Fabrication of Photo-Crosslinkable Poly(Trimethylene Carbonate)/Polycaprolactone Nanofibrous Scaffolds for Tendon Regeneration.

BACKGROUND: The treatment of tendon injuries remains a challenging problem in clinical due to their slow and insufficient natural healing process. Scaffold-based tissue engineering provides a promising strategy to facilitate tendon healing and regeneration. However, many tissue engineering scaffolds have failed due to their poor and unstable mechanical properties. To address this, we fabricated nanofibrous polycaprolactone/methacrylated poly(trimethylene carbonate) (PCL/PTMC-MA) composite scaffolds via electrospinning.
MATERIALS AND METHODS: PTMC-MA was characterized by nuclear magnetic resonance. Fiber morphology of composite scaffolds was evaluated using scanning electron microscopy. The monotonic tensile test was performed for determining the mechanical properties of composite scaffolds. Cell viability and collagen deposition were assessed via PrestoBlue assay and enzyme-linked immunosorbent assay, respectively.
RESULTS: These PCL/PTMC-MA composite scaffolds had an increase in mechanical properties as PTMC-MA content increase. After photo-crosslinking, they showed further enhanced mechanical properties including creep resistance, which was superior to pure PCL scaffolds. It is worth noting that photo-crosslinked PCL/PTMC-MA (1:3) composite scaffolds had a Young's modulus of 31.13 ± 1.30 MPa and Max stress at break of 23.80 ± 3.44 MPa that were comparable with the mechanical properties of native tendon (Young's modulus 20-1200 MPa, max stress at break 5-100 MPa). In addition, biological experiments demonstrated that PCL/PTMC-MA composite scaffolds were biocompatible for cell adhesion, proliferation, and differentiation.

Introduction

Tendon injuries are frequently-occurring diseases in our daily life. It was estimated that over 25% of adults have suffered a tendon-related musculoskeletal problem.1 In case of severe tendon injuries, their natural healing processes are slow and insufficient due to the low innervation, hypovascularity, and hypocellularity of these tissues.2,3 Surgical intervention is the main reconstruction strategy, including autografts, allografts, and xenografts.4,5 However, the clinical outcomes of tendon repair remain unsatisfactory as a result of the limit of grafts’ donor source, the existence of immune rejection and inflammation.6

Tissue engineering, also known as“regenerative medicine”, is an interdisciplinary field which combines life sciences, materials science, and engineering towards the development of biological substitutes that restore, maintain, or replace diseased tissues.7 Scaffold-based tissue engineering which is the use of a combination of cells, scaffolds and growth factors to develop functional replacements for traditional grafts, shows great potential for application in the clinical treatment of tendon injuries.6,8,9 Currently, techniques for fabricating tissue engineering scaffolds mainly include 3D printing,10 knitting machine,11 and electrospinning.12 Electrospinning is a versatile and simple approach for generating nanofibrous scaffolds that mimic the structure of native extracellular matrix.13,14 As such, electrospinning has been widely used for fabricating tissue engineering scaffolds.

How to improve mechanical properties of electrospun scaffolds is one of the most important aspects in tissue engineering.1,6,15,16 In general, there are three methods to improve the mechanical properties of electrospun scaffolds including design specific structure for scaffolds, composite scaffolds with other materials and photo-crosslinking. The process of constructing specific structure is complicated.17 While it is simple and common to improve the mechanical properties of polymer scaffolds by physically mixing them with other materials to form composite scaffolds and via photo-crosslinking.18 Forming a polymer network by chemically crosslinked polymers is beneficial to improve creep resistance of scaffolds.19 For example, Chen et al used a photo-crosslinkable polymer, polylactic acid-co-acrylate, to prepare an electrospun-oriented fiber scaffold. Cyclic tensile test results showed that photo-crosslinked scaffold has good creep resistance.20

Polycaprolactone (PCL) is an aliphatic linear polyester which has been approved by the FDA for clinical application due to its biocompatibility, bioabsorbability and low cost.21 However, PCL exhibits a low degradation rate due to its semi-crystalline and hydrophobic nature resulting in a relatively slow healing process of injured tendons.22–24 Besides, poor mechanical tolerance of PCL is another drawback for its further application in clinical.25

Poly(trimethylene carbonate) (PTMC), another FDA approved material, is a rubber-like polymer with a glass transition temperature (Tg) around −20°C.19,21 PTMC has unique degradation behaviors such as enzymatic degradation with a surface erosion process, without formation of acidic products, and resistance to non-enzymatic hydrolysis.26,27 Therefore, this elastomeric polymer has been exploited in regenerative medicine, including vascular tissue engineering,28,29 bone defect repair,30,31 and drug loading implantation.32,33 To the best of our knowledge, the use of PTMC in tendon regeneration have not been investigated. It is worth noting that PTMC is soft with a poor rigidity.23,28 Thus, fabrication of PTMC nanofibrous scaffolds by electrospinning remains a challenge since PTMC fibers will collapse and fuse together.33–35

