JOURNAL OF TISSUE ENGINEERING AND REGENERATIVE MEDICINE RESEARCH J Tissue Eng Regen Med (2015) Published online in Wiley Online Library (wileyonlinelibrary.com) DOI: 10.1002/term.1999
ARTICLE
Potential of VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation of human mesenchymal stem cells Dan Kai1,4, Molamma P. Prabhakaran2*, Guorui Jin3,4, Lingling Tian2 and Seeram Ramakrishna2,3 1
NUS Graduate School for Integrative Sciences and Engineering, National University of Singapore Centre for Nanofibers and Nanotechnology, Nanoscience and Nanotechnology Initiative, National University of Singapore 3 Department of Mechanical Engineering, National University of Singapore 4 Institute of Materials Research and Engineering (IMRE), Singapore 2
Abstract Heart disease, especially myocardial infarction (MI), has become the leading cause of death all over the world, especially since the myocardium lacks the ability to regenerate after infarction. The capability of mesenchymal stem cells (MSCs) to differentiate into the cardiac lineage holds great potential in regenerative medicine for MI treatment. In this study, we investigated the potential of human MSCs (hMSCs) to differentiate into cardiomyogenic cell lineages, using 5-azacytidine (5-aza) on electrospun poly(ε-caprolactone)–gelatin (PCL–gelatin) nanofibrous scaffolds. Immunofluorescence staining analysis showed that after 15 days of in vitro culture the hMSCs differentiated to cardiomyogenic cells on PCL–gelatin (PG) nanofibers and expressed a higher level of cardiacspecific proteins, such as α-actinin and troponin-T, compared to the MSC-differentiated CMs on tissue culture plates (control). To further induce the cardiac differentiation, vascular endothelial growth factor (VEGF) was incorporated into the nanofibers by blending or co-axial electrospinning, and in vitro release study showed that the growth factor could cause sustained release of VEGF from the nanofibers for a period of up to 21 days. The incorporation of VEGF within the nanofibers improved the proliferation of MSCs and, more importantly, enhanced the expression of cardiacspecific proteins on PG–VEGF nanofibers. Our study demonstrated that the electrospun PG nanofibers encapsulated with VEGF have the ability to promote cardiac differentiation of hMSCs, and might be promising scaffolds for myocardial regeneration. Copyright © 2015 John Wiley & Sons, Ltd. Received 18 July 2014; Revised 13 November 2014; Accepted 12 December 2014
Keywords
stem cells; gelatin; electrospinning; release rate; elastic modulus
1. Introduction Despite the advancement of the medical industry and clinical practice due to the burst of nanotechnology and biomedical research in the past 20 years, heart failure remains a leading cause of morbidity and mortality among patients suffering from cardiovascular disease (Singla, 2010). Myocardial infarction (MI) progresses from the acute death of cardiomyocytes (CMs), leading to scar *Correspondence to: Molamma P. Prabhakaran, Centre for Nanofibers and Nanotechnology, E3-05-14, Nanoscience and Nanotechnology Initiative, Faculty of Engineering, National University of Singapore, 2 Engineering Drive 3, Singapore 117576. E-mail: [email protected] Copyright © 2015 John Wiley & Sons, Ltd.
tissue formation accompanied by the rearrangement of cardiac structure, including infarct wall thinning, cardiac dilatation and remodelling (Jugdutt, 2003; Wei et al., 2006). Although several cardiovascular therapies, such as percutaneous treatments, implantable defibrillators, cardiac resynchronization therapy and left ventricular assist devices, have been applied to treat MI, such procedures are associated with several disadvantages, and difficulties still persist to completely restore or regenerate the injured myocardial tissue (Kuraitis et al., 2011). Over recent years, attention has turned to the potential of stem cells to repair the damaged myocardium, and multiple stem cell types have been considered for MI therapies, including mesenchymal stem cells (MSCs), haematopoietic stem cells, embryonic stem cells and
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induced pluripotent stem cells (Vunjak-Novakovic et al., 2010). Attempting to amplify and/or recreate the natural processes of tissue formation, the therapeutic stem cells are expected to augment the creation of myocardium in vivo, and even replace damaged or dead cells. MSCs isolated from bone marrow have been widely used for the regeneration of a variety of mesenchymal tissues, including bone, skin, muscle and cartilage, mainly because these cells possess specific advantages, such as multipotency and established isolation/culture technology; thus, this is the most preferable stem cell model for MI regeneration (Pittenger and Martin, 2004). Recently, a number of studies have suggested that MSCs could be made to differentiate into CMs by treatment with 5-azacytidine (5-aza), bone morphogenetic protein-2 or fibroblast growth factor-4, as well as by co-culture with CMs (Kai et al., 2010). It is reported that mere cell disposition to the recipient myocardium via myocardial injection has several drawbacks, including a high cell death ratio and insufficient prevention of progressive left ventricular dilatation (Muller-Ehmsen et al., 2002; Reinecke and Murry, 2002). To overcome these limitations, we here hypothesize a novel approach of seeding the stem cells on ‘smart’ biodegradable polymeric nanofibrous scaffolds and applying them as cardiac patches for regeneration of the infarcted heart. Electrospinning has gained increasing attention as a method for fabricating nanofibrous scaffolds with micro/nanoscale features similar to the architecture of the native extracellular matrix (ECM). The ability to mimic the structural organization of the ECM is a ’golden’ criterion to design a cell-responsive scaffold platform upon which additional functionalities can be incorporated (Lim and Mao, 2009). Electrospun polymeric nanofibers offer a higher surface area for protein and cell attachment, a highly porous scaffold for oxygen and nutrition transportation, with flexibility of loading growth factors or drugs. Electrospun nanofibers have shown several advantages, such as a drug delivery system for the targeted delivery of proteins or drugs (Ji et al., 2011). For example, their release profile can be finely tailored by tuning the composition, morphology and porosity of the fibrous membrane, while their small fibre diameter might help to provide a short diffusion passage length, and their high surface area could be beneficial for mass transfer and efficient drug release. In this study, PCL–gelatin (PG) nanofibers were fabricated as a substrate for cardiomyogenic differentiation of hMSCs. A biodegradable synthetic polymer PCL with good mechanical properties and the natural polymer, gelatin, with cell interaction cues were blended by electrospinning, and PG nanofibers were demonstrated to promote the attachment of CMs, together with the expression of cardiac-specific proteins (Kai et al., 2011). Vascular endothelial growth factor (VEGF) is an important signal protein that stimulates vasculogenesis and angiogenesis. However, VEGF has a short half-life, and hence it cannot meet long-term requirements in TE applications. Here we incorporated VEGF into nanofibrous scaffolds to induce the Copyright © 2015 John Wiley & Sons, Ltd.
cardiac differentiation of MSCs seeded on the scaffolds. The nanofibers could assist in the controlled release of VEGF, and might effectively retain its bioactivity and optimize the dosage effect. We hypothesize that the synergetic effect of the topographical cues of nanofibers and the biological cues of VEGF could promote the cardiomyogenic differentiation of hMSCs for application of PG–VEGF as a substrate for myocardial regeneration.
2. Materials and methods 2.1. Materials PCL (molecular weight = 80 000), gelatin type A, 1,1,1,3,3,3hexafluoro-2-propanol (HFP), Dulbecco’s modified Eagle’s medium (DMEM), 5-aza, monoclonal anti-α-actinin, anti-troponin-T and anti-myosin heavy chain antibody produced in mouse, were all obtained from Sigma-Aldrich (St. Louis, MO, USA). Hexamethyldisilazane and polyvinyl alcohol mounting medium were purchased from Fluka. Fetal bovine serum (FBS), phosphate-buffered saline (PBS) and trypsin– EDTA were purchased from Gibco (Singapore). CellTiter 96 AQueous One solution reagent was purchased from Promega (Singapore). Methanol was purchased from Sinopharm Chemical Reagent Co. (Singapore).
