Effective naked plasmid DNA delivery into stem cells by microextrusion-based transient-transfection system for in situ cardiac repair

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Effective naked plasmid DNA delivery into stem cells by microextrusionbased transient-transfection system for in situ cardiac repair Nien-Chi Huanga, Chii-Ming Leeb, Shan-hui Hsua,c,d,* a

Institute of Polymer Science and Engineering, National Taiwan University, Taipei, Taiwan, R.O.C. Division of Cardiology, Department of Internal Medicine, National Taiwan University Hospital, Taipei, Taiwan, R.O.C. c Center of Tissue Engineering and 3D Printing, National Taiwan University, Taipei, Taiwan, R.O.C. d Institute of Cellular and System Medicine, National Health Research Institutes, Zhunan, Taiwan, R.O.C. b

A R T I C L E

I N F O

Article History: Received 28 August 2019 Accepted 3 December 2019 Available online xxx Key Words: cardiac repair cell reprogramming in situ transfection naked plasmid delivery polyurethane hydrogels

A B S T R A C T

Background aims. Combining the use of transfection reagents and physical methods can markedly improve the efficiency of gene delivery; however, such methods often cause cell damage. Additionally, naked plasmids without any vector or physical stimulation are difficult to deliver into stem cells. In this study, we demonstrate a simple and rapid method to simultaneously facilitate efficient in situ naked gene delivery and form a bioactive hydrogel scaffold. Methods. Transfecting naked GATA binding protein 4 (GATA4) plasmids into human umbilical cord-derived mesenchymal stem cells (hUC-MSCs) by co-extruding naked plasmids and hUC-MSCs with a biomimetic and negatively charged water-based biodegradable thermo-responsive polyurethane (PU) hydrogel through a microextrusion-based transient-transfection system can upregulate the other cardiac marker genes. Results. The PU hydrogels with optimized physicochemical properties (such as hard-soft segment composition, size, hardness and thermal gelation) induced GATA4-transfected hUCMSCs to express the cardiac marker proteins and then differentiated into cardiomyocyte-like cells in 15 days. We further demonstrated that GATA4-transfected hUC-MSCs in PU hydrogel were capable of in situ revival of heart function in zebrafish in 30 days. Conclusions. Our results suggest that hUC-MSCs and naked plasmids encapsulated in PU hydrogels might represent a new strategy for in situ tissue therapy using the microextrusion-based transient-transfection system described here. This transfection system is simple, effective and safer than conventional technologies. © 2019 International Society for Cell and Gene Therapy. Published by Elsevier Inc. All rights reserved.

Introduction Gene therapy involves the delivery of biological macromolecules (e.g., DNA, RNA, and protein) to target cells to make them capable of inducing cell transformation or the expression of certain proteins for the treatment of diseases through intracellular encoding [1]. However, these exogenous molecules are difficult to transfer into cells without vectors and are often rapidly degraded by nucleases or proteases [2]. Therefore, viral and nonviral vectors have been developed to protect and enhance the delivery of these molecules into the target cells [3,4]. Although viral vectors exhibit high transfection efficiency of small DNA fragments, the associated safety concerns and immunogenicity present clinical issues that remain unresolved [5]. Nonviral vectors or transfection reagents influence transfection efficiency but also cause cell cytotoxicity [5]. It remains difficult to deliver naked plasmids into stem cells without using a vector. * Correspondence: Shan-hui Hsu, Institute of Polymer Science and Engineering, National Taiwan University, No. 1, Sec. 4, Roosevelt Road, Taipei 10617, Taiwan, R.O.C. E-mail address: [email protected] (S.-h. Hsu).

In addition to vector-based delivery, physical or mechanical methods (such as microinjection, electroporation and magnetoreception) have been used to enhance the delivery of biological molecules. However, these methods might result in alterations in the permeability and structural integrity of the cell membrane, thereby causing cytotoxicity [6,7]. Other physical transfection platforms, such as a microextrusion-based device, can produce a bioactive scaffold comprising a combination of cells or macromolecules, can facilitate continuous and localized release exogenous genes, and can print tissue analogues capable of engraftment with endogenous tissue [8 12]. Microfluidicbased devices enhance in vitro and in vivo macromolecule delivery efficiency and cell viability [13 16]. Furthermore, a microfluidicbased device enhanced the delivery efficiency of naked biological macromolecules into primary cells compared with an endocytosis group by squeezing cells [13]. Cell squeezing methods in combination with microfluidic and microextrusion-based devices might provide a new strategy for gene delivery into the cell nucleus. The materials necessary for generation of bioactive scaffolds must be highly injectable and have suitable physicochemical properties such as gel formation dynamics and crosslinking density. They should

https://doi.org/10.1016/j.jcyt.2019.12.003 1465-3249/© 2019 International Society for Cell and Gene Therapy. Published by Elsevier Inc. All rights reserved.

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also have similar mechanical and degradation properties to mimic the respective architecture in organs (such as native architecture and cell composition) [17]. The materials should be well suited for cell encapsulation and provide a suitable microenvironment and nanotopography for cell adhesion, proliferation and differentiation [8 25]. Synthetic hydrogels are widely used to form bioactive scaffolds due to their advantageous physicochemical properties (e.g., contractility, density, hardness and flexibility). It is easy to control their composition, and they also have lower immunoreactivity [26 29]. In our previous studies, waterborne biodegradable polyurethane (PU) hydrogels have been used in bioprinting for reprogramming human fibroblasts into neural-like cells, owing to their outstanding biocompatibility, high degree of water content and the specific viscoelastic properties [30 32]. However, the efficiency of in situ delivery using a microextrusion-based device and PU hydrogels and tissue function repair by naked plasmids transfected into human mesenchymal stem cells (hMSCs) remain unclear. On the basis of the advantages of microextrusion-based device and PU hydrogels, we applied these techniques for in situ gene delivery and cardiac repair. In this study, we used a self-developed microextrusion-based transient-transfection system to deliver naked GATA binding protein 4 (GATA4) plasmids. GATA4 is transcription factor involved in cardiac development in the embryo and promotes the binding of other cardiac transcription factors [33 35], into human umbilical cord-derived mesenchymal stem cells (hUC-MSCs). Both hUC-MSCs and naked plasmids were encapsulated into biomimetic and negatively charged water-based biodegradable thermo-responsive PU hydrogels (PU1 and PU2) before gelation in the absence of cross-linkers. The hUC-MSCs- and naked GATA4 plasmid-encapsulating hydrogel was extruded by the microextrusion-based transienttransfection system to generate three-dimensional bioactive scaffolds, where hUC-MSCs might be reprogrammed in situ during the extruding process. Furthermore, we speculated that the elasticity of the PU hydrogel might promote a squeezing effect on hUC-MSCs by a micronozzle to enhance the efficiency of naked plasmid delivery during the extrusion period. We then evaluated cell viability and cardiac gene expression in vitro and used a heart-defect model of adult zebrafish to assess the in situ cardiac repair potential of the hUCMSC- and naked GATA4 plasmid-encapsulated PU hydrogels. Our findings suggested that this innovative technology and its potential ability to reprogram cells might have prospective applications for in situ naked gene delivery and tissue therapy. Methods Synthesis of negatively charged water-based biodegradable thermoresponsive PU The hUC-MSC- and plasmid-encapsulating hydrogel was formed by heating PU nanoparticle (NP) dispersion. The preparation process for two types of PU NPs, PU1 and PU2, is shown in Figure 1A. For PU1, soft segments were synthesized by poly (e-caprolactone) diol (PCL diol MW: »2000 g/mol) and D,L-polylactide diol (PDLLA diol MW: »1500 g/mol; PDLLA1). For PU2, soft segments were synthesized by PCL diol and DLLA diol (MW: »2000 g/mol; PDLLA2). These PDLLA diols were synthesized from D,L-lactide, 1,3-propanediol and stannous octoate (0.05%) at 95°C for 8 h (PDLLA1) and 10 h (PDLLA2), followed by purification in ethanol. The hard segment comprised isophorone diisocyanate (IPDI), two-chain extender 2,2-bis(hydroxymethyl) propionic acid (DMPA) and ethylenediamine (EDA). The oligodiols, IPDI and the 0.03% catalyst tin(II)2-ethylhexanoate were placed in a round-bottomed, four-necked flask with continuous stirring at 180 rpm. The reaction temperature was between 70°C and 80° C until the diols were dissolved and mixed. After 3 h, DMPA and methylethylketone (MEK) were added to the flask with constant stirring at 180 rpm. After another 1 h, the temperature was reduced to

