct neurite outgrowth and extension. Aligned fibrous neural conduits influenced endogenous neural repair mechanisms and are much more conducive to neurite growth with no need for additional exogenous growth factors. In addition, the aligned fibers benefit the formation of longitudinally oriented Bunger bands. The Bunger bands include aligned strands of Schwann cells and laminin, which are a key element in nerve repair [3]. This study demonstrated the advantages of aligned nanocomposite nanofibers over randomly aligned nanofibers. Figure 4. Confocal microscopy images of PC-12 cell line. Staining of MAP2 and TuJ1 detected PC-12 differentiation on various scaffolds following 7 days of culture. [3]. In another study, random and aligned nanofiber scaffolds were also fabricated from PCL. However, an emulsion electrospinning technique was used where BSA and NGF formed the core, while PCL formed the shell. Emulsion electrospinning has been developed to prepare core-shell structured nanofibers as drug delivery vehicles [4]. Random and aligned pure PCL, PCL-BSA-NGF, PCL-BSA, and PCL-NGF nanofibers were produced for comparison. Figure 4 shows a schematic illustration of the emulsion preparation and electrospinning set up. Figure 4. Schematic illustration of (A) emulsion preparation process, and the electrospinning set-up to produce (B) random, and (C) aligned nanofibers. Red = water phase. Green = oil phase [4]. PC-12 cells were cultured for eight days on the surface of each nanofiber created for this study. After 8 days of culture, some PC12 cells grown on NGF-added (R/A)-PCL, (R/A)-PCL-BSA scaffolds showed elongation. PC12 cells on (R/A)-PCL-NGF and (R/A)-PCL-NGF&BSA scaffolds projected the neurites [4]. These results suggest that NGF stimulated PC12 differentiation [4]. However, NGF released from the NGF encapsulated nanofibers were more effective on PC12 differentiation compared to NGF that was only added directly in the culture medium [4]. 4. Blended electrospun nanofibers In another study by Lins et al, they developed electrospun nanofiber scaffolds that showed similar structure to the ECM present in neural tissues. This group used poly(lactic acid) (PLA)/poly(lactide-b-ethylene glycol-b-lactide) block copolymer (PELA) and PLA/polyethylene glycol (PEG) as the blended polymer materials. An advantage for PLA in neural tissue engineering applications is its hydrolytic degradation kinetics [5]. This allows PLA to be eliminated as carbon dioxide and water in the Krebs’ cycle. The limitations of PLA include, slow biodegradation process, high stiffness, and is hydrophobic. To increase the hydrophilicity and reduce the brittleness of the PLA-based scaffold, soft PEG was blended with PLA to increase the biocompatibility of the scaffold. PEG has good biocompatibility and low toxicity and has been investigated for its treatment of injury to neuronal membranes [5]. PELA is a block copolymer that is good for tissue engineering applications because it possesses intermediate physicochemical characteristics and has a good balance between degradation rate and hydrophilicity. The PLA, PLA/PELA, and PLA/PEG-based membranes were prepared by electrospinning. This paper focused on the comparison between the blends of PLA with varying molecular weights of PEG and PELA. Following the electrospinning of the polymer blends, 3D interconnected fibers with smooth and round shapes were formed. The PLA/PELA2k and PLA/PEG2k blended nanofibers showed more homogenous morphologies than PLA, PLA/PELA20k and PLA/PEG20k blended nanofibers. To test the cell affinity of the electrospun fibers, monkey embryonic stem cells (ESCs) and neural stem cells (NSCs) were used. Monkey ESCs and NSCs share the same properties as human ESCs and NSCs, which made them an appropriate cell type for this study. NSCs were first cultured on the PLA, PLA/PELA, and PLA/PEG blended scaffolds, and immunostained for NSC marker SOX2 (Figure 6B). TAU-green fluorescent protein (TAU-GFP) allowed the visualization of cell morphology, which bind GFP to microtubules [5]. Neat PLA expressed SOX2 at 71%, and PLA/PELA20k expressed SOX2 at 67% (Fig 7C). The cell density on the PLA and PLA/PELA20k was similar as well