2.2. Fabrication of electrospun nanofibers PG–VEGF nanofibers were prepared by blending and core– shell electrospinning processes. To fabricate PG–VEGF blending nanofibers (PGV/B), PCL, gelatin and VEGF was dissolved in HFP to obtain the electrospinning solution, with a concentration of 8 wt% at a weight ratio of 1 g:1 g:10 μg. After being stirred for 24 h, the solution was electrospun from a 5 ml syringe with a 27G blunted stainless steel needle at a flow rate of 1.0 ml/h. A high voltage of 15 kV (Gamma High Voltage Research, USA) was applied, when the polymer solution was drawn into fibres and collected on an aluminium foil-wrapped collector kept at a distance of 10 cm from the needle tip. PG nanofibers were also electrospun, using the same parameters as the control. A two-fluid co-axial spinneret was used for electrospinning core–shell nanofibers, where the inner needle had a diameter of 0.4 mm and the outer needle had a diameter of 1.8 mm. 8 wt% PCL–gelatin (1:1) was dissolved in HFP to form the shell solution, while VEGF was dissolved in PBS in the concentration of 1 μg/ml to form the core solution. During the core–shell electrospinning, the flow rate was set at 0.4 ml/h for the core and 1.0 ml/h for the shell solution. DC power supply was used to provide a high voltage of 18 kV, and a collector plate was placed at a distance of 15 cm from the tip of the spinneret to collect the PG–VEGF core–shell nanofibers (PGV/CS). The fibres produced were subsequently vacuum-dried so as to remove any residual solvents. J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
2.3. Characterization of electrospun nanofibers The morphology of electrospun nanofibrous scaffolds were studied by scanning electron microscopy (JSM5600, Jeol, Japan) at an accelerating voltage of 15 kV. Before observation, the scaffolds were coated with gold, using a sputtercoater (Jeol JFC-1200 Fine Coater, Japan). The diameter of the fibres was measured from the SEM micrographs, using image analysis software (ImageJ, National Institutes of Health, USA). Verification of core–shell structure was carried out using transmission electronic microscopy (TEM; JEM-3010, Jeol, Tokyo, Japan) at 300 kV, and the samples for TEM observations were prepared by collecting the nanofibers on carbon-coated Cu grids. In addition, FITC was added into the core solution during the co-axial electrospinning, and the FITC included core–shell nanofibers were viewed under confocal laser scanning microscopy (CLSM; Olympus Fluoview FV1000, Olympus Optical, Tokyo, Japan) at excitation wavelengths of 488 nm, to observe the fluorescent core of the electrospun coaxial nanofibers. The wettability of the electrospun nanofibers was evaluated by drop water contact angle measurement, using a VCA Optima Surface Analysis System (AST Products, Billerica, MA, USA). The droplet size was 0.5 μl and at least five samples were used for each test. Tensile properties of electrospun nanofibrous scaffolds were determined using a tabletop tensile tester (Instron 3345, Canton, MA, USA) at a load cell capacity of 10 N. Test specimens of dimensions 10 mm breadth × 30 mm length were tested at a crosshead speed of 5 mm/min, at a gauge length of 20 mm, under ambient conditions. A minimum of six specimens of individual scaffolds were tested and the results obtained were plotted for the stress–strain curves of the scaffolds.
2.4. In vitro VEGF release study In the VEGF release behaviour study, 20 mg PG–VEGF nanofibrous mat was soaked in 10 ml PBS at 37°C with shaking, in an incubator shaker, at 100 rpm. At each time point, 1 ml supernatant was retrieved from the vial and an equal volume of fresh PBS was replaced. The concentration of VEGF in the supernatant was determined by Human VEGF ELISA Kit (Immunocell, Singapore).
2.5. Human MSCs culture and differentiation The electrospun scaffolds were sterilized under UV light for 2 h, washed three times with PBS for 20 min each, and incubated in DMEM for 2 h before cell seeding. hMSCs (Lonza, Walkersville, MD, USA) at an early passage (< P4) were cultured in DMEM supplemented with 10% FBS and 1% penicillin–streptomycin. After reaching 70–80% confluence, the cells were detached using trypsin–EDTA and viable cells were counted by Trypan blue assay. The cells were further seeded onto the Copyright © 2015 John Wiley & Sons, Ltd.
scaffolds at a cell density of 1 × 104 cells/well. The medium was changed every 2 days. Similarly, neonatal rat CMs (Lonza) were cultured on PG nanofibers under the same conditions as the control. To induce the cardiac differentiation of hMSCs, 24 h after seeding cells were exposed to α-MEM containing 10 μM 5-aza for 24 h and then washed three times with PBS. hMSCs continued to be cultured on the scaffolds for 2 weeks until the differentiation was evaluated by immunostaining.
2.6. Cell proliferation assay Cell proliferation, with and without 5-aza treatment, was analysed after 5, 10 and 15 days, using the colorimetric 3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)2-(4-sulphophenyl)-2H-tetrazolium (MTS) assay. At each time point, cells were washed with PBS and incubated with 20% MTS reagent containing serum-free medium. After 4 h of incubation at 37°C in 5% CO2 incubator, aliquots were pipetted into a 96-well plate and the absorbance of the content of each well was measured at 492 nm, using a spectrophotometric plate reader (Fluostar Optima, BMG Lab Technologies, Germany).
2.6. Cell morphology A morphological study of hMSCs cultured on different scaffolds, with and without 5-aza treatment, was carried out after 15 days of cell culture by SEM. The cell–scaffold constructs were rinsed twice with PBS and fixed in 3% glutaraldehyde for 3 h. The scaffolds were further rinsed in de-ionized water and dehydrated with increasing concentrations of ethanol (50%, 70%, 90% and 100%) for 20 min each. Finally, the cell–scaffold constructs were treated with hexamethyldisilazane and air-dried in a fume hood overnight.
2.7. Immunocytochemistry Double immunofluorescent staining was performed to confirm the cardiomyogenic differentiation of hMSCs. After 15 days of culture, cells were fixed using 2.5% paraformaldehyde and permeated using 0.1% Triton X-100. The cells were further blocked with 3% bovine serum albumin (BSA) and incubated individually with the mouse-derived primary antibodies α-actinin (α-ACT) and troponin-T (TP-T) for 2 h at dilutions of 1:100 each. After washing with PBS, the cells were incubated with the secondary antibody AlexaFluor 488 (green) at a dilution of 1:400 for 1.5 h at room temperature. Then the samples were washed three times with PBS and treated with the hMSC-specific marker CD44 for 2 h at room temperature. This was followed by the addition of the secondary antibody, AlexaFluor 546 (red) and counterstained with DAPI (1:500) at room temperature. The samples were then mounted on glass slides, using mounting medium, and J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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imaged using CLSM. The signal intensity of the cardiac markers was measured using ImageJ. Rat CMs cultured on PG nanofibers were also immunostained under the same conditions as the control.
2.8. Statistical analysis All the data presented are expressed as mean ± standard deviation (SD) of the mean. Single-factor analysis of variance was carried out to compare the means of different datasets; p ≤ 0.05 was considered statistically significant.