50°C, and then triethylamine (TEA) was added to the flask to neutralize the carboxylic acid groups. After 30 min, the temperature was decreased to 45°C, and the speed of the stirrer was increased to 1100 rpm. Deionized water was added to disperse the prepolymer, and the chain-extender EDA was dropped into the flask to complete the reaction. The stoichiometric ratio of IPDI:oligodiols:DMPA:EDA: TEA was 3.52:1:1:1.52:1. The residual MEK and TEA were removed by vacuum distillation at 80°C for 2 h. Physical characterization The measurement of hydrodynamic diameter of the PU NPs was obtained by diluting the dispersion to 3000 ppm by adding deionized water, followed by dynamic light scattering assessment (165° backscatter). The surface zeta potential of the PU NPs was determined using electrophoretic light scattering instrument (Delsa Nano C; Beckman Coulter, Brea, CA, USA). The dynamic rheological properties of the NP dispersions were investigated at 37°C using a cone-and-plate rheometer (RS-5; TA Instruments, New Castle, DE, USA). The diameter of the cone was 40 mm, and the angle was 2°. The NP dispersion (700 mL) was loaded, and the configuration of the cone and plate was maintained at 1% shear strain with a 1-Hz frequency. The storage modulus (G0 ) and loss modulus (G00 ) were measured against time. The point of sol gel transition was defined when G0 started to exceed G00 . Construction of cardiac-specific plasmids Plasmids containing human cardiac-specific sequences were constructed as follows. We used pTrace-CMV/Bsd vector that encodes for cytomegalovirus (CMV) promoter and green fluorescent protein (GFP). The human GATA-binding protein 4 (GATA4) coding sequence was amplified by polymerase chain reaction (PCR) using the cDNA from hMSCs as template and specific primers. The amplified sequences were confirmed by agarose gel electrophoresis. To construct the plasmid, the GATA4 gene was ligated into the pTrace-CMV/Bsd vector following restriction digestion with EcoRI and EcoRV. The GATA4 plasmids were then cloned into Escherichia coli and extracted using a plasmid mini-prep purification kit (GeneMark, Taipei, Taiwan). The extracted clones were confirmed by agarose gel electrophoresis. The concentration and purity of the plasmids were determined spectrophotometrically (NanoDrop 1000; Thermo Fisher Scientific, Wilmington, DE, USA). Plasmids with absorbance A260/A280 ratio of 1.8-2.0 were considered to be of sufficient purity for use in subsequent transfection studies. The base composition of the cloned gene was confirmed by sequencing (Supplementary Figure 1). Culture of hUC-MSCs The first passage of hUC-MSCs was provided by BIONET Corp. (Taipei, Taiwan). The identity of the hUC-MSCs was confirmed by flow cytometry using a panel of 19 surface markers (Supplementary Figure 2). The hUC-MSCs were cultured in a-minimum essential medium (complete medium) (Invitrogen, Carlsbad, CA, USA) containing 10% fetal bovine serum (Gibco; Thermo Fisher Scientific), 1.7 g/L Na2CO3 and 1% penicillin-streptomycin solution (Invitrogen) at 37°C with 5% CO2. The medium was changed three times every week. The hUC-MSCs from passages 5 through 8 were used for subsequent experiments. Cell labeling For studying cell morphology, hUC-MSCs were labeled with red fluorescent dye PKH26 according to manufacturer’s instructions (Sigma-Aldrich, St. Louis, MO, USA). The hUC-MSCs at a concentration of 1 £ 107 cells/mL were labelled by mixing with 2 £ 10 6 mol/L PKH26. The reaction was stopped by adding the complete medium.

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Figure 1. (A) Synthesis of water-based biodegradable polyurethane (PU) dispersions (PU1 and PU2). Chain extender: 2,2-bis(hydroxymethyl) propionic acid (DMPA) and ethylenediamine (EDA). Hard segment: isophorone diisocyanate (IPDI). Soft segments: poly(e-carpolactone)diol (PCL diol), D,L-polylactide diol (PDLLA diol MW: »1500 g/mol; PDLLA1), and DLLA diol (MW: »2000 g/mol; PDLLA2). Waterborne biodegradable polyurethane (PU1 and PU2). Triethylamine (TEA). (B) The rheological properties of two PU dispersions, 25% PU1 and 25% PU2, as a function of time (time sweep) in water after placing at 37°C. Time sweep tests were performed at 1-Hz frequency and 1% shear strain.

Labeled hUC-MSCs were washed with phosphate-buffered saline (PBS) and used for viewing. Transfection by microextrusion-based transient-transfection system The PU NP dispersion with a solid content of 25% in serum-free medium was preheated at 60°C for 10 min. Subsequently, the temperature was decreased to 37°C and the PU dispersion was mixed with hUCMSCs (2.5 £ 106 cells/20 mL serum-free medium) and 150 mg/mL GATA4 plasmids before gelation. The hydrogel scaffolds were fabricated using a self-developed microextrusion-based fused-deposition manufacturing instrument that included a handheld microextrusion-

injection system, a computer and an x-y-z motion platform with a heater [31]. The computer was used to design and construct the scaffolds, paths and motion of the platform. A 200-mm nozzle was fitted with a 5-mL syringe and connected to a pressure controller at 55 kPa for use as the handheld injection system. The hUC-MSCs and naked GATA4 plasmids were co-extruded with hydrogels and then printed as crossing fibers with a fiber diameter of »200 mm and a width £ length of »1.5 £ »1.5 cm into a 3-cm Petri dish placed on the 37°C platform. After printing, the hydrogels were immersed in serum-free medium, which was changed to complete medium after 24 h. After another 48 h, samples were harvested for assaying. For the tissue culture polystyrene (TCPS) group, 2.5 £ 106 cells in serum-free medium with 150 mg/mL