at 541 and 494 cells/mm2 respectively. Cell density was quantified by counting the number of living cells per mm2. The PLA/PELA2k cell density was reduced at 339 cells/mm2 and SOX2 expression was lower at 57%. In contrast to the previously described fibers, the PLA/PEG2k and 20k resulted in very low cell density. These results indicated that PLA/PEG2K and 20K were not suitable scaffolds for neural cell maintenance and differentiation. Figure 6. ESC-NSC characterization upon culture on PLA, PLA/PEG2K, PLA/PEG20K, PLA/PELA2K, and PLA/PELA 20K scaffolds. (B) Immunostaining for TAU-GFP, SOX2, and Ki67 in EXS-NSCs cultured for 2 and 5 DIV on the different scaffolds. Scale bars, 50um (2DIV); 20 um (5 DIV). (C) Cell density and SOX-2 positive cell percentage after 2 DIV. DIV = days in vitro. [5] In addition, PGS, PMMA, and gelatin were electrospun to form blended nanofibers for neural tissue engineering. PGS is a biodegradable and elastic polymer, but uncured PGS is difficult to electrospin into nanofibers. PGS was modified by using atom transfer radical polymerization (ATRP) to synthesize PGS-based copolymers with MMA. PGS-PMMA was easily electrospun into nanofibers with a diameter of 167 +/- 33 nm [6]. Rat PC12 cells were seeded onto the PGS- PMMA/gelatin nanofibers and analyzed for its efficacy in nerve regeneration. As shown in Figure 8 below, the gelatin-containing PGS based nanofibers showed the greatest amount of PC-12 cell proliferation after 8 days. This was most likely due to the preferential proliferation of neuronal cell on nanofibers containing natural polymers [6]. Gelatin is a derivative of collagen, and present amide functional groups on the surface of the nanofibrous scaffolds, which provide chemical and biological cues for cell adhesion and growth [6]. This study indicated that by modifying PGS and synthesizing a new PGS-PMMA copolymer material, may be a promising approach for novel biomaterial applications. Figure 8. Proliferation of rat PC12 cells on PGS-PMMA/gelatin nanofibers as determined by alamarBlue. #p 3´) Collagen I F: TGGAGCAAGAGGCGAGAG Base(length bp) 121 R: CACCAGCATCACCCTTAGC Runx2 F: GCCTTCAAGGTGGTAGCCC 66 R: CGTTACCCGCCATGACAGTA Osteonectin F: AGGTATCTGTGGGAGCTAATC 224 R: ATTGCTGCACACCTTCTC Osteocalcin F: GCAAAGGTGCAGCCTTTGTG 80 R: GGCTCCCAGCCATTGATACAG β2M F: GCCTTCACCCCAGAGAAAGG R: GCGGTTGGGATTTACATGTTG This article is protected by copyright. All rights reserved 101 Figure 1 This article is protected by copyright. All rights reserved Figure 2 This article is protected by copyright. All rights reserved Figure 3 This article is protected by copyright. All rights reserved Figure 4 This article is protected by copyright. All rights reserved Figure 5 This article is protected by copyright. All rights reserved Figure 6 This article is protected by copyright. All rights reserved Figure 7 This article is protected by copyright. All rights reserved Xue et al. Stem Cell Research & Therapy (2017) 8:148 DOI 10.1186/s13287-017-0588-0 RESEARCH Open Access Polycaprolactone nanofiber scaffold enhances the osteogenic differentiation potency of various human tissue-derived mesenchymal stem cells Ruyue Xue1, Yuna Qian2, Linhao Li3, Guidong Yao1, Li Yang2 and Yingpu Sun1* Abstract Background: Polycaprolactone (PCL) has been regarded as a promising synthetic material for bone tissue engineering application. Owing to its unique biochemical properties and great compatibility, PCL fibers have come to be explored as a potential delivering scaffold for stem cells to support bone regeneration during clinical application. Methods: The human derived mesenchymal stem cells (MSCs) were obtained from umbilical cord (UC), bone marrow (BM), and adipose tissue (AD), respectively. The osteogenic differentiation potency of various human MSCs on this novel synthetic biomaterial was also investigated in vitro. Results: Here, we illustrated that those human UC-, BM-, and AD-derived MSCs exhibited fibroblast-like morphology and expressed characteristic markers. Impressively, PCL nanofiber scaffold could support those MSC adhesion and proliferation. Long-term culture on PCL nanofiber scaffold maintained the viability as well as accelerated the