3. Results 3.1. Characterization of electrospun nanofibers Electrospinning parameters were optimized to obtain beadless and uniform PG and PG–VEGF nanofibers, while PG–VEGF was prepared using two different methods, (a) blend spinning and (b) core–shell spinning. Figure 1A–C shows the morphology of PG, PGV/B and PGV/CS nanofibers, with average fibre diameters of 265 ± 72, 199 ± 37 and 297 ± 79 nm, respectively. To verify the core–shell structure of PGV/CS fibres, TEM and CLSM images were
taken. As shown in Figure 1D, the core–shell interface with core-phase VEGF encased by PG shells were exhibited by TEM analysis. Figure 1E displays the CLSM image of PGV/CS with a fluorescent stain of FITC, where the CLSM images showed a green core with a colourless outer shell. The surface hydrophilicity of electrospun scaffolds were determined by water contact angle test. The results (Table 1) show that PG, PGV/B and PGV/CS scaffolds had hydrophilic surfaces, with average water contact angles of 35.2 ± 2.7°, 21.7 ± 1.3° and 31.4 ± 2.5°, respectively. The incorporation of VEGF into the nanofibers slightly increased the hydrophilicity of the VEGF blended nanofibers, but the surfaces of the core–shell fibres had no significant differences in contact angle compared to PG fibres. The mechanical properties of electrospun nanofibers were evaluated from the stress–strain curves obtained from the Instron universal mechanical test machine, and the Young’s modulus, tensile strength and elongation at break of the scaffolds are shown in Figure 1F and Table. 1. The average Young’s modulus of the PG scaffold was 7.69 ± 0.80 MPa, with a tensile strength of 4.07 ± 1.19 MPa and elongation at break of 52.0 ± 6.5%. The addition of VEGF weakened the mechanical properties compared to electrospun PG nanofibers. PGV/B showed the lowest mechanical properties, with Young’s modulus of 5.54 ± 0.64 MPa, tensile
Figure 1. SEM micrographs of electrospun nanofibers: (A) PG; (B) PGV/B; (C) PGV/CS. (D) TEM micrograph. (E) CLSM image of PGV/ CS. (F) Stress–strain curve of electrospun PG, PGV/B and PGV/CS nanofibers
Table 1. Fibre diameter, water contact angle and mechanical properties of electrospun PG, PGV/B and PGV/CS nanofibers Samples
Fibre diameter (nm)
Water contact angle (º)
Young’s modulus (MPa)
Tensile strength (MPa)
Elongation at break (%)
PG PGV/B PGV/CS
265 ± 72 199 ± 37 297 ± 79
35.2 ± 2.7 21.7 ± 1.3 31.4 ± 2.5
7.69 ± 0.80 5.54 ± 0.64 8.48 ± 0.53
4.07 ± 1.19 1.91 ± 0.43 3.82 ± 0.91
52.0 ± 6.5 32.6 ± 7.2 35.8 ± 4.8
Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
strength of 1.91 ± 0.43 MPa and elongation at break of 32.6 ± 7.2%, while PGV/CS exhibited a similar stiffness (8.48 ± 0.53 MPa) and tensile strength (3.82 ± 0.91 MPa) but lower elongation at break (35.8 ± 4.8%) compared to PG nanofibers.
the core–shell structured PGV/CS nanofibers were comparatively more desirable for preventing the initial burst release and for the sustained release of VEGF.
3.3. Cell proliferation and morphology 3.2. In vitro release study Release behaviours of VEGF from both kinds of nanofibers (PGV/B and PGV/CS) were investigated during this study. As shown in Figure 2A, PGV/B released 42.8 ± 7.2% VEGF in the first 24 h while PGV/CS released 22.8 ± 1.3%, indicating that the initial burst release of the VEGF was controlled effectively from the core–shell structure. For the following 2 weeks, at every time point of study, PGV/B showed significantly higher release of VEGF than PGV/CS. After 504 h (21 days) of the test, the total release amounts of PGV/B and PGV/CS were 81.6 ± 9.6% and 76.9 ± 5.8%, respectively. The results indicated that
In order to understand the effect of 5-aza treatment, as well as to evaluate the proliferation of hMSCs, we performed the MTS assay for cells on different scaffolds, either with or without 5-aza treatment, for a period of up to 15 days. As shown in Figure 2B, cell proliferations on both PGV/B and PGV/CS scaffolds, with or without 5aza treatment, was significantly higher than on PG scaffolds (p ≤ 0.05). At the same time, there was no significant difference in cell proliferations between PGV/B and PGV/CS scaffolds. The results indicated that the VEGFcontaining scaffolds benefited towards the proliferation of cells, due to the presence of the growth factor. In addition, 5-aza treatment significantly reduced cell growth on
Figure 2. (A) Release profiles of VEGF from PGV/B and PGV/CS nanofibers. (B) Proliferation of hMSCs, with or without 5-aza treatment, on PG, PGV/B and PGV/CS nanofibers, as determined by MTS assay: significantly different from cell proliferation on corre# sponding sample with 5-aza treatment at *p ≤ 0.05; PG scaffolds with 5-aza treatment at p ≤ 0.05; PG scaffolds without 5-aza $ treatment at p ≤ 0.05. (C) Morphology of hMSCs on (C1, C4) TCP; (C2, C5) PG; (C3, C6) PGV/B; and (C4, C8) PGV/CS without 5-aza treatment (C1–C4) and with 5-aza treatment (C5, C6). (C9) CMs on nanofibers as positive control; scale bar = 100 μm Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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scaffolds compared to cell growth in medium devoid of 5-aza reagent, suggesting that the incorporation of 5-aza in the culture medium induced cytotoxicity to the hMSCs and reduced the cell proliferation rate, and this was especially true in the early days of cell culture (day 5). Figure 2C shows the morphology of hMSCs on TCP and nanofibrous scaffolds, with or without 5-aza treatment. hMSCs cultured on the scaffolds without the presence of 5-aza showed elongated and extended filopodia and the cells remained undifferentiated, with a fibroblastic phenotype. However, after 5-aza treatment, a few hMSCs on the nanofibers exhibited an entirely different morphology, whereby they appeared triploid in shape and were very similar to the morphology of native CMs (Figure 2C9).
were found to express higher α-ACT than on PGV/CS, and the reason might be the faster release of VEGF from ‘blend’ nanofibers. Figure 4 exhibits the expression of TP-T of hMSCs on TCP and nanofibers, where hMSCs on nanofibers showed significantly higher expression of TPT than cells on TCP, but there was no significant difference in protein expression between PG and PG–VEGF scaffolds (Figure 4N). We performed immunostaining of α-ACT and TP-T on neonatal rat CMs to gain an understanding of the striation patterns. The results also showed that our hMSCs were positive to the cardiac proteins (see supporting information, Figure S1); however, the distribution of the protein was somewhat different from that expressed by the mature CMs.