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GATA4 plasmids were seeded in 3-cm cell culture dishes through a 200mm nozzle. After 24 h, the serum-free medium was changed with complete medium with 150 mg/mL GATA4 plasmids. After another 48 h, cells were harvested for assays. For the PolyFect (a dendrimer-based transfection reagent; Qiagen, Hilden, Germany) group, cells were transfected with plasmid/PolyFect complexes according to manufacturer protocol, and 2.5 £ 106 hUC-MSCs were seeded in a 3-cm cell culture dishes with complete medium. After cell attachment, the complete medium was replaced with the complete medium containing 150 mg GATA4 plasmids and 30 mg of PolyFect. After another 24 h, the medium with plasmid/PolyFect complexes was changed with the complete medium. After 72 h, the cells were harvested. Induction of cardiomyogenesis differentiation For preparation of the induction medium for cardiomyogenesis, 10 mmol/L of 5-azacytidine was added to the complete medium. Once the hydrogel scaffolds were printed as crossing fibers and cultured for 3 days, the medium was changed to the induction medium. After 3 days of treatment, the medium was regularly changed with the complete medium three times at 3-day intervals. The total culture period was 1 day in serum-free medium, then 2 days in complete medium, followed by 3 days in induction medium, and finally 9 days in complete medium. For the TCPS and PolyFect groups, cells were cultured on TCPS in the same pattern. Subsequently, gene and protein expressions were analyzed. Determination of cell viability After 1 and 3 days of incubations, the hUC-MSCs were stained with calcein AM and ethidium homodimer-1, according to manufacturer’s kit protocol (Live/Death viability/cytotoxicity kit, Invitrogen). Samples were washed with PBS and then stained with the staining solution (5 mL calcein AM and 20 mL ethidium homodimer-1 in 10 mL PBS) for 1 h at 25°C. Using fluorescence microscope (Leica, Wetzlar, Germany), dead and live cells were identified by their red and green fluorescence respectively, at days 1 and 3 of experiment. Cell viability was examined by 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide (MTT; Sigma-Aldrich) assay at days 1, 2 and 3. Samples were washed with PBS and stained with 0.5 mg/mL MTT/PBS for 4 h at 37°C. Crystal violet was dissolved by immersion with dimethyl sulfoxide. The absorbance of crystal violet was determined using a microplate reader at 540 nm (SpectraMax M5; Molecular Devices, Sunnyvale, CA, USA). Data at days 1, 2 and 3 were normalized to the values at day 0 to evaluate cell viability. The amount of hUC-MSCs in the PU hydrogel after 1, 2 and 3 days of culture was determined using DNA quantitation kit (Sigma-Aldrich). The result was calculated from a DNA standard curve of identical cells and expressed as DNA content. Gene expression analysis Real-time reverse-transcription polymerasechain reaction (RTPCR) was used to analyze gene expression. Details of primer sequences are presented in Supplementary Table I. The samples were analyzed for the expression of GATA4, myocyte enhancer factor 2C (MEF2C), T-box 5 (TBX5) and glyceraldehyde 3-phosphate dehydrogenase (GAPDH) at day 3 and GATA4, MEF2C, TBX5, NK-2 transcriptionfactor-related locus 5 (NKX2.5), myosin heavy chain 6 (MYH6), troponin T2 (TNNT2) and GAPDH at day 15. Total RNA was extracted from the cells by adding Trizol reagent (Invitrogen). The cDNA was synthesized by the RevertAid first-strand cDNA synthesis kit (MBI Fermentas, St. Leon-Rot, Germany) and amplified by PCR at 42°C for 60 min and 72°C for 10 min. RT-PCR was performed with 25 mL of 2 £ SYBR Green PCR master mix (Finnzymes Oy, Espoo, Finland) and 1 nmol/L of forward and reverse primers using Chromo 4 PTC200 thermal cycler (MJ Research, Waltham, MA, USA). The following conditions

were used: 30 cycles, denaturation at 94°C for 30 s, annealing at 62°C for 30 s, and extension at 72°C for 50 s. Expression levels were analyzed and normalized to the housekeeping gene GAPDH. Immunofluorescence Samples were washed with PBS, fixed with 4% paraformaldehyde (PFA)/PBS for 30 min and permeabilized with 2.5% Triton X-100 in PBS (PBST) for 30 min. After blocking with 1% bovine serum albumin (BSA) in PBST for 24 h, the samples were incubated with primary antibodies for GATA4, NKX2.5, MYH6, and zonula occludens-1 (ZO-1; Santa Cruz Biotechnology, Dallas, TX, USA) at 4°C for 24 h. The samples were then washed five times with PBS to remove unbound antibodies and immersed in the appropriate secondary antibody solutions for 24 h at 4°C. After subsequent washing with PBS for three times, cell nuclei were stained with 40 ,6-diamidino-2-pheny lindole and observed by fluorescence microscope. Data was saved in 8-bit tiff image format. The semi-quantification of representative marker protein expression level was expressed by counting the number of cells with red fluorescence in a 200-mm square. Each sample was randomly selected from 10 microscopic fields under a 100 £ magnification. Western blot For Western blot analysis of GATA4, NKX2.5, MYH6, ZO-1 and GAPDH, samples were homogenized in RIPA lysis buffer containing protease inhibitors (PhosSTOP, Roche). After centrifugation of cell lysates (13 000 rpm for 40 min), supernatants were collected. Total proteins from each sample were separated on 8% sodium dodecyl sulfate polyacrylamide gel electrophoresis. Western blot was then performed for GATA4 (EPR4768, ab134057), NKX2.5 (2E1, ab91196), MYH6 (EPR10891(2), ab185967; Abcam, Cambridge, MA) and ZO-1 (61-7300; Thermo Scientific, Amsterdam, the Netherlands). The separated proteins on the gel were blotted to polyvinylidene difluoride membranes. The membranes were blocked with 5% BSA in PBST at room temperature for 1.5 h to avoid unspecific binding and then immersed in the primary antibody solution specific for each protein at 4°C for 24 h. GAPDH antibody (14C10; Cell Signaling Technology, Danvers, MA, USA) was used as the internal control. The samples were then washed with PBS to remove unbound antibodies and immersed again in the appropriate horseradish peroxidase conjugated secondary antibody solutions at room temperature for 1.5 h. ECL DualVue Western Blotting Marker (GE Healthcare, Little Chalfont, Buckinghamshire, UK) was finally added and the images were obtained using a chemiluminescence analyzer (UVP Biospectrum Imaging System; UVP, Inc., Upland, CA). In vivo experiments The method to create the heart defect in zebrafish was reported by Gonzalez-Rosa et al. [36]. Briefly, adult zebrafish (wild-type AB strain) between the ages of 6 and 18 months and harboring a heart defect were used for in vivo experiments. Zebrafish were immersed in 0.04% tricaine (Sigma-Aldrich) and immobilized with the ventral portion facing upward into a foam holder. Forceps were used to tear the tissue through the body wall and pericardium, and the heart was exposed by gently squeezing the abdomen. A copper filament (length: »1.5 cm; diameter: »250 mm) was immersed in liquid nitrogen for 5 min and placed on the heart surface for 10 s to cause the defect. The sample was then injected into the heart ventricle using the self-developed microextrusion-based fused-deposition manufacturing instrument with a handheld microextrusion-injection system with a 90 mm nozzle. The sample included hUC-MSCs mixed with each type of PU hydrogels (PU1 and PU2) and PBS and hUCMSCs and naked plasmids mixed with each type of PU hydrogels and PBS. For each group, six zebrafish were used for in vivo experiments. After the operation, the fish were placed in a tank with fresh water

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and reanimated until the fish were swimming normally after 30 min. Zebrafish appeared less active and uncoordinated for the first few days after the operation [37]. Zebrafish with hearts not operated on were used as the normal group. All animal experiments were performed according to regulations approved by the Animal Ethical Committee of National Taiwan University (NTU105-EL-00152). The heartbeat rate of the zebrafish was counted before sacrifice (at day 30) for 30 sec at 25 § 0.2°C, and the counter was multiplied by 2 to yield beats per minute. Six zebrafish hearts from each group were counted. The number of zebrafish heartbeats was expressed as beats per minute [38]. Zebrafish were euthanized at day 30 by immersion in 0.16% tricaine, and the heart was dissected for histological staining. Histological staining Six zebrafish hearts from each group were used for histological staining. Zebrafish hearts were fixed in 4% PFA/PBS solution at 4°C for a day and washed with PBS. The heart samples were then dehydrated with 20%, 40%, 60%, 80% and 95% sucrose solution for 1 h, immersed in Hemo-De (Scientific Safety Solvents, Keller, TX, USA) for 1 h, and embedded in paraffin wax at 60°C overnight. The samples were sectioned into 3-mm sections, and dried at 37°C for 1 day. The sections were deparaffinized by immersion in Hemo-De, rehydrated with 95%, 80% and 75% sucrose solution for 1 h and washed with distilled water. The sections were subsequently stained with hematoxylin-eosin. The quantification of zebrafish heart size and trabecular thickness were determined by ImageJ. Data were then saved in 8-bit tiff image format to count the percentage of trabeculae and cavity in whole heart. Statistical analysis Data are expressed as the mean § standard deviation. Multiple samples were used in each in vitro and in vivo experiment. Reproducibility was confirmed in a minimum of three to six independent experiments and using cells from at least three donors. Statistical differences were determined by one-way analysis of variance. Differences were considered significant at a P < 0.05. Results Physical properties of PU NP dispersions The results of particle size and zeta potential of the PU NP dispersions are displayed in Table 1. The hydrodynamic diameter of PU1 NPs (46.0 § 1.8 nm) was slightly higher than that of PU2 NPs (40.2 § 1.5 nm). The zeta potentials of the PU1 and PU2 NPs were 39.6 § 3.8 mV and 55.3 § 2.6 mV, respectively. The rheological properties of PU1 and PU2 are shown in Figure 1B. The sol gel transition was observed within 15 min for PU1 and PU2 NP dispersions, with PU1 and PU2 displaying phase transitions at 194 s and 769 s, respectively. After 20 min, the shear storage modulus of PU1 (»13 000 Pa) was higher than that of PU2 (»900 Pa). Cell morphology and viability The morphologies of hUC-MSCs in hydrogels were observed for 1 and 3 days, and the results are shown in Figure 2A. The images Table 1 Physical properties of the PU nanoparticle dispersions.