In the present work, photo-crosslinkable PTMC (PTMC-MA) macromers were prepared by ring-opening polymerization of 1,3-Dioxane-2-one (TMC) monomers and subsequently functionalized with methacrylate. PTMC-based nanofibrous scaffolds were prepared by electrospinning of mixed solutions of PCL and PTMC-MA. Our hypothesis was that the introduction of PCL would facilitate PTMC-MA electrospinning. Furthermore, it was expected that the use of photo-crosslinkable PTMC-MA instead of PTMC would enhance the creep resistance of composite scaffolds. The properties of PCL/PTMC-MA composite scaffolds, including fiber morphology, fiber orientation, and mechanical properties were investigated. To determine the cellular compatibility of PCL/PTMC-MA composite scaffolds for their potential applications in tendon tissue engineering, mouse mesenchymal stem cells were cultured on the scaffolds, and their cell proliferation and differentiation were evaluated later on.

Materials and Methods

Materials

1,3-Dioxane-2-one (trimethylene carbonate, TMC) was obtained from Jinan Daigang Biomaterial Co., Ltd. (Jinan, China). Polycaprolactone (PCL, Mn≈80 kg/mol), 1,1,1-tris(hydroxymethyl) propane (trimethylolpropane, TMP), Tin (II) ethylhexanoate (Sn(Oct)2), methacrylic anhydride (94%, MMAh), triethylamine (TEA), ascorbic acid (AA) and deuterated chloroform were all purchased from Sigma-Aldrich (MO, USA). Calcium hydride and Irgacure 2959 (2-hydroxy-4-(2-hydroxyethoxy)-2-methylpropiophenone, D2959) were purchased from Aladdin Inc. (Shanghai, China). Dichloromethane (DCM), methanol, and dimethylformamide (DMF) were received from Tianjin Zhiyuan Chemical Reagent Factory (Tianjin, China). DCM was dried by calcium hydride and distilled under N2 atmosphere before use. Hydroquinone was purchased from Tianjin Damao Chemical Reagent Factory (Tianjin, China). High-glucose Dulbecco’s Modified Eagle’s Medium (DMEM) and Fetal bovine serum (FBS) were supplied from Biological Industries (Israel), Trypsin and Penicillin-Streptomycin-Gentamicin Solution were purchased from Solarbio Inc. (Beijing, China). PrestoBlue™ HS Cell Viability Reagent was purchased from Thermo Fisher Scientific (Massachusetts, USA). Mouse collagen type I ELISA kit was purchased from Qchen Bio Co., Ltd. (Shanghai, China). All other reagents and organic solvents are of analytic reagent (AR) without further treatment prior to use.

Synthesis of Methacrylated Poly(Trimethylene Carbonate)

Three-armed PTMC was prepared by ring-opening polymerization (ROP) of TMC initiated by TMP as previously described.36 PTMC’s molecular weight was controlled by varying the amount of TMP. Briefly, TMC/TMP, in a molar ratio of 500/1, was loaded in a dry eggplant-shaped flask under dry N2 atmosphere. The polymerization was carried out at 130°C for 3 days using 0.13 wt% of Sn(Oct)2 as a catalyst. The resulting polymer was purified by dissolving in dry DCM and followed by precipitating in cold methanol for 3 times. The precipitate was dried at room temperature in vacuo. In order to obtain a photo-crosslinkable compound, PTMC-MA was subsequently synthesized using an established protocol with slight modification.37 Three-armed PTMC was functionalized in DCM (5 mL/g, PTMC) by reaction with MAAh (4 mol/mol) in the presence of TEA (4 mol/mol) and hydroquinone (0.03 mol/mol). The functionalized reaction was conducted at room temperature for 5 days avoiding light. Reacted PTMC-MA solution was extracted with demineralized water 3 times and the organic phase was precipitated in cold methanol. The precipitate was firstly dried in the dark at ambient conditions overnight and then further dried in vacuo at room temperature until its weight kept constant.

Both PTMC and PTMC-MA were characterized by nuclear magnetic resonance (NMR) spectroscopy (Mercury-plus 300, Varian, USA) and gel permeation chromatography (GPC, Waters Breeze, Waters, USA; eluent: tetrahydrofuran; flow rate: 1.0 mL per minute; column calibration: polystyrene standards). The degree of functionalization of PTMC-MA was determined by 1H NMR, and deuterated chloroform was used as a solvent. The average-number molecular weight and molar-mass dispersity were measured using GPC.