proliferation of those three different kinds of human MSCs. More importantly, the osteogenic differentiation potency of those human MSCs was increased significantly by culturing on PCL nanofiber scaffold. Of note, BM-derived MSCs demonstrated greater differentiation potency among the three kinds of MSCs. The Wnt/β-catenin and Smad3 signaling pathways contributed to the enhanced osteogenesis of human MSCs, which was activated consistently by PCL nanofiber scaffold. Conclusions: The utilization of PCL nanofiber scaffold would provide a great application potential for MSC-based bone tissue repair by enhancing the osteogenic differentiation of human MSCs. Keywords: Polycaprolactone, Mesenchymal stem cells, Osteogenic differentiation, Wnt/β-catenin signaling pathway, Smad3 Background Mesenchymal stem cells (MSCs) are a kind of cell population with multi-differentiation potential, which were first identified in bone marrow (BM)-MSCs more than 40 years ago [1]. It was soon noted that MSCs display not only regenerative properties, but also remarkable differentiation ability. Growing evidence showed that human MSCs are able to differentiate into derivatives of all germ layers including bone, cartilage, adipose tissue, cardiomyocyte, and neurocyte [2, 3]. Those properties have * Correspondence: [email protected] 1 Reproductive Medical Center, First Affiliated Hospital of Zhengzhou University, Zhengzhou 450052, China Full list of author information is available at the end of the article endowed MSC-based therapy with a great potential therapeutic strategy for wound healing and tissue regeneration, which is also thanks to few ethical issues and low risk of tumorigenesis in contrast to embryonic stem cells and induced pluripotent stem cells [4–6]. Even though bone marrow is referred to as the most extensively used source for MSC isolation, the frequency of BMMSCs has been estimated to be in the order of 0.001– 0.01% of total uncleated cells in bone marrow [7, 8]. Instead, several other tissue-derived MSCs have been identified and investigated. Among them, umbilical cord (UC)-MSCs and adipose tissue (AD)-MSCs exert great potential as ideal allergenic and autologous cell types for clinical application owing to their easy cell isolation and © The Author(s). 2017 Open Access This article is distributed under the terms of the Creative Commons Attribution 4.0 International License (, which permits unrestricted use, distribution, and reproduction in any medium, provided you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license, and indicate if changes were made. The Creative Commons Public Domain Dedication waiver ( applies to the data made available in this article, unless otherwise stated. Xue et al. Stem Cell Research & Therapy (2017) 8:148 sufficient frequency of clean sources [9–11]. Increasing studies have indicated that BM-MSCs, UC-MSCs, and AD-MSCs are able to differentiate into bone tissues in vitro and in vivo. But there still exist concerns regarding the osteogenic efficiency of those most widely used types of MSCs during clinical application. To address this concern, various synthetic biomaterials have been employed to support the adhesion and accelerate the osteogenic differentiation in MSC-based regenerative therapy for treating bone impairment. Notably, the use of polycaprolactone (PCL) for tissue engineering increased significantly over the past decade [12–14]. PCL, a biodegradable polyester material, has been widely exploited in several kinds of implants, adhesion barriers, and drug delivery devices. Since the 1980s, this material has been approved by the Food and Drug Administration (FDA), which is thanks to its excellent electrospinnability, good mechanical features, good compatibility, and biodegradable properties. But the effects of this biomaterial on the osteogenic potency of various human MSCs were still elusive. Here we used a novel scaffold consisting of PCL-based electrospun nanofibers and investigated the bioactivity and osteogenic potency of three different tissue-derived human MSCs on this synthetic biomaterial. It was illustrated that human-derived MSCs obtained from umbilical cord, bone marrow, and adipose tissue