3.4. Expression of cardiac-specific proteins by immunostaining
4. Discussion
To demonstrate the cardiomyogenic differentiations of hMSCs, dual immunofluorescence staining of cardiacspecific proteins, α-ACT and TP-T, as well as the MSCspecific marker CD44, was carried out for cells grown on the scaffolds after 15 days of culture with 5-aza treatment. Figure 3 shows that hMSCs on nanofibrous scaffolds expressed significantly higher levels of α-ACT than cells on TCP. Higher expression of α-ACT was found for cells cultured on PGV/B and PGV/CS compared to those on TCP, indicating that VEGF could promote the cardiac differentiation of hMSCs. Furthermore, cells on PGV/B
Stem cell therapy has been widely accepted as an effective approach for treating heart disease, and MSCs with capacity for self-renewal and cardiomyogenic differentiation offer an extremely promising therapeutic strategy for myocardial regeneration. Instead of directly injecting stem cells into the heart, applying a patch of stem cellpopulated bio-graft might be more versatile in controlling the properties of the tissue graft. Furthermore, a wellengineered scaffold might serve as a ’stem cell niche’ and provide a suitable microenvironment for the cells to proliferate and differentiate into specialized mature cells,
Figure 3. Dual immunocytochemical analysis for the expression of (A–D) α-ACT and (E–H) CD44. (I–L) Merged images showing the dual expression on (A, E, I) TCP, (B, F, J) PG, (C, G, K) PGV/B and (D, H, L) PGV/CS nanofibers; scale bar = 100 μm. (M) Expression of α-ACT in CMs on nanofibers as positive control; scale bar = 40 μm. (N) Signal intensity of α-ACT expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
Figure 4. Dual immunocytochemical analysis for the expression of (A–D) TP-T and (E–H) CD44. (I–L) Merged images showing the dual expression on (A, E, I) TCP, (B, F, J) PG, (C, G, K) PGV/B and (D, H, L) PGV/CS nanofibers; scale bar =50 μm. (M) Expression of TP-T in CMs on nanofibers as positive control; scale bar =40 μm. (N) Signal intensity of TP-T expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers
such as the CMs. In this study, we fabricated electrospun PG nanofibers containing VEGF as an artificial niche, and evaluated their potential towards the cardiomyogenic differentiation of hMSCs. In vitro cardiomyogenic differentiation of MSCs by treatment with 5-aza was first reported by Makino et al. (1999). After treatment with 3 μM 5-aza in culture medium, it was found that almost 30% of fibroblast-shaped mouse MSCs changed to a ball-like phenotype or formed the characteristic rod-shaped CMs, and they also expressed the cardiac-specific markers, such as the sarcomeres (Makino et al., 1999). Further, 5-aza treatment became a classical approach for inducing MSCs to CMs. However, most of these differentiation studies were carried out in flat culture flasks or TCP. Studies of such a kind on TCP might not be suitable for any implantations, and hence we carried out the differentiation of hMSCs to cardiomyocytes on electrospun substrates. We chose to blend a synthetic polymer, PCL, with good elasticity and biostability, with a natural polymer, gelatin, which has hydrophilic properties and excellent biocompatibility. The composite scaffold might benefit from the properties of both types of polymer and might counteract its singular defects. In our previous study, we have also demonstrated that PG nanofibers possess suitable hydrophilic and mechanical degradation properties and favour the growth of CMs (Kai et al., 2011). Here we demonstrated the capability of differentiation of hMSCs to the cardiac lineage on these nanofibrous scaffolds in the presence of small amounts of 5-aza reagent. α-ACT is an important Copyright © 2015 John Wiley & Sons, Ltd.
constituent of the contractile apparatus in the native CMs, and TP-T is a related protein that regulates the force and velocity of myocardial contraction (Di Domenico et al., 2012; Gallina et al., 2012). Immunostaining results show that MSCs on PG nanofibers expressed significantly higher levels of cardiac-specific proteins than those on TCP. Our quantification results showed that the expression of αACT by the MSC-differentiated CMs on PG was 73.2% higher than its expression on TCP, while TP-T was expressed three times higher than on TCP, indicating that the cardiac substrate with nanofibrous topography could provide additional chemical and mechanical cues to promote the cardiac differentiation of MSCs. Similarly, Yang et al. (2009) found that the differentiated rat MSCs cultured on silk fibroin–chitosan–hyaluronic acid cardiac patches expressed higher cardiac proteins (TP-T, cardiotin and connexin 43) and improved the expressions of cardiac genes (Gata4, Nkx2.5, Tnnt2 and Actc1) compared to TCP and pure silk fibroin scaffolds. To improve cardiac differentiation, VEGF was incorporated in the nanofibers by two different approaches, blend spinning and core–shell spinning. VEGF is as an important angiogenic growth factor, has been used in improving cardiac function after MI in animal models and has been shown to stimulate angiogenesis with the formation of vascular structures in the infarcted area (Chiu and Radisic, 2010; Guo et al., 2012). During our study, PGV/B had a reduced fibre diameter of approximately 199 ± 37 nm, a reduced water contact angle as well as reduced mechanical properties compared to PG (Table 1). On the other hand, J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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encapsulating VEGF within the core of the PG nanofibers did not significantly change the mechanical properties of the nanofibers, except for a slight decrease in its elongation at break compared to PG (Table 1). In TE the speed of cell growth is important, as it is related to the speed of regeneration of new tissue, and a rapid proliferation rate would obviously result in an improved tissue development. Figure 2B shows that the VEGF-containing nanofibers significantly promoted the proliferation of hMSCs compared to pure PG nanofibers. Similarly, Huang et al. (2010) demonstrated that the profiled addition of VEGF in the culture medium increased the proliferation of mouse MSCs by 1.9-fold. The coinjection of MSCs with VEGF into MI hearts was also described to increase the grafted cell viability and improved cardiac function (Pons et al., 2008). VEGF could not only promote cell proliferation but, more importantly, regulate cellular differentiation. As shown in Figure 3, cells on both PGV/B and PGV/CS expressed higher levels of αACT than those on TCP and PG nanofibers, and the increased protein expression of cells on VEGF-containing nanofibers further confirmed that critical function of VEGF in the developmental stages of cardiomyogenesis (see supporting information, Figure S2). Chen et al. (2006) previously reported that adding VEGF to the culture medium significantly enhanced α-MHC, Troponin-I and Nkx2.5 expression in differentiated embryonic stem cells. Later, studies by Song et al. (2007) also showed that VEGF was a critical factor for the spontaneous differentiation of adipose tissue-derived stem cells, and the growth factor was responsible for the expression of cardiac markers via a paracrine mechanism. Here we demonstrated the positive function of VEGF towards the
cardiomyogenic differentiation of hMSCs, in a different perspective, by performing the experiments on VEGF incorporated PG nanofibers. Our study shows that the PG– VEGF substrates hold great potential application as a bioengineered cardiac patch for myocardial regeneration.
5. Conclusion PG, PGV/B and PGV/CS nanofibrous scaffolds were fabricated to evaluate their effects on the growth and cardiomyogenic differentiation of hMSCs in vitro. The incorporation of VEGF into the nanofibers improved the proliferation of MSCs and, more importantly, improved the expression of cardiac-specific markers. After 15 days, hMSCs cultured on PGV/B and PGV/CS scaffolds showed higher expression of α-ACT and TP-T proteins compared to cells grown on pure PG nanofibers, indicating the potential of VEGF-encapsulated PG nanofibers as a suitable scaffold for myocardial regeneration.
Conflict of interest The authors have declared that there is no conflict of interest.
Acknowledgements This research was supported by NRF–Technion (Grant No. WBS R-398-001-065-592) and the Nanoscience and Nanotechnology Initiative in NUS, Singapore.