Fabrication of Photo-Crosslinked Scaffolds with Aligned Orientation

Polymer solution (10%, w/v) was prepared by dissolving PCL and PTMC-MA in DCM/DMF (v/v = 3:1) with different weight ratios of PTMC-MA to PCL. The resultant mixtures were stirred for 12 h to obtain homogenous solutions before electrospinning. For electrospinning, polymer solution was extruded through a 20-gauge blunt-tip needle at a flow rate of 5 mL/h with a distance of 15 cm between spinneret and mandrel. The spinneret was charged with a voltage of 25 kV, and aligned nanofibers were collected using a mandrel at speed of 2000 rpm. Electrospinning was performed at ambient condition (room humidity Hroom = 30–50% and room temperature Troom = 23–28°C).

To obtain photo-crosslinked nanofibers, electrospun scaffolds were immersed in D2959 aqueous solution (1 wt%) for 2 h, followed by exposure to UV light source (0.072 W*cm−2) for 10 min each side. After crosslinking, electrospun scaffolds were gently washed with ethanol and PBS. The final electrospun scaffolds were dried under vacuum at room temperature (Troom = 23–28°C) for 12 h.

Characterization of PCL/PTMC-MA Photo-Crosslinked Scaffolds

The morphology of the scaffold was characterized using scanning electron microscope (SEM, Phenom G2 Pro, Phenom-World, The Netherlands) at 5 kV. Samples were fixed at the stage with electroconductive paste and sputter coated with platinum (≈2 nm, 120 s, 20 mA) before SEM imaging. The average fiber diameter and orientation of scaffold were determined by analyzing at least 500 fibers (n≈120, 5 images per scaffold) from SEM images using ImageJ software.

The monotonic tensile test was performed for determining the mechanical properties of before photocrosslinking and after crosslinking scaffolds. Briefly, each scaffold was cut into 15 mm × 50 mm strips with the same thickness (50~70 μm) before testing. All strips were strained at a rate of 20 mm/min using a pneumatically powered tensile tester (Instron 5697, USA). Moreover, to further confirm the influence of photocrosslinking treatment on mechanical behavior, strips with 50% content PTMC-MA (PCL/PTMC-MA (1:1)) were pulled to 20% strain for 20 cycles, and the change of gauge length was recorded every cycle. All tests were performed in the same external conditions (Troom = 20–30°C and Hroom = 50–60%, n = 3).

Cell Culture and Seeding on Scaffolds

Green fluorescent protein (GFP)+ C57BL/6 mouse mesenchymal stem cells (MSCs) were obtained from Cyagen Biosciences (Guangzhou, China). MSCs were expanded in T-25 flask, and cultured in a medium, comprised of high-glucose DMEM, 10% (v/v) FBS, 1% (v/v) penicillin-streptomycin-nystatin and 0.01% (v/v) ascorbic acid. Cells were harvested at approximately 80% confluence for further scaffold-based study.

Photo-crosslinked electrospun scaffolds were punched into discs (diameter = 1 cm) and placed inside a 48-well plate (n = 3, for each group). For sterilization, the scaffolds were immersed in 75% ethanol for 45 min for 3 times, followed by washing with PBS 3 times for 15min each to remove remaining ethanol. After that, sterile scaffolds were immersed in the cell culture medium overnight before cell seeding. After removing the medium from scaffolds, MSCs were seeded on each scaffold at a density of 15,000 cells/well in 100 μL cell culture media. Cell-scaffold constructs were incubated for 2 h to allow cell attachment and topped up to 0.5 mL cell culture media. The cell-scaffold constructs were cultured in an incubator with 5% CO2 humid atmosphere at 37°C for up to 21 days. Cell culture media was refreshed every two days.

Cell Distribution and Viability

The distribution of MSCs on scaffolds was analyzed using an Inverted fluorescence microscope (AX10, ZEISS, German). The viability of cells seeded on scaffolds was analyzed by PrestoBlueTM assay according to the vendor’s protocol. Briefly, cell culture media was replaced with a medium containing 10% (v/v) PrestoBlueTM HS reagent, and samples were wrapped with aluminum foil to avoid light, followed by incubating at 37°C for 45 min. Two blank wells added with 500 μL Presto BlueTM HS reagent worked as a control. Fluorescence intensity was measured in a microplate reader (Biotek Cytation 5, USA; Excitation 560 nm/Emission 590 nm). Cell viability was measured at days 1, 4, and 7.