Hydrodynamic diameter (nm) Zeta potential (mV) Phase transition (s) Shear storage modulus (Pa)

PU1

PU2

46.0 § 1.8 39.6 § 3.8 194 »13,000

40.2 § 1.5 55.3 § 2.6 769 »900

5

revealed that the hUC-MSCs were printed with PU hydrogels during the extrusion process. hUC-MSCs in both PU hydrogels exhibited spheroid morphology and did not display spike-like membranes or vesicles, both of which were signs of apoptosis at day 1 and day 3. For the TCPS and PolyFect groups, hUC-MSCs revealed a well-spread morphology; however, those in the PolyFect group displayed lower cell number compared with other groups at days 1 and 3. The cell death in PolyFect group may be caused by the excessive and lethal damage to the membranes of cells and aggregation of the transfection reagents in cellular organelles after dissociation with plasmids [39 41]. The samples were also stained with live/dead fluorescence kit and the images are displayed in Figure 2B. A greater number of live cells was observed when hUC-MSCs were cultured in PU1 hydrogel and the TCPS group, compared with that in PU2 hydrogel and the PolyFect group. The cell viability in PU hydrogel, TCPS and PolyFect groups are presented in Figure 2C. At day 1, the viability of hUC-MSCs in PU1 (»100%) was similar to that of hUC-MSCs on TCPS (»95%), whereas hUC-MSCs in PU2 showed lower viability (»50%). At day 2, the viability of cells in PU1 (»160%) was similar to that of cells on TCPS (»155%), followed by cells in PU2 (»95%). At day 3, cells in PU1 displayed the highest cell viability (»340%) compared with those in PU2 (»220%) and the TCPS group (»190%). For the PolyFect group, cell viability was lower than that of the other groups during the observation period (<25%). The amount of DNA content showed the same tendency as the cell viability analyzed by MTT assay after 1, 2 and 3 days of culture (Figure 2D). Previous studies reported that the modulus of hydrogel may affect the viability and morphology of cells [42]. Here, we tried to explain the relationship between cell proliferation and hydrogel stiffness. Our results were consistent with previous reports regarding the higher proliferation rate of hUC-MSCs in PU1 hydrogels (»13 000 Pa, the stiffer hydrogel) than that of cells in PU2 hydrogels (»900 Pa). Expression of cardiac marker genes The expression of GATA4, MEF2C and TBX5 genes in untransfected and GATA4-transfected groups for 3 days is shown in Figure 3A. The expression levels of cardiac marker genes in untransfected hUC-MSCs in PU1 and PU2 hydrogels (»0.4) were similar to those in the TCPS group (»0.3). GATA4-transfected hUC-MSCs showed higher cardiac gene expression (»1.8) than any of the untransfected groups (»0.4). GATA4-transfected hUC-MSCs in PU1 hydrogels displayed higher gene expression levels (»1.8) than those in PU2 hydrogels (»1.0) and the TCPS group (»0.5). Furthermore, the cardiac gene expression levels from GATA4-transfected hUC-MSCs in PU1 hydrogels were similar to those in the PolyFect group (»1.8). These results indicate that the transfection efficiency of hUC-MSCs and GATA4 plasmids co-extruded with PU1 hydrogels was greater than that with PU2 hydrogels. In particular, the transfection efficiency for transfected hUC-MSCs in PU1 hydrogels reached a similar level as that of the PolyFect group. The expression levels of cardiac marker genes (GATA4, MEF2C, TBX5, NKX2.5, MYH6 and TNNT2) in untransfected and GATA4-transfected groups at day 15 are shown in Figure 3B. Untransfected hUCMSCs in PU1 hydrogels displayed slightly higher expression of cardiac genes (»50) than hUC-MSCs in PU2 hydrogels (»40), followed by the TCPS group (»30). The expression levels of cardiac genes in both GATA4-transfected groups (greater than »90) were higher than that in any of the untransfected groups (»50). GATA4-transfected hUC-MSCs in PU1 hydrogels (»180 1200) displayed the highest gene expression levels, followed by those in PU2 hydrogels (»150 800) and the TCPS (»50 100) group. Moreover, the expression levels of cardiac genes in transfected hUC-MSCs in PU1 hydrogels were similar to those in the PolyFect group (»180 1200). For GATA4-transfected cells in PU1 hydrogels in particular, the gene expression levels of NKX2.5 (»700), MYH6 (»1100) and TNNT2

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Figure 2. (A) The morphology of human umbilical cord derived mesenchymal stem cells (hUC-MSCs) cultured on TCPS and those cultured in PU1 and PU2 hydrogels. Cells were stained with red fluorescent dye (PKH26). (B) The live/dead stain of hUC-MSCs. Live cells were stained with green fluorescent dye (calcein AM), and dead cells were stained with red fluorescent dye (ethidium homodimer-1). Scale bar: 200 mm. (A and B) Examined at day 1 and 3 after cell seeding. (C) Cell viability was examined by 3-(4,5-dimethylthiazol-2yl)-2,5-diphenyltetrazolium bromide (MTT) assay. (D) DNA quantification analysis was carried out at days 1, 2 and 3. **P < 0.01, ***P< 0.001.

(»350) were higher than those of other cardiac marker genes (»230). We then found that the expression level of cardiac genes in naked GATA4 plasmid-transfected hUC-MSCs in PU1 hydrogels was higher than that of naked GATA4 plasmid-transfected cells in PU2 hydrogels and the TCPS group. Immunofluorescence images and Western blotting Immunofluorescence images of cardiac marker proteins (GATA4, NKX2.5, MYH6 and ZO-1) are shown in Figure 4A. Semi-quantification of representative marker proteins is presented in Figure 4B.

Untransfected hUC-MSCs showed weak expression and intensity (»20%) for all cardiac marker proteins according to staining results for all groups. GATA4-transfected hUC-MSCs were positive for GATA4, NKX2.5, MYH6 and ZO-1 expression. Furthermore, the expression levels in the PU1 and PolyFect groups were similar (»75%) and also higher than those in the PU2 group (»55%). For the TCPS group, these marker proteins displayed weaker intensity (»40%). Western blots of GATA4, NKX2.5, MYH6 and ZO-1 are demonstrated in Figure 4C, and the quantification conducted is shown in Figure 4D. The expression levels of all marker proteins in the PU1 and PolyFect groups were higher than those in the PU2 group. For the untransfected and TCPS groups, these

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Figure 3. (A) The gene expression of GATA4, MEF2C and TBX5 for untransfected and naked GATA4-transfected hUC-MSCs cultured on TCPS and those cultured in PU hydrogels at day 3. (B) The expression levels of specific cardiac marker genes (GATA4, MEF2C, TBX5, NKX2.5, MYH6 and TNNT2) for the untransfected and transfected groups at day 15. The expression levels were normalized to the housekeeping gene GAPDH. *P < 0.05, **P < 0.01, ***P < 0.001.