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Supporting information on the internet The following supporting information may be found in the online version of this article: Figure S1. Dual immunocytochemical analysis for the expression of myosin heavy chain (MHC) and CD44; merged images showing dual expression on TCP, PG, PGV/B and PGV/CS nanofibers; expression of MHC in CMs on nanofibers as positive control; and signal intensity of MHC expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers Figure S2. Normalized gene expression of Tnni3 in differentiated hMSCs on TCP and PG, PGV/B and PGV/CS nanofibers with 5-aza treatment
Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
ARTICLE
Potential of VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation of human mesenchymal stem cells Dan Kai1,4, Molamma P. Prabhakaran2*, Guorui Jin3,4, Lingling Tian2 and Seeram Ramakrishna2,3 1
NUS Graduate School for Integrative Sciences and Engineering, National University of Singapore Centre for Nanofibers and Nanotechnology, Nanoscience and Nanotechnology Initiative, National University of Singapore 3 Department of Mechanical Engineering, National University of Singapore 4 Institute of Materials Research and Engineering (IMRE), Singapore 2
Abstract Heart disease, especially myocardial infarction (MI), has become the leading cause of death all over the world, especially since the myocardium lacks the ability to regenerate after infarction. The capability of mesenchymal stem cells (MSCs) to differentiate into the cardiac lineage holds great potential in regenerative medicine for MI treatment. In this study, we investigated the potential of human MSCs (hMSCs) to differentiate into cardiomyogenic cell lineages, using 5-azacytidine (5-aza) on electrospun poly(ε-caprolactone)–gelatin (PCL–gelatin) nanofibrous scaffolds. Immunofluorescence staining analysis showed that after 15 days of in vitro culture the hMSCs differentiated to cardiomyogenic cells on PCL–gelatin (PG) nanofibers and expressed a higher level of cardiacspecific proteins, such as α-actinin and troponin-T, compared to the MSC-differentiated CMs on tissue culture plates (control). To further induce the cardiac differentiation, vascular endothelial growth factor (VEGF) was incorporated into the nanofibers by blending or co-axial electrospinning, and in vitro release study showed that the growth factor could cause sustained release of VEGF from the nanofibers for a period of up to 21 days. The incorporation of VEGF within the nanofibers improved the proliferation of MSCs and, more importantly, enhanced the expression of cardiacspecific proteins on PG–VEGF nanofibers. Our study demonstrated that the electrospun PG nanofibers encapsulated with VEGF have the ability to promote cardiac differentiation of hMSCs, and might be promising scaffolds for myocardial regeneration. Copyright © 2015 John Wiley & Sons, Ltd. Received 18 July 2014; Revised 13 November 2014; Accepted 12 December 2014
Keywords
stem cells; gelatin; electrospinning; release rate; elastic modulus
1. Introduction Despite the advancement of the medical industry and clinical practice due to the burst of nanotechnology and biomedical research in the past 20 years, heart failure remains a leading cause of morbidity and mortality among patients suffering from cardiovascular disease (Singla, 2010). Myocardial infarction (MI) progresses from the acute death of cardiomyocytes (CMs), leading to scar *Correspondence to: Molamma P. Prabhakaran, Centre for Nanofibers and Nanotechnology, E3-05-14, Nanoscience and Nanotechnology Initiative, Faculty of Engineering, National University of Singapore, 2 Engineering Drive 3, Singapore 117576. E-mail: [email protected] Copyright © 2015 John Wiley & Sons, Ltd.
tissue formation accompanied by the rearrangement of cardiac structure, including infarct wall thinning, cardiac dilatation and remodelling (Jugdutt, 2003; Wei et al., 2006). Although several cardiovascular therapies, such as percutaneous treatments, implantable defibrillators, cardiac resynchronization therapy and left ventricular assist devices, have been applied to treat MI, such procedures are associated with several disadvantages, and difficulties still persist to completely restore or regenerate the injured myocardial tissue (Kuraitis et al., 2011). Over recent years, attention has turned to the potential of stem cells to repair the damaged myocardium, and multiple stem cell types have been considered for MI therapies, including mesenchymal stem cells (MSCs), haematopoietic stem cells, embryonic stem cells and
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induced pluripotent stem cells (Vunjak-Novakovic et al., 2010). Attempting to amplify and/or recreate the natural processes of tissue formation, the therapeutic stem cells are expected to augment the creation of myocardium in vivo, and even replace damaged or dead cells. MSCs isolated from bone marrow have been widely used for the regeneration of a variety of mesenchymal tissues, including bone, skin, muscle and cartilage, mainly because these cells possess specific advantages, such as multipotency and established isolation/culture technology; thus, this is the most preferable stem cell model for MI regeneration (Pittenger and Martin, 2004). Recently, a number of studies have suggested that MSCs could be made to differentiate into CMs by treatment with 5-azacytidine (5-aza), bone morphogenetic protein-2 or fibroblast growth factor-4, as well as by co-culture with CMs (Kai et al., 2010). It is reported that mere cell disposition to the recipient myocardium via myocardial injection has several drawbacks, including a high cell death ratio and insufficient prevention of progressive left ventricular dilatation (Muller-Ehmsen et al., 2002; Reinecke and Murry, 2002). To overcome these limitations, we here hypothesize a novel approach of seeding the stem cells on ‘smart’ biodegradable polymeric nanofibrous scaffolds and applying them as cardiac patches for regeneration of the infarcted heart. Electrospinning has gained increasing attention as a method for fabricating nanofibrous scaffolds with micro/nanoscale features similar to the architecture of the native extracellular matrix (ECM). The ability to mimic the structural organization of the ECM is a ’golden’ criterion to design a cell-responsive scaffold platform upon which additional functionalities can be incorporated (Lim and Mao, 2009). Electrospun polymeric nanofibers offer a higher surface area for protein and cell attachment, a highly porous scaffold for oxygen and nutrition transportation, with flexibility of loading growth factors or drugs. Electrospun nanofibers have shown several advantages, such as a drug delivery system for the targeted delivery of proteins or drugs (Ji et al., 2011). For example, their release profile can be finely tailored by tuning the composition, morphology and porosity of the fibrous membrane, while their small fibre diameter might help to provide a short diffusion passage length, and their high surface area could be beneficial for mass transfer and efficient drug release. In this study, PCL–gelatin (PG) nanofibers were fabricated as a substrate for cardiomyogenic differentiation of hMSCs. A biodegradable synthetic polymer PCL with good mechanical properties and the natural polymer, gelatin, with cell interaction cues were blended by electrospinning, and PG nanofibers were demonstrated to promote the attachment of CMs, together with the expression of cardiac-specific proteins (Kai et al., 2011). Vascular endothelial growth factor (VEGF) is an important signal protein that stimulates vasculogenesis and angiogenesis. However, VEGF has a short half-life, and hence it cannot meet long-term requirements in TE applications. Here we incorporated VEGF into nanofibrous scaffolds to induce the Copyright © 2015 John Wiley & Sons, Ltd.
cardiac differentiation of MSCs seeded on the scaffolds. The nanofibers could assist in the controlled release of VEGF, and might effectively retain its bioactivity and optimize the dosage effect. We hypothesize that the synergetic effect of the topographical cues of nanofibers and the biological cues of VEGF could promote the cardiomyogenic differentiation of hMSCs for application of PG–VEGF as a substrate for myocardial regeneration.
2. Materials and methods 2.1. Materials PCL (molecular weight = 80 000), gelatin type A, 1,1,1,3,3,3hexafluoro-2-propanol (HFP), Dulbecco’s modified Eagle’s medium (DMEM), 5-aza, monoclonal anti-α-actinin, anti-troponin-T and anti-myosin heavy chain antibody produced in mouse, were all obtained from Sigma-Aldrich (St. Louis, MO, USA). Hexamethyldisilazane and polyvinyl alcohol mounting medium were purchased from Fluka. Fetal bovine serum (FBS), phosphate-buffered saline (PBS) and trypsin– EDTA were purchased from Gibco (Singapore). CellTiter 96 AQueous One solution reagent was purchased from Promega (Singapore). Methanol was purchased from Sinopharm Chemical Reagent Co. (Singapore).