Enzyme-Linked Immunosorbent Assay (ELISA)

The concentration of collagen type I in cell culture medium produced from MSCs was quantified by ELISA. Briefly, the same volume of cell culture medium was collected at time points of days 1, 4, 7, 14 and 21, and preserved at −80°C before further analysis. For collagen type I quantification, samples were assessed using a collagen type I Elisa kit following the vendor’s protocol. The absorbance was recorded at 450 nm in a microplate reader (Biotek Cytation 5, USA).

Statistical Analysis

The statistical analysis of data was performed using the GraphPad PRISM version 7.00. Two-way analysis of variance (ANOVA) followed by the Tukey post hoc test for multiple comparisons was performed, unless specified otherwise. The results were presented as mean ± standard deviation, and p < 0.05 meant statistical significance.

Results and Discussion

Synthesis of PTMC-MA

In the present work, three-armed PTMC was prepared by ROP of TMC using TMP as an initiator. The product was subsequently functionalized with methacrylic anhydride to yield photo-crosslinked PTMC-MA (Figure 1A). The obtained PTMC macromers had a relatively low molecular weight (Mn = 21.90 kg/mol or MW = 34.30 kg/mol) and moderate molar-mass dispersity (PDI = 1.56) measured by GPC (). A relatively low molecular weight would benefit the increase of scaffold’s mechanical properties after crosslinking, since the number of methyl propylene functional groups depend on the number of polymers.19 By comparing the 1H NMR spectra of PTMC-MA with PTMC, the appearance of the double bond proton peaks at 5.57 ppm and 6.11 ppm in PTMC-MA clearly confirmed the end-group functionalization of PTMC with methacrylic anhydride (Figure 1B). The degree of functionalization was determined following a similar method as reported by Geven et al.36 An average degree of functionalization of approximately 75% could be calculated. To evaluate the crosslink density of PTMC-MA network, gel content was measured at different time points of UV exposure (). The results demonstrated that gel content of PTMC-MA was the highest up to 99.37% after exposure to UV for 10 min. Gel content slightly decreased as the exposure time increase to 15min, which may be attributed to the degradation of network.38 Thus, the UV exposure time for PCL/PTMC-MA composite scaffolds was set at 10 min.

Figure 1

(A) The schematic synthesis of the PTMC and its subsequent functionalization using methacrylic anhydride for PTMC-MA. (B) Representative 1H NMR spectrum of PTMC and PTMC-MA.

Fabrication of PCL/PTMC-MA Photo-Crosslinked Scaffolds

Within the native tendon, spindle-shaped tenocytes are organized in linear arrays aligned with and interspersed between nano collagen fibers.2 Recent studies have demonstrated that there was a strong relationship between scaffold topography and cell behaviors.39 For example, the orientation of electrospun fibers has a close connection to the differentiation, elongation, and migration of seed cells.24,40 Hence, a tissue engineering scaffold should be ideally designed to imitate the microenvironment of ECM, which is beneficial for the treatment of tendon injury. Here, we used electrospinning, a versatile and simple approach, to generate aligned nanofibrous composite scaffolds.

SEM micrographs of composite nanofibers obtained at different mass ratios of PCL to PTMA-MA are shown in Figure 2. The fiber morphology of composite nanofibers was strongly affected by the portion of PTMC-MA. When the portion of PTMC-MA increased from 0% to 50% (pure PCL, PCL/PTMC-MA (3:1), and PCL/PTMC-MA (1:1)), a homogeneous and ultrafine fiber morphology was observed (Figure 2A–C). As the weight ratio reached up to 75% (PCL/PTMC-MA (1:3)), the fibers presented a molten morphology with an increase in fiber interconnection (Figure 2D). No fibrous structure was observed in the fabrication of pure PTMC-MA due to its natural amorphism and poor stiffness resulting in fiber fusing together, especially at relatively lower molecular weight (Figure 2E).29,33 No significant difference in the morphology of scaffolds before and after immersion in PBS (), demonstrating the successful crosslinking of PCL/PTMC-MA composite nanofibers. Figure 2F–H shows the average fiber diameters and diameter distributions of electrospun mats. A similar average fiber diameter (around 600 nm) was detected in spite of an increase of PTMC-MA content from 0% to 50%. The dimension of composite fibers obtained in the present work was comparable to collagen fibers in native tendons.41 Fiber orientations were analyzed by the statistics of fiber angles in their SEM micrographs (Figure 2I–K). Most PCL fibers were oriented with an angle distribution from −30° to 50°. PCL/PTMC-MA (3:1) fibers (−50° to 50°) had a nearly similar angle distribution range to PCL/PTMC-MA (1:1) fibers (−40° to 50°). However, the latter one presented a more equal distribution at all angles indicating more randomly oriented fibers. Taken together, the increase of three-armed PTMC-MA content would sacrifice the fiber orientation of composite fibers.