all could form colonies, show fibroblast-like morphology, and be characterized by expressing CD90, CD105, and CD73 but not CD31, CD34, and CD45. Long-term culture on this novel synthetic scaffold did not reduce human MSC viability. Moreover, the proliferation potency of human UC-MSCs, BM-MSCs, and AD-MSCs was increased after adhesion on the surface of the PCL nanofiber scaffold. We further investigated the osteogenic differentiation of human UCMSCs, BM-MSCs, and AD-MSCs respectively with and without PCL nanofiber scaffold, then we found that this synthetic scaffold indeed enhanced the osteogenic potency of those three different tissue-derived human MSCs in vitro. In addition, human BM-MSCs exert the greatest increase of osteogenic potency among those three kinds of human MSCs when cultured on PCL nanofiber scaffold. The Wnt/β-catenin and Smad3 signaling pathways contributed to the osteogenesis of human MSCs, which was activated consistently by the PCL nanofiber scaffold. Therefore, our study indicated that PCL nanofiber scaffold exerts a great compatibility with human MSCs, which could maintain cell viability as well as promote the proliferation potency. More importantly, the osteogenic potency of human UC-MSCs, BM-MSCs, and AD-MSCs could be enhanced by culturing on PCL nanofiber scaffold, implying the great potential of this novel synthetic biomaterial combined with MSC-based therapy for bone repair in clinical application. Page 2 of 9 Methods Polycaprolactone (PCL) PCL is a kind of polymer with mechanical properties, miscibility and biodegradability. PCL is an aliphatic linear polyester, with a glass transition temperature of about −60 °C and a melting point of 55–60 °C, depending on the degree of crystallinity, which in turn is dictated by the molecular weight and the scaffold fabrication process. It is also a kind of biocompatible, absorbable and low-cost synthetic polymer. The way to obtain high molecular weight PCL is mainly dependent of the ring-opening polymerization of ε-caprolactone (Fig. 2a). Due to its semi-crystalline and hydrophobic nature, the mechanical property of PCL is suitable for a variety of applications, and this material demonstrates a slow degradation rate (2–4 years). PCL has been clinically used as a slow-release drug delivery device and suture material that has been approved by the FDA since the 1980s. In this study, the PCL nanofiber scaffold was fabricated by the Key Laboratory of Biorheological Science and Technology, Ministry of Education, Bioengineering College, Chongqing University. Culture of human-derived MSCs UC-MSCs (c-12971), BM-MSCs (c-12974), and AD-MSCs (c-12977) were purchased from Promocell (Miaotong (Shanghai) Biological Science & Technology Co., Ltd. Shanghai, China). Cells were cultured with Dulbecco’s modified Eagle’s medium (DMEM) supplemented with 10% FBS, 10 ng/mL bFGF and 5% penicillin/streptomycin and incubated at 37 °C and 5% CO2. As shown in the manufacturer’s instructions, the three kinds of tissue were obtained from healthy volunteers after informed consent according to the Helsinki declaration. MSC surface marker identification The MSCs surface markers were examined by flow cytometry using PE-conjugated anti-human CD45, CD34, CD31, CD73, CD90, and CD105 antibodies were used for staining human mesenchymal stem cells according to the manufacturer’s instructions. All antibodies were purchased from BD Biosciences (Franklin Lakes, NJ, USA). Briefly, 1 × 105 MSCs were harvested with 0.25% Trypsin/EDTA, washed with PBS twice, and then incubated with monoclonal antibody in 100 μL PBS at 4 °C for 30 minutes. Then the cells were washed in PBS three times and resuspended in a volume of 300 μL PBS. Fluorescence of the cells was detected by flow cytometer (FACS Caliber, BD). Cell viability detection MSC viability was examined by flow cytometry analysis using LIVE/DEAD® Fixable Dead Cell Stain Kit, which was purchased from Life Technologies (Thermo Fisher Xue et al. Stem Cell Research & Therapy (2017) 8:148 Scientific, Waltham, MA, USA). Briefly, 1 × 105 MSCs were harvested with 0.25% Trypsin/EDTA, washed with PBS twice, and then incubated with 1 μL of DMSOdiluted stain (Cat. No. L23101) in 