2.2. Fabrication of electrospun nanofibers PG–VEGF nanofibers were prepared by blending and core– shell electrospinning processes. To fabricate PG–VEGF blending nanofibers (PGV/B), PCL, gelatin and VEGF was dissolved in HFP to obtain the electrospinning solution, with a concentration of 8 wt% at a weight ratio of 1 g:1 g:10 μg. After being stirred for 24 h, the solution was electrospun from a 5 ml syringe with a 27G blunted stainless steel needle at a flow rate of 1.0 ml/h. A high voltage of 15 kV (Gamma High Voltage Research, USA) was applied, when the polymer solution was drawn into fibres and collected on an aluminium foil-wrapped collector kept at a distance of 10 cm from the needle tip. PG nanofibers were also electrospun, using the same parameters as the control. A two-fluid co-axial spinneret was used for electrospinning core–shell nanofibers, where the inner needle had a diameter of 0.4 mm and the outer needle had a diameter of 1.8 mm. 8 wt% PCL–gelatin (1:1) was dissolved in HFP to form the shell solution, while VEGF was dissolved in PBS in the concentration of 1 μg/ml to form the core solution. During the core–shell electrospinning, the flow rate was set at 0.4 ml/h for the core and 1.0 ml/h for the shell solution. DC power supply was used to provide a high voltage of 18 kV, and a collector plate was placed at a distance of 15 cm from the tip of the spinneret to collect the PG–VEGF core–shell nanofibers (PGV/CS). The fibres produced were subsequently vacuum-dried so as to remove any residual solvents. J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
2.3. Characterization of electrospun nanofibers The morphology of electrospun nanofibrous scaffolds were studied by scanning electron microscopy (JSM5600, Jeol, Japan) at an accelerating voltage of 15 kV. Before observation, the scaffolds were coated with gold, using a sputtercoater (Jeol JFC-1200 Fine Coater, Japan). The diameter of the fibres was measured from the SEM micrographs, using image analysis software (ImageJ, National Institutes of Health, USA). Verification of core–shell structure was carried out using transmission electronic microscopy (TEM; JEM-3010, Jeol, Tokyo, Japan) at 300 kV, and the samples for TEM observations were prepared by collecting the nanofibers on carbon-coated Cu grids. In addition, FITC was added into the core solution during the co-axial electrospinning, and the FITC included core–shell nanofibers were viewed under confocal laser scanning microscopy (CLSM; Olympus Fluoview FV1000, Olympus Optical, Tokyo, Japan) at excitation wavelengths of 488 nm, to observe the fluorescent core of the electrospun coaxial nanofibers. The wettability of the electrospun nanofibers was evaluated by drop water contact angle measurement, using a VCA Optima Surface Analysis System (AST Products, Billerica, MA, USA). The droplet size was 0.5 μl and at least five samples were used for each test. Tensile properties of electrospun nanofibrous scaffolds were determined using a tabletop tensile tester (Instron 3345, Canton, MA, USA) at a load cell capacity of 10 N. Test specimens of dimensions 10 mm breadth × 30 mm length were tested at a crosshead speed of 5 mm/min, at a gauge length of 20 mm, under ambient conditions. A minimum of six specimens of individual scaffolds were tested and the results obtained were plotted for the stress–strain curves of the scaffolds.
2.4. In vitro VEGF release study In the VEGF release behaviour study, 20 mg PG–VEGF nanofibrous mat was soaked in 10 ml PBS at 37°C with shaking, in an incubator shaker, at 100 rpm. At each time point, 1 ml supernatant was retrieved from the vial and an equal volume of fresh PBS was replaced. The concentration of VEGF in the supernatant was determined by Human VEGF ELISA Kit (Immunocell, Singapore).
2.5. Human MSCs culture and differentiation The electrospun scaffolds were sterilized under UV light for 2 h, washed three times with PBS for 20 min each, and incubated in DMEM for 2 h before cell seeding. hMSCs (Lonza, Walkersville, MD, USA) at an early passage (< P4) were cultured in DMEM supplemented with 10% FBS and 1% penicillin–streptomycin. After reaching 70–80% confluence, the cells were detached using trypsin–EDTA and viable cells were counted by Trypan blue assay. The cells were further seeded onto the Copyright © 2015 John Wiley & Sons, Ltd.
scaffolds at a cell density of 1 × 104 cells/well. The medium was changed every 2 days. Similarly, neonatal rat CMs (Lonza) were cultured on PG nanofibers under the same conditions as the control. To induce the cardiac differentiation of hMSCs, 24 h after seeding cells were exposed to α-MEM containing 10 μM 5-aza for 24 h and then washed three times with PBS. hMSCs continued to be cultured on the scaffolds for 2 weeks until the differentiation was evaluated by immunostaining.
2.6. Cell proliferation assay Cell proliferation, with and without 5-aza treatment, was analysed after 5, 10 and 15 days, using the colorimetric 3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)2-(4-sulphophenyl)-2H-tetrazolium (MTS) assay. At each time point, cells were washed with PBS and incubated with 20% MTS reagent containing serum-free medium. After 4 h of incubation at 37°C in 5% CO2 incubator, aliquots were pipetted into a 96-well plate and the absorbance of the content of each well was measured at 492 nm, using a spectrophotometric plate reader (Fluostar Optima, BMG Lab Technologies, Germany).
2.6. Cell morphology A morphological study of hMSCs cultured on different scaffolds, with and without 5-aza treatment, was carried out after 15 days of cell culture by SEM. The cell–scaffold constructs were rinsed twice with PBS and fixed in 3% glutaraldehyde for 3 h. The scaffolds were further rinsed in de-ionized water and dehydrated with increasing concentrations of ethanol (50%, 70%, 90% and 100%) for 20 min each. Finally, the cell–scaffold constructs were treated with hexamethyldisilazane and air-dried in a fume hood overnight.
2.7. Immunocytochemistry Double immunofluorescent staining was performed to confirm the cardiomyogenic differentiation of hMSCs. After 15 days of culture, cells were fixed using 2.5% paraformaldehyde and permeated using 0.1% Triton X-100. The cells were further blocked with 3% bovine serum albumin (BSA) and incubated individually with the mouse-derived primary antibodies α-actinin (α-ACT) and troponin-T (TP-T) for 2 h at dilutions of 1:100 each. After washing with PBS, the cells were incubated with the secondary antibody AlexaFluor 488 (green) at a dilution of 1:400 for 1.5 h at room temperature. Then the samples were washed three times with PBS and treated with the hMSC-specific marker CD44 for 2 h at room temperature. This was followed by the addition of the secondary antibody, AlexaFluor 546 (red) and counterstained with DAPI (1:500) at room temperature. The samples were then mounted on glass slides, using mounting medium, and J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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imaged using CLSM. The signal intensity of the cardiac markers was measured using ImageJ. Rat CMs cultured on PG nanofibers were also immunostained under the same conditions as the control.
2.8. Statistical analysis All the data presented are expressed as mean ± standard deviation (SD) of the mean. Single-factor analysis of variance was carried out to compare the means of different datasets; p ≤ 0.05 was considered statistically significant.