proteins displayed weak intensity relative to all marker proteins. These results indicated that GATA4-transfected hUC-MSCs in PU hydrogels exhibited differentiation into cardiomyocyte-like cells. Heartbeat rate and histology of zebrafish heart The heartbeat rate of zebrafish injected with untransfected and GATA4-transfected hUC-MSCs is shown in Figure 5A. The heartbeat

rate of the normal group (without liquid nitrogen treatment) was »58 beats/min. For the untransfected hUC-MSC-PU hydrogel-treated group, the heartbeat rate was »25 beats/min. For the GATA4-transfected group, the heartbeat rate following transfer of cells in PU1 (»55 beats/min) was higher than that associated with cells in PU2 (»38 beats/min) and PBS (»30 beats/min). The histology of zebrafish hearts from the untransfected and GATA4-transfected groups are shown in Figure 5B. The quantification

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Figure 4. (A) The images of immunofluorescence staining, (B) the semi-quantification of immunofluorescence intensities, (C) representative photographs of Western blotting bands and (D) semi-quantification of the specific protein expression levels for untransfected and naked GATA4-transfected hUC-MSCs on TCPS and those in PU hydrogels. *P < 0.05, **P < 0.01, ***P < 0.001. All cells were analyzed at day 15.

of heart size, trabecular network and the percentage of heart cavity and trabeculae from the histology images by ImageJ are displayed in Figure 5C E. For the untransfected hUC-MSC-PBS-treated group, the size of zebrafish hearts was significantly smaller, and the percentage

of cavity in whole heart was larger than that observed in the normal and PU1 hydrogel groups. The trabecular thickness of the PBS group was lowest in all groups. For the untransfected-PU1 group, the size of heart and the percentage of trabeculae in whole heart were larger

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Figure 5. (A) The heartbeat rate. (B) Histological images of zebrafish injected with PBS and each type of PU hydrogels with untransfected or GATA4-transfected hUC-MSCs. Scale bar: 1 mm. The quantification of (C) heart size and (D) trabecular thickness. (E) Percentage of trabeculae and cavity in whole heart from the histological images. ns = non-significant, *P < 0.05, **P < 0.01, ***P < 0.001.

than that of the normal group. The trabecular thickness of the untransfected-PU1 group was more than that of the untransfectedPU2 group. For the untransfected-PU2 group, the heart size was larger than the untransfected-PBS group but was smaller than the normal and untransfected-PU1 groups. The cavity in whole heart was larger than that of the normal and untransfected-PU1 groups, and the trabecular thickness was lesser than that of the normal and untransfected-PU1 groups. The heart morphologies following transfer of GATA4-transfected cells with PBS and GATA4-transfected cells in PU2 showed a larger heart cavity, whereas those associated with PU1 were similar to the normal group. The heart size of the GATA4transfected PU1 group was similar to that of the normal group, and the trabecular thickness was almost similar to the normal group. These results indicated that transfer of GATA4-transfected hUC-MSCs encapsulated by PU1 hydrogels was capable of reviving heart function in zebrafish. Discussion Studies have reported the efficient transfection of plasmids into cells through temporarily applying a squeezing force by microextrusion and microfluidic-based device [8,15,16,20]. A microfluidic-based device enhances the delivery efficiency of naked biological macromolecules by allowing the cells to pass through a 6-mm channel [13]. Additionally, the delivery of GFP plasmids into Chinese hamster ovary cells using a thermal syringe needle of 85 mm diameter exhibited low apoptosis (»3.5%) and high transfection efficiency (»31.5%), reflecting the self-repair capacity of the cell membranes [15]. The shear stress and heat generated by the nozzle size and firing of the transfection device might cause short-term disruption of the cell membrane while allowing the complexes to enter the cells through transient pores [15]. Moreover, the degree of deformation and selfrepair capacity of the cell membrane affects cell transfection

efficiency and survival rate [8,15,16]. Herein, we used the self-developed microextrusion-based transient-transfection system as the naked plasmid delivery device to enhance the transfection efficiency of hUC-MSCs while protecting the cells from heat shock. We verified that apoptosis was not initiated in squeezed hUC-MSCs based on the absence of spike-like membranes or vesicle morphology [43]. Moreover, nozzles with smaller inner diameters (200 and 90 mm) might have the ability to increase the squeezing force and thus induce transient permeability in the cell membranes, enhancing the naked plasmid delivery efficiency (Supplementary Figure 3). In this study, we optimized the conditions for using of a synthetic hydrogel for simultaneous transfection of cardiac marker gene into hUC-MSCs, printing tissue analogues capable of engraftment with endogenous tissue and promoting transfected hUC-MSCs differentiation into cardiomyocyte-like cells for cardiac-related applications. The surface design of gene vectors or NPs, the bioprinting ink or culture medium-containing transfection reagents and the use of cationic polymers or fibrin-hyaluronic hydrogels to further enhance plasmid delivery efficiency [8,9,13,15,16,19 21,44,45]. Positively charged materials like NPs and transfection reagents interact with negatively charged plasmids to enhance the delivery efficiency by binding to the negatively charged cell surface but frequently cause cytotoxicity [46,47]. The chitosan-based NPs that showed a slightly negative charge (<15 mV) and »150 nm size, efficiently enhanced the uptake in tumor cells [48]. In our study, naked plasmids were extruded with PU hydrogels and hUC-MSCs. We found that the PU NPs with different soft-segment content resulted in various NP sizes and charges, which might have influenced the delivery efficiency of naked plasmids (Supplementary Figure 4 and Supplementary Table II). We subsequently found that the transfection efficiency and hUCMSC viability were positively correlated with the zeta potential and particle size of the PU NPs within a range of negative charge between 60 mV and 40 mV and a particle size of »45 nm. The particle size

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and charge of the PU2 NPs were within the same range; however, the cell viability and transfection efficiency were low. The soft segment can be combined with the hard segment to affect the hardness, elastic modulus, gelation, particle size and surface charges of PU [30,49 53]. Our previous study discovered that the PU NPs might swell due to the increased mobility of molecular chains, allowing water molecules to penetrate the PU NPs as the temperature increased [32,54]. The increased NP swelling might be associated with Na+ ions in culture medium interacting with COO groups in the hard segments of PU NPs. Moreover, the surface charge of PU NPs is also associated with the dissociation of COOH groups in the hard segment [54]. We also found that PU2 might harbor additional dissociated COOH groups, resulting in an increased negative charge and PU1 exhibiting higher degrees of binding involving COO , resulting in larger particle size. The greater secondary force measured in PU1 might enhance chain-folding and swelling [55], which might enhance the ability of PU1 to more easily form a hydrogel. Previous studies reported that increase in the molecular weight of the soft segment resulted in decrease in hard segment content [56,57], leading to a decrease in hydrogel stiffness [23,24]. Moreover, PU NPs containing PDLLA1 reportedly resulted in a rod-like shape and increased elongation of the PDLLA1 PU NPs [32]. NPs with a predominant rod-like morphology were more easily packaged together than those with a sphere-like morphology [58], resulting in a more compact PU1 package that exhibited a larger elastic modulus than that of PU2. These results implied that the soft-hard segment composition of the PU1 might be more suitable than PU2 for use in the microextrusion-based transient-transfection system to enhance naked plasmid delivery efficiency and reduce cytotoxicity. Transfection through the microextrusion-based system was achieved in situ and was more convenient and faster than previous transfection methods, such as those involving transfection reagents. Moreover, we did not use any transfection reagents but rather used PU NP dispersions with negative charges and flexibility [30]. Accordingly, the PU NP dispersions formed gels at 37°C without the use of an additional crosslinker. Comparisons of the physical properties and transfection efficiencies of PU1 and PU2 revealed that rheological properties affected naked plasmid delivery efficiency. Furthermore, the expression of cardiac-associated markers of transfected hUCMSCs cultured in PU1 exceeded those of transfected cells in PU2. The stiffness of printed scaffolds influenced cell behavior, including proliferation, migration and cytoskeletal architecture, all of which could affect plasmid delivery efficiency and cell differentiation [14,29,59 61]. A previous study reported that the moduli of hydrogels fell between 12 and 24 kPa, and the stem cells differentiated into skeletal muscle cells [59]. Moreover, the stiffness of healthy heart muscle ranged from »10 to 20 kPa at the early stage of diastole and then rose to 50 kPa at the end of diastole [62]. Our results agreed with those of previous studies, showing that hydrogels with optimal moduli provided a platform for transfected hUC-MSC growth, influenced naked plasmid delivery into hUC-MSC nuclei and further enhanced transfected hUC-MSC differentiation. This might explain our observation of markedly elevated levels of cardiac-specific genes and proteins, including GATA4, MEF2C, TBX5, NKX2.5, MYH6 and TNNT2, in hUC-MSCs in PU1 hydrogels (stiffer) relative to those in PU2 hydrogels (softer). Exogenous molecules are important tools for therapeutic applications in stem cells because they can enter and affect almost all cells [63,64]. The transient perturbation of transcriptome could trigger the differentiation of MSCs [65 67]. A recent study reported that transfection of nanoparticles composed of polyethyleneimine and plasmid of the chondrogenic transcription factor into hMSCs could enhance the expression levels of chondrogenic marker gene and protein after 3 weeks [68]. In this study, we delivered the cardiac transcription factor GATA4, which can promote the binding of other cardiac transcription factors, such as MEF2C, TBX5 and MYOCD, to their respective