1 mL protein-free buffer at 4 °C for 30 minutes. Then the cells were washed in PBS three times and resuspended in a volume of 300 μL PBS. Fluorescence of the cells was detected by flow cytometer (FACS Caliber, BD). Cell proliferation assay 1 × 104 MSCs were seeded in a six-well plate and then the cells were continuously cultured for a further 14 days. Each of the three wells of MSCs were harvested with 0.25% Trypsin/EDTA, washed with PBS twice, and then counted by cell counting counter (Countstar, Shanghai Ruiyu BioTech Co., Ltd., Shanghai, China) every 2 days. The proliferation curve was made by using GraphPad Prism (v 6.0, GraphPad Software, Inc., San Diego, CA, USA). Osteogenic differentiation of MSCs MSCs were seeded in 24-well culture plates in a density of 2 × 104 cells for each well with osteogenic culture medium to induce osteogenic differentiation for 21 days. The medium was refreshed every 2 days. Inducing osteogenic differentiation medium contained 10% FBS, 100 μg/mL streptomycin, 100 U/ml penicillin, 10 mM βglycerolphosphate, 0.1 μM dexamethasone, and 0.2 mM ascorbate. At the 21th day of differentiation, the MSCs were fixed with 4% formaldehyde and Alizarin Red S staining was employed to examine the osteogenic differentiation of MSCs. Real-time PCR Trizol reagent (Invitrogen, Carlsbad, CA, USA) was used to isolate total RNA according to the introductions from the manufacturer. RevertAid RT-PCR system (Fermentas, Waltham, MA, USA) was employed to reverse-transcribe the total RNA into cDNA. Then the cDNA was mixed with primers and Maxima SYBR Green qPCR Master Mix in the Real-time PCR Stratagene Mx3000P System (Applied Biosystems, Foster City, CA, USA). The levels of mRNA in each group were compared after normalization by GAPDH. The primer sequences are listed in Table 1. Western blotting analysis Total proteins extraction from MSCs and Western blot analysis were performed as described [15]. Briefly, MSCs were lysed in RIPA buffer with 1 mM PMSF at 0 °C. The cell lysates were centrifuged in 4 °C at 12,000 rpm for 10 minutes. Then the protein supernatants were transferred into new tubes. The concentration of protein in each sample was assessed by BCA Protein Assay Kit. The protein was mixed with laemmli sample buffer, heated at for 10 minutes at 65 °C. The samples were loaded (20 μg Page 3 of 9 Table 1 Sequence of the oligonucleotides for real-time PCR Gene Sequence (5′ → 3′) β-actin F GTGGGGCGCCCCAGGCACCA R CTTCCTTAATGTCACGCACGATTTC F ACGACAACCGCACCATGGT R CTGTAATCTGACTCTGTCCT F TGGAGCTTCAGAAGCTCAACACCA R ATCTCGTTGTCTGAGTACCAGTCC F CCTGAGCCAGCAGATTGA R TCCGCTCTTCCAGTCAG F GAGGTCCTGAGCGAGTTCGA R ACCTGAGTGCCTGCGATACA Runx-2 ALP Collagen I BMP-2 per sample) and separated by sodium dodecyl sulfatepolyacrylamide gel (7.5%) electrophoresis in denaturing conditions and electroblotted on nitrocellulose membranes. The membranes were blocked by Tris-buffered saline containing Tween 20 (TBST) with 5% nonfat milk at room temperature for 2 hours. Primary antibodies of βcatenin (610154, BD Biosciences), Smad3 (ab40854, Abcam, Cambridge, MA, USA), and p-Smad3 (ab52903, Abcam) were incubated with the membranes at 4 °C overnight. Then the membranes were incubated with horseradish peroxidase (HRP)-conjugated secondary antibodies (Dako, Glostrup, Denmark) and the results were observed by enhanced chemiluminescence. GAPDH was employed as internal control to normalize the loading protein. Scanning electron microscope The scanning electron microscope (SEM; JEOL 5300, JEOL USA Inc., Peabody, MA, USA) was used to observe the MSCs on PLC. Specimens with MSCs were rinsed with PBS buffer and fixed with 2.5% (v/v) glutaraldehyde in a 0.1 mol/L sodium cacodylate buffer for 2 hours and then were post fixed in 1% (w/v) OsO4 for 1 hour. The specimens were subjected to graded alcohol dehydration, washed with hexamethyldisilazane, coated with gold, and observed by SEM. Statistical analysis All data, expressed as mean ± standard deviation (SD), were from at least three separate experiments. Statistical analysis was performed by t test (two-tailed) by SPSS 18.0 software (SPSS, Inc., Chicago, IL, USA). P



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