3. Results 3.1. Characterization of electrospun nanofibers Electrospinning parameters were optimized to obtain beadless and uniform PG and PG–VEGF nanofibers, while PG–VEGF was prepared using two different methods, (a) blend spinning and (b) core–shell spinning. Figure 1A–C shows the morphology of PG, PGV/B and PGV/CS nanofibers, with average fibre diameters of 265 ± 72, 199 ± 37 and 297 ± 79 nm, respectively. To verify the core–shell structure of PGV/CS fibres, TEM and CLSM images were
taken. As shown in Figure 1D, the core–shell interface with core-phase VEGF encased by PG shells were exhibited by TEM analysis. Figure 1E displays the CLSM image of PGV/CS with a fluorescent stain of FITC, where the CLSM images showed a green core with a colourless outer shell. The surface hydrophilicity of electrospun scaffolds were determined by water contact angle test. The results (Table 1) show that PG, PGV/B and PGV/CS scaffolds had hydrophilic surfaces, with average water contact angles of 35.2 ± 2.7°, 21.7 ± 1.3° and 31.4 ± 2.5°, respectively. The incorporation of VEGF into the nanofibers slightly increased the hydrophilicity of the VEGF blended nanofibers, but the surfaces of the core–shell fibres had no significant differences in contact angle compared to PG fibres. The mechanical properties of electrospun nanofibers were evaluated from the stress–strain curves obtained from the Instron universal mechanical test machine, and the Young’s modulus, tensile strength and elongation at break of the scaffolds are shown in Figure 1F and Table. 1. The average Young’s modulus of the PG scaffold was 7.69 ± 0.80 MPa, with a tensile strength of 4.07 ± 1.19 MPa and elongation at break of 52.0 ± 6.5%. The addition of VEGF weakened the mechanical properties compared to electrospun PG nanofibers. PGV/B showed the lowest mechanical properties, with Young’s modulus of 5.54 ± 0.64 MPa, tensile
Figure 1. SEM micrographs of electrospun nanofibers: (A) PG; (B) PGV/B; (C) PGV/CS. (D) TEM micrograph. (E) CLSM image of PGV/ CS. (F) Stress–strain curve of electrospun PG, PGV/B and PGV/CS nanofibers
Table 1. Fibre diameter, water contact angle and mechanical properties of electrospun PG, PGV/B and PGV/CS nanofibers Samples
Fibre diameter (nm)
Water contact angle (º)
Young’s modulus (MPa)
Tensile strength (MPa)
Elongation at break (%)
PG PGV/B PGV/CS
265 ± 72 199 ± 37 297 ± 79
35.2 ± 2.7 21.7 ± 1.3 31.4 ± 2.5
7.69 ± 0.80 5.54 ± 0.64 8.48 ± 0.53
4.07 ± 1.19 1.91 ± 0.43 3.82 ± 0.91
52.0 ± 6.5 32.6 ± 7.2 35.8 ± 4.8
Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
strength of 1.91 ± 0.43 MPa and elongation at break of 32.6 ± 7.2%, while PGV/CS exhibited a similar stiffness (8.48 ± 0.53 MPa) and tensile strength (3.82 ± 0.91 MPa) but lower elongation at break (35.8 ± 4.8%) compared to PG nanofibers.
the core–shell structured PGV/CS nanofibers were comparatively more desirable for preventing the initial burst release and for the sustained release of VEGF.
3.3. Cell proliferation and morphology 3.2. In vitro release study Release behaviours of VEGF from both kinds of nanofibers (PGV/B and PGV/CS) were investigated during this study. As shown in Figure 2A, PGV/B released 42.8 ± 7.2% VEGF in the first 24 h while PGV/CS released 22.8 ± 1.3%, indicating that the initial burst release of the VEGF was controlled effectively from the core–shell structure. For the following 2 weeks, at every time point of study, PGV/B showed significantly higher release of VEGF than PGV/CS. After 504 h (21 days) of the test, the total release amounts of PGV/B and PGV/CS were 81.6 ± 9.6% and 76.9 ± 5.8%, respectively. The results indicated that
In order to understand the effect of 5-aza treatment, as well as to evaluate the proliferation of hMSCs, we performed the MTS assay for cells on different scaffolds, either with or without 5-aza treatment, for a period of up to 15 days. As shown in Figure 2B, cell proliferations on both PGV/B and PGV/CS scaffolds, with or without 5aza treatment, was significantly higher than on PG scaffolds (p ≤ 0.05). At the same time, there was no significant difference in cell proliferations between PGV/B and PGV/CS scaffolds. The results indicated that the VEGFcontaining scaffolds benefited towards the proliferation of cells, due to the presence of the growth factor. In addition, 5-aza treatment significantly reduced cell growth on
Figure 2. (A) Release profiles of VEGF from PGV/B and PGV/CS nanofibers. (B) Proliferation of hMSCs, with or without 5-aza treatment, on PG, PGV/B and PGV/CS nanofibers, as determined by MTS assay: significantly different from cell proliferation on corre# sponding sample with 5-aza treatment at *p ≤ 0.05; PG scaffolds with 5-aza treatment at p ≤ 0.05; PG scaffolds without 5-aza $ treatment at p ≤ 0.05. (C) Morphology of hMSCs on (C1, C4) TCP; (C2, C5) PG; (C3, C6) PGV/B; and (C4, C8) PGV/CS without 5-aza treatment (C1–C4) and with 5-aza treatment (C5, C6). (C9) CMs on nanofibers as positive control; scale bar = 100 μm Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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scaffolds compared to cell growth in medium devoid of 5-aza reagent, suggesting that the incorporation of 5-aza in the culture medium induced cytotoxicity to the hMSCs and reduced the cell proliferation rate, and this was especially true in the early days of cell culture (day 5). Figure 2C shows the morphology of hMSCs on TCP and nanofibrous scaffolds, with or without 5-aza treatment. hMSCs cultured on the scaffolds without the presence of 5-aza showed elongated and extended filopodia and the cells remained undifferentiated, with a fibroblastic phenotype. However, after 5-aza treatment, a few hMSCs on the nanofibers exhibited an entirely different morphology, whereby they appeared triploid in shape and were very similar to the morphology of native CMs (Figure 2C9).
were found to express higher α-ACT than on PGV/CS, and the reason might be the faster release of VEGF from ‘blend’ nanofibers. Figure 4 exhibits the expression of TP-T of hMSCs on TCP and nanofibers, where hMSCs on nanofibers showed significantly higher expression of TPT than cells on TCP, but there was no significant difference in protein expression between PG and PG–VEGF scaffolds (Figure 4N). We performed immunostaining of α-ACT and TP-T on neonatal rat CMs to gain an understanding of the striation patterns. The results also showed that our hMSCs were positive to the cardiac proteins (see supporting information, Figure S1); however, the distribution of the protein was somewhat different from that expressed by the mature CMs.