promoter regions, inhibiting non-cardiac gene expression in hUCMSCs [33 35]. Combinations of GATA4, MEF2C and TBX5 (GMT) genes transdifferentiate tail-tip, cardiac and neonatal fibroblasts to cardiomyocyte-like cells in vitro [69 71]. A recent study also confirmed the cardiomyocyte differentiation ability of UC-MSCs [72]. Although the exogenous GATA4 gene may be degraded by the enzymes of hUCMSCs and extruded from cells, the transient perturbation may still induce the expression of endogenous cardiac specific genes and proteins in hUC-MSCs, enhancing the cardiac differentiation of GATA4transfected hUC-MSCs. We simultaneously observed the dynamic interaction of these genes in stem cells. Our results showed that the expression of cardiac-specific genes and proteins were upregulated in GATA4-transfected hUC-MSCs. In contrast, the expression of GATA4 was quickly upregulated for GATA4-transfected human fibroblasts at day 3, which was low (»0.03) at day 15. Moreover, the expression of MEF2C and TBX5 genes was not upregulated (only »0.05) in GATA4-transfected human fibroblasts at day 3 and 15 (Supplementary Figure 5). These results suggested that the transfection was a transient effect in non stem cells and reprogramming efficiency was highly dependent on the specific genetic, epigenetic or transcriptome status of the cells. A previous study reported that the average electrical activity of cardiomyocytes induced from human embryonic stem cells was »300 mV [73]. Moreover, the amplitude of electrical activity of hMSCs (»64 mV) was lower than that of cardiomyocytes (»478 mV) [74]. In the present study, examination of the electrical activity of untransfected and GATA4-transfected hUC-MSCs in PU1 hydrogels (Supplementary Figure 6) showed higher activity associated with GATA4-transfected hUC-MSCs in PU1 hydrogels (»230 mV) relative to that in the untransfected group (»180 mV). Consistent with these findings, the hydrogel with optimal moduli might promote GATA4-transfected hUC-MSCs reprogramming into cardiomyocyte-like cells. We then used zebrafish as an in situ cardiac repair model. Zebrafish are vertebrates, making them a more suitable genetic model based on similarity to humans than Drosophila or Caenorhabditis elegans. Moreover, zebrafish have a shorter generation time (3 months) than other vertebrate research models [75]. Zebrafish heart, as mammalian heart, develops from the cardiac progenitor cells. The variations in the heartbeat of zebrafish can be related to heart defects, such as irregular arrhythmia, fibrillation, arrest of the ventricular beat, asystole or heart contractility [76 78]. The size and shape of heart are also associated with the cardiac function [79]. Trabeculae in adult zebrafish heart were reported to be the functional equivalent of the His-Purkinje fibers in the mammalian heart, which are essential for maintaining a consistent heart rhythm and increasing myocardial surface area for blood oxygenation [80,81]. We compared the recovery of heart function in zebrafish heart-defect models transplanted with untransfected and GATA4-transfected hUC-MSCs. The abnormal heart size and heartbeat rate might be associated with irregular cardiac function, and the percentage of trabeculae might be related to insufficient blood oxygen in the untransfected group. Moreover, we observed that the heartbeat rates and histological profiles of zebrafish hearts with GATA4-transfected hUC-MSCs in PU1 hydrogels were similar to those observed in the normal group. Conclusively, GATA4 was sufficient to promote stem cell differentiation into cardiomyocyte-like cells and facilitate recovery of heart function in zebrafish. In conclusion, our results demonstrated the efficacy of PU hydrogels using the microextrusion-based transient-transfection system with regard to naked GATA4 plasmid transfection efficiency, hUCMSC viability and differentiation efficiency. Additionally, the negatively charged thermo-responsive PU dispersions were easily mixed with cells and naked plasmids for in vitro and in situ tissue repair. The naked plasmid delivery efficiency in the absence of the PolyFect transfection reagent was enhanced by suitable hard-soft segment composition of PU. The plasmids and hUC-MSCs were co-extruded with the flexible PU1 hydrogel through the 200-mm nozzle without

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subsequent signs of cytotoxicity and induced differentiation of GATA4-transfected hUC-MSCs into cardiomyocyte-like cells reviving the heart function of zebrafish. Our findings suggested that the PU hydrogel microextrusion-based transient-transfection system represented a simple, reproducible, nontoxic, effective method that showed potential for use in in situ naked gene delivery and hUC-MSC reprogramming. The system also represented a promising strategy for gene therapy and in situ tissue engineering applications. Declaration of Competing Interest The authors have no commercial, proprietary, or financial interest in the products or companies described in this article. Author Contributions Conception and design of the study: Nien-Chi Huang, Chii-Ming Lee, and Shan-hui Hsu. Acquisition of data: Nien-Chi Huang. Analysis and interpretation of data: Nien-Chi Huang. Drafting or revising the manuscript: Nien-Chi Huang, Chii-Ming Lee, and Shan-hui Hsu. All authors have approved the final article. Acknowledgements This work was supported by the Program Project for Regenerative Medicine (MOST-106-3114-Y-043-021), Ministry of Science and Technology, Taiwan, R.O.C. Supplementary materials Supplementary material associated with this article can be found in the online version at doi:10.1016/j.jcyt.2019.12.003. References [1] Rolland A. Gene medicines: the end of the beginning? Adv Drug Deliv Rev 2005;57:669–73. [2] Peng CH, Cherng JY, Chiou GY, Chen YC, Chien CH, Kao CL, et al. Delivery of Oct4 and SirT1 with cationic polyurethanes-short branch PEI to aged retinal pigment epithelium. Biomaterials 2011;32:9077–88. [3] Chuah MK, Collen D, VandenDriessche T. Biosafety of adenoviral vectors. Curr Gene Ther 2003;3:527–43. [4] Boussif O, Lezoualc’h F, Zanta MA, Mergny MD, Scherman D, Demeneix B, Behr JP. A versatile vector for gene and oligonucleotide transfer into cells in culture and in vivo: polyethylenimine. PNAS 1995;92:7297–301. [5] Mintzer MA, Simanek EE. Nonviral vectors for gene delivery. Chem Rev 2009;109:259–302. [6] Kim TK, Eberwine JH. Mammalian cell transfection: the present and the future. Anal Bioanal Chem 2010;397:3173–8. [7] Mehier-Humbert S, Guy RH. Physical methods for gene transfer: improving the kinetics of gene delivery into cells. Adv Drug Deliv Rev 2005;57:733–53. [8] Ozbolat IT, Hospodiuk M. Current advances and future perspectives in extrusionbased bioprinting. Biomaterials 2016;76:321–43. [9] Chang R, Nam J, Sun W. Effects of dispensing pressure and nozzle diameter on cell survival from solid freeform fabrication based direct cell writing. Tissue Eng Part A 2008;14:41–8. [10] Zhu W, Ma X, Gou M, Mei D, Zhang K, Chen S. 3D printing of functional biomaterials for tissue engineering. Curr Opin Biotechnol 2016;40:103–12. [11] Evans CH, Ghivizzani SC, Robbins PD. Progress and prospects: genetic treatments for disorders of bones and joints. Gene Ther 2009;16:944–52. [12] Lin H, Tang Y, Tan J, Wang B, Tuan R. 474. One-step BMP-2 gene-activated 3dimensional matrix for bone formation. Mol Ther 2014;22:S182. [13] Sharei A, Zoldan J, Adamo A, Sim WY, Cho N, Jackson E, et al. A vector-free microfluidic platform for intracellular delivery. PNAS 2013;110:2082–7. [14] Hofmann M, Wollert KC, Meyer GP, Menke A, Arseniev L, Hertenstein B, et al. Monitoring of bone marrow cell homing into the infarcted human myocardium. Circulation 2005;111:2198–202. [15] Cui X, Dean D, Ruggeri ZM, Boland T. Cell damage evaluation of thermal inkjet printed Chinese hamster ovary cells. Biotechnol Bioeng 2010;106:963–9. [16] Balbino TA, Azzoni AR, de la Torre LG. Microfluidic devices for continuous production of pDNA/cationic liposome complexes for gene delivery and vaccine therapy. Colloids Surf B Biointerfaces 2013;111:203–10.

11

[17] Ma X, Qu X, Zhu W, Li YS, Yuan S, Zhang H, et al. Deterministically patterned biomimetic human iPSC-derived hepatic model via rapid 3D bioprinting. PNAS 2016;113:2206–11. [18] George MW. The origins and the future of microfluidics. Nature 2006;442:368–73. [19] Andrew JD. Control and detection of chemical reactions in microfluidic systems. Nature 2006;442:394–402. [20] Xu T, Rohozinski J, Zhao W, Moorefield EC, Atala A, Yoo JJ. Inkjet-mediated gene transfection into living cells combined with targeted delivery. Tissue Eng. Part A 2008;15:95–101. [21] Hou D, Youssef EA, Brinton TJ, Zhang P, Rogers P, Price ET, et al. Radiolabeled cell distribution after intramyocardial, intracoronary, and interstitial retrograde coronary venous delivery: implications for current clinical trials. Circulation 2005;112:150–6. [22] Kai D, Prabhakaran MP, Stahl B, Eblenkamp M, Wintermantel E, Ramakrishna S. Mechanical properties and in vitro behavior of nanofiber-hydrogel composites for tissue engineering applications. Nanotechnology 2012;23:095705. [23] Liow SS, Dou Q, Kai D, Karim AA, Zhang K, Xu F, Loh XJ. Thermogels: in situ gelling biomaterial. ACS Biomater Sci Eng 2016;2:295–316. [24] Wang LL, Burdick JA. Engineered hydorgels for local and sutained delivery of RNA-interference therapies. Adv Healthc Mater 2017;1:1601041. [25] Fedorovich NE, Alblas J, de Wijn JR, Hennink WE, Verbout AJ, Dhert WJ. Hydrogels as extracellular matrices for skeletal tissue engineering: state-of-the-art and novel application in organ printing. Tissue Eng 2007;13:1905–25. [26] Zhang SB, Xu YM, Wang B, Qiao WH, Liu DL, Li ZS. Cationic compounds used in lipoplexes and polyplexes for gene delivery. J Control Release 2004;100:165–80. [27] Pack DW, Hoffman AS, Pun S, Stayton PS. Design and development of polymers for gene delivery. Nat Rev Drug Discov 2005;4:581–93. [28] Lv H, Zhang S, Wang B, Cui S, Yan J. Toxicity of cationic lipids and cationic polymers in gene delivery. J Control Release 2006;114:100–9. [29] Place ES, George JH, Williams CK, Stevens MM. Synthetic polymer scaffolds for tissue engineering. Chem Soc Rev 2009;38:1139–51. [30] Hsu Sh, Hung KC, Lin YY, Su CH, Yeh HY, Jeng US, et al. Water-based synthesis and processing of novel biodegradable elastomers for medical applications. J Mater Chem B 2014;2:5083–92. [31] Hung KC, Tseng CS, Hsu Sh. Synthesis and 3D printing of biodegradable polyurethane elastomer by a water-based process for cartilage tissue engineering applications. Adv Healthc Mater 2014;3:1578–87. [32] Ho L, Hsu Sh. Cell reprogramming by 3D bioprinting of human fibroblasts in polyurethane hydrogel for fabrication of neural-like constructs. Acta Biomater 2018;70:57–70. [33] Lickert H, Takeuchi JK, Von Both I, Walls JR, McAuliffe F, Adamson SL, et al. Baf60c is essential for function of BAF chromatin remodelling complexes in heart development. Nature 2004;432:107–12. [34] Houweling AC, Borren MM, Moorman AFM, Christoffels VM. Expression and regulation of the atrial natriuretic factor encoding gene Nppa during development and disease. Cardiovasc Res 2005;67:583–93. [35] Garg V, Kathiriya IS, Barnes R, Schluterman MK, King IN, Butler CA, et al. GATA4 mutations cause human congenital heart defects and reveal an interaction with TBX5. Nature 2003;424:443–7. [36] Gonzalez-Rosa JM, Martin V, Peralta M, Torres M, Mercader N. Extensive scar formation and regression during heart regeneration after cryoinjury in zebrafish. Development 2011;138:1663–74. [37] Poss KD, Wilson LG, Keating MT. Heart regeneration in zebrafish. Science 2002;298:2188–90. [38] Zhu JJ, Xu YQ, He JH, Yu HP, Huang CJ, Gao JM, et al. Human cardiotoxic drugs delivered by soaking and microinjection induce cardiovascular toxicity in zebrafish. J Appl Toxicol 2014;34:139–48. [39] Gheisari Y, Soleimani M, Azadmanesh K, Zeinali S. Multipotent mesenchymal stromal cells: optimization and comparison of five cationic polymer-based gene delivery methods. Cytotherapy 2008;10:815–23. [40] Zhang K, Fan H, Wang Z, Taylor JSA, Wooley KL. Cationic shell-crosslinked knedellike nanoparticles for highly efficient gene and oligonucleotide transfection of mammalian cells. Biomaterials 2009;30:968–77. [41] Pouton CW, Seymour LW. Key issues in non-viral gene delivery. Adv Drug Deliver Rev 2001;46:187–203. [42] Park JY, Yoo SJ, Lee EJ, Lee DH, Kim JY, Lee SH. Increased poly(dimethylsiloxane) stiffness improves viability and morphology of mouse fibroblast cells. BioChip J 2010;4:230–6. [43] Sun Y, Wang W, Li B, Wu Y, Wu H, Shen W. Synchronized expression of two caspase family genes, ice-2 and ice-5, in hydrogen peroxide-induced cells of the silkworm, Bombyx mori. J Insect Sci 2010;10:1–10. [44] Lei P, Padmashali R, Andreadis ST. Cell-controlled and spatially arrayed gene delivery from fibrin hydrogels. Biomaterials 2009;30:3790–9. [45] Lei Y, Rahim M, Ng Q, Segura T. Hyaluronic acid and fibrin hydrogels with concentrated DNA/PEI polyplexes for local gene delivery. J Control Release 2011;153:255–61. [46] Wilson CJ, Clegg RE, Leavesley DI, Pearcy MJ. Mediation of biomaterial cell interactions by adsorbed proteins: a review. Tissue Eng 2005;11:1–18. [47] Nan A, Bai X, Son SJ, Lee SB, Ghandehari H. Cellular uptake and cytotoxicity of silica nanotubes. Nano Lett 2008;8:2150–4. [48] He C, Hu Y, Yin L, Tang C, Yin C. Effects of particle size and surface charge on cellular uptake and biodistribution of polymeric nanoparticles. Biomaterials 2010;31:3657–66. [49] Lin HH, Hsieh FY, Tseng CS, Hsu Sh. Preparation and characterization of biodegradable polyurethane hydrogel and the hybrid gel with soy protein for 3D cellladen bioprinting. J Mater Chem B 2016;4:6694–705.