3.4. Expression of cardiac-specific proteins by immunostaining
4. Discussion
To demonstrate the cardiomyogenic differentiations of hMSCs, dual immunofluorescence staining of cardiacspecific proteins, α-ACT and TP-T, as well as the MSCspecific marker CD44, was carried out for cells grown on the scaffolds after 15 days of culture with 5-aza treatment. Figure 3 shows that hMSCs on nanofibrous scaffolds expressed significantly higher levels of α-ACT than cells on TCP. Higher expression of α-ACT was found for cells cultured on PGV/B and PGV/CS compared to those on TCP, indicating that VEGF could promote the cardiac differentiation of hMSCs. Furthermore, cells on PGV/B
Stem cell therapy has been widely accepted as an effective approach for treating heart disease, and MSCs with capacity for self-renewal and cardiomyogenic differentiation offer an extremely promising therapeutic strategy for myocardial regeneration. Instead of directly injecting stem cells into the heart, applying a patch of stem cellpopulated bio-graft might be more versatile in controlling the properties of the tissue graft. Furthermore, a wellengineered scaffold might serve as a ’stem cell niche’ and provide a suitable microenvironment for the cells to proliferate and differentiate into specialized mature cells,
Figure 3. Dual immunocytochemical analysis for the expression of (A–D) α-ACT and (E–H) CD44. (I–L) Merged images showing the dual expression on (A, E, I) TCP, (B, F, J) PG, (C, G, K) PGV/B and (D, H, L) PGV/CS nanofibers; scale bar = 100 μm. (M) Expression of α-ACT in CMs on nanofibers as positive control; scale bar = 40 μm. (N) Signal intensity of α-ACT expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
VEGF-encapsulated electrospun nanofibers for in vitro cardiomyogenic differentiation
Figure 4. Dual immunocytochemical analysis for the expression of (A–D) TP-T and (E–H) CD44. (I–L) Merged images showing the dual expression on (A, E, I) TCP, (B, F, J) PG, (C, G, K) PGV/B and (D, H, L) PGV/CS nanofibers; scale bar =50 μm. (M) Expression of TP-T in CMs on nanofibers as positive control; scale bar =40 μm. (N) Signal intensity of TP-T expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers
such as the CMs. In this study, we fabricated electrospun PG nanofibers containing VEGF as an artificial niche, and evaluated their potential towards the cardiomyogenic differentiation of hMSCs. In vitro cardiomyogenic differentiation of MSCs by treatment with 5-aza was first reported by Makino et al. (1999). After treatment with 3 μM 5-aza in culture medium, it was found that almost 30% of fibroblast-shaped mouse MSCs changed to a ball-like phenotype or formed the characteristic rod-shaped CMs, and they also expressed the cardiac-specific markers, such as the sarcomeres (Makino et al., 1999). Further, 5-aza treatment became a classical approach for inducing MSCs to CMs. However, most of these differentiation studies were carried out in flat culture flasks or TCP. Studies of such a kind on TCP might not be suitable for any implantations, and hence we carried out the differentiation of hMSCs to cardiomyocytes on electrospun substrates. We chose to blend a synthetic polymer, PCL, with good elasticity and biostability, with a natural polymer, gelatin, which has hydrophilic properties and excellent biocompatibility. The composite scaffold might benefit from the properties of both types of polymer and might counteract its singular defects. In our previous study, we have also demonstrated that PG nanofibers possess suitable hydrophilic and mechanical degradation properties and favour the growth of CMs (Kai et al., 2011). Here we demonstrated the capability of differentiation of hMSCs to the cardiac lineage on these nanofibrous scaffolds in the presence of small amounts of 5-aza reagent. α-ACT is an important Copyright © 2015 John Wiley & Sons, Ltd.
constituent of the contractile apparatus in the native CMs, and TP-T is a related protein that regulates the force and velocity of myocardial contraction (Di Domenico et al., 2012; Gallina et al., 2012). Immunostaining results show that MSCs on PG nanofibers expressed significantly higher levels of cardiac-specific proteins than those on TCP. Our quantification results showed that the expression of αACT by the MSC-differentiated CMs on PG was 73.2% higher than its expression on TCP, while TP-T was expressed three times higher than on TCP, indicating that the cardiac substrate with nanofibrous topography could provide additional chemical and mechanical cues to promote the cardiac differentiation of MSCs. Similarly, Yang et al. (2009) found that the differentiated rat MSCs cultured on silk fibroin–chitosan–hyaluronic acid cardiac patches expressed higher cardiac proteins (TP-T, cardiotin and connexin 43) and improved the expressions of cardiac genes (Gata4, Nkx2.5, Tnnt2 and Actc1) compared to TCP and pure silk fibroin scaffolds. To improve cardiac differentiation, VEGF was incorporated in the nanofibers by two different approaches, blend spinning and core–shell spinning. VEGF is as an important angiogenic growth factor, has been used in improving cardiac function after MI in animal models and has been shown to stimulate angiogenesis with the formation of vascular structures in the infarcted area (Chiu and Radisic, 2010; Guo et al., 2012). During our study, PGV/B had a reduced fibre diameter of approximately 199 ± 37 nm, a reduced water contact angle as well as reduced mechanical properties compared to PG (Table 1). On the other hand, J Tissue Eng Regen Med (2015) DOI: 10.1002/term
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encapsulating VEGF within the core of the PG nanofibers did not significantly change the mechanical properties of the nanofibers, except for a slight decrease in its elongation at break compared to PG (Table 1). In TE the speed of cell growth is important, as it is related to the speed of regeneration of new tissue, and a rapid proliferation rate would obviously result in an improved tissue development. Figure 2B shows that the VEGF-containing nanofibers significantly promoted the proliferation of hMSCs compared to pure PG nanofibers. Similarly, Huang et al. (2010) demonstrated that the profiled addition of VEGF in the culture medium increased the proliferation of mouse MSCs by 1.9-fold. The coinjection of MSCs with VEGF into MI hearts was also described to increase the grafted cell viability and improved cardiac function (Pons et al., 2008). VEGF could not only promote cell proliferation but, more importantly, regulate cellular differentiation. As shown in Figure 3, cells on both PGV/B and PGV/CS expressed higher levels of αACT than those on TCP and PG nanofibers, and the increased protein expression of cells on VEGF-containing nanofibers further confirmed that critical function of VEGF in the developmental stages of cardiomyogenesis (see supporting information, Figure S2). Chen et al. (2006) previously reported that adding VEGF to the culture medium significantly enhanced α-MHC, Troponin-I and Nkx2.5 expression in differentiated embryonic stem cells. Later, studies by Song et al. (2007) also showed that VEGF was a critical factor for the spontaneous differentiation of adipose tissue-derived stem cells, and the growth factor was responsible for the expression of cardiac markers via a paracrine mechanism. Here we demonstrated the positive function of VEGF towards the
cardiomyogenic differentiation of hMSCs, in a different perspective, by performing the experiments on VEGF incorporated PG nanofibers. Our study shows that the PG– VEGF substrates hold great potential application as a bioengineered cardiac patch for myocardial regeneration.
5. Conclusion PG, PGV/B and PGV/CS nanofibrous scaffolds were fabricated to evaluate their effects on the growth and cardiomyogenic differentiation of hMSCs in vitro. The incorporation of VEGF into the nanofibers improved the proliferation of MSCs and, more importantly, improved the expression of cardiac-specific markers. After 15 days, hMSCs cultured on PGV/B and PGV/CS scaffolds showed higher expression of α-ACT and TP-T proteins compared to cells grown on pure PG nanofibers, indicating the potential of VEGF-encapsulated PG nanofibers as a suitable scaffold for myocardial regeneration.
Conflict of interest The authors have declared that there is no conflict of interest.
Acknowledgements This research was supported by NRF–Technion (Grant No. WBS R-398-001-065-592) and the Nanoscience and Nanotechnology Initiative in NUS, Singapore.
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Supporting information on the internet The following supporting information may be found in the online version of this article: Figure S1. Dual immunocytochemical analysis for the expression of myosin heavy chain (MHC) and CD44; merged images showing dual expression on TCP, PG, PGV/B and PGV/CS nanofibers; expression of MHC in CMs on nanofibers as positive control; and signal intensity of MHC expressed in differentiated MSCs on TCP and PG, PGV/B and PGV/CS nanofibers Figure S2. Normalized gene expression of Tnni3 in differentiated hMSCs on TCP and PG, PGV/B and PGV/CS nanofibers with 5-aza treatment
Copyright © 2015 John Wiley & Sons, Ltd.
J Tissue Eng Regen Med (2015) DOI: 10.1002/term