ARTICLE IN PRESS 12

N.-C. Huang et al. / Cytotherapy 00 (2019) 1 12

[50] Wu GH, Hsu Sh. Synthesis of water-based cationic polyurethane for antibacterial and gene delivery applications. Colloids Surf B Biointerfaces 2016;146:825–32. [51] De D, Gaymans RJ. Thermoplastic polyurethanes with TDI-based monodisperse hard segments. Macromol Mater Eng 2009;294:405–13. [52] Hossieny NJ, Barzegari MR, Nofar M, Mahmood SH, Park CB. Crystallization of hard segment domains with the presence of butane for microcellular thermoplastic polyurethane foam. Polymer 2014;55:651–62. € rgens E. Recent developments in aqueous [53] Melchiors M, Sonntag M, Kobusch C, Ju two-component polyurethane (2K-PUR) coatings. Prog Org Coat 2000;40: 99–109. [54] Ou CW, Su CH, Jeng US, Hsu Sh. Characterization of biodegradable polyurethane nanoparticles and thermally induced self-assembly in water dispersion. ACS Appl Mater Interfaces 2014;6:5685–94. [55] Mondal T, Dan K, Deb J, Jana SS, Ghosh S. Hydrogen-bonding-induced chain folding and vesicular assembly of an amphiphilic polyurethane. Langmuir 2013;29:6746–53. [56] Pei A, Malho JM, Ruokolainen J, Zhou Q, Berglund LA. Strong nanocomposite reinforcement effects in polyurethane elastomer with low volume fraction of cellulose nanocrystals. Macromolecules 2011;44:4422–7. [57] Lin JR, Chen LW. Study on shape-memory behavior of polyether-based polyurethanes. II. Influence of soft-segment molecular weight. J Appl Polym Sci 1998;69:1575–86. [58] Albert PP, Anieke MW. On the density and structure formation in gels and clusters of colloidal rods and fibers. Langmuir 1998;14:49–54. [59] Peyton SR, Raub CB, Keschrumrus VP, Putnam AJ. The use of poly(ethylene glycol) hydrogels to investigate the impact of ECM chemistry and mechanics on smooth muscle cells. Biomaterials 2006;27:4881–93. [60] Kong HJ, Liu J, Riddle K, Matsumoto T, Leach K, Mooney DJ. Non-viral gene delivery regulated by stiffness of cell adhesion substrates. Nat Mate 2005;4:460–4. [61] Keeney M, Zhang Z, Yang F. Modulating polymer chemistry to enhance non-viral gene delivery inside hydrogels with tunable matrix stiffness. Biomaterials 2013;34:9657–65. [62] Chen QZ, Harding SE, Ali NN, Lyon AR, Boccaccini AR. Biomaterials in cardiac tissue engineering: ten years of research survey. Mater Sci Eng B 2008;59:1–37. [63] Lukaszewicz AI, McMillan MK, Kahn M. Small molecules and stem cells. Potency and lineage commitment: the new quest for the fountain of youth. J Med Chem 2010;53:3439–53. [64] Xu Y, Shi Y, Ding S. A chemical approach to stem-cell biology and regenerative medicine. Nature 2008;453:338–44. [65] Takahashi K, Tanabe K, Ohnuki M, Narita M, Ichisaka T, Tomoda K, Yamanaka S. Induction of pluripotent stem cells from adult human fibroblasts by defined factors. Cell 2007;131:861–72.

[66] Wernig M, Lengner CJ, Hanna J, Lodato MA, Steine E, Foreman R, et al. A druginducible transgenic system for direct reprogramming of multiple somatic cell types. Nat Biotechnol 2008;26:916–24. [67] Lia Z, Zhang C, Weiner LP, Zhang Y, Zhon JF. Molecular characterization of heterogeneous mesenchymal stem cells with single-cell transcriptomes. Biotechnol Adv 2013;31:312–7. [68] Yi SW, Kim HJ, Oh HJ, Shin H, Lee JS, Park JS, Park KH. Gene expression profiling of chondrogenic differentiation by dexamethasone-conjugated polyethyleneimine with SOX trio genes in stem cells. Stem Cell Res Ther 2018;9:341. [69] Ieda M, Fu JD, Delgado-Olguin P, Vedantham V, Hayashi Y, Bruneau BG, Srivastava D. Direct reprogramming of fibroblasts into functional cardiomyocytes by defined factors. Cell 2010;142:375–86. [70] Qian L, Huang Y, Spencer CI, Foley A, Vedantham V, Liu L, et al. In vivo reprogramming of murine cardiac fibroblasts into induced cardiomyocytes. Nature 2012;485:593–8. [71] Song K, Nam YJ, Luo X, Qi X, Tan W, Huang GN, et al. Heart repair by reprogramming non-myocytes with cardiac transcription factors. Nature 2012;485:599–604. [72] Iachininoto MG, Capodimonti S, Podda MV, Valentini CG, Bianchi M, Leone AM, et al. In vitro cardiomyocyte differentiation of umbilical cord blood cells: crucial role for c-kit(+) cells. Cytotherapy 2015;17:1627–37. [73] Anderson D, Self T, Mellor IR, Goh G, Hill SJ, Denning C. Transgenic enrichment of cardiomyocytes from human embryonic stem cells. Mol Ther 2007;15:2027–36. [74] Beeres SLMA, Atsma DE, Laarse A, Pijnappels DA, Tuyn J, Fibbe WE, et al. Human adult bone marrow mesenchymal stem cells repair experimental conduction block in rat cardiomyocyte cultures. J Am Coll Cardiol 2005;46:1943–52. [75] Flinn L, Bretaud S, Lo C, Ingham PW, Bandmann O. Zebrafish as a new animal model for movement disorders. J Neurochem 2008;106:1991–7. [76] Luca ED, Zaccaria GM, Hadhoud M, Rizzo G, Ponzini R, Morbiducci U, Santoro MM. ZebraBeat: a flexible platform for the analysis of the cardiac rate in zebrafish embryos. Sci. Rep. 2014;4:4898. [77] Li K, Wu JQ, Jiang LL, Shen LZ, Li J-Y, He ZH, et al. Developmental toxicity of 2,4dichlorophenoxyacetic acid in zebrafishembryos. Chemosphere 2017;171:40–8. [78] Langheinrich U, Vacun G, Wagner T. Zebrafish embryos express an orthologue of HERG and are sensitive toward a range of QT-prolonging drugs inducing severe arrhythmia. Toxicol Appl Pharm 2003;193:370–82. [79] Hu N, Yost HJ, Clark EB. Cardiac morphology and blood pressure in the adult zebrafish. Anat Rec 2011;264:1–12. [80] Sedmera D, Reckova M, deAlmeida A, Sedmerova M, Biermann M, Volejnik J, et al. Functional and morphological evidence for a ventricular conduction system in zebrafish and Xenopus hearts. Am J Physiol Heart Circ Physiol 2003;284:H1152–60. [81] Samsa LA, Yang B, Liu J. Embryonic cardiac chamber maturation: Trabeculation, conduction, and cardiomyocyte proliferation. Am J Med Genet C 2013;163C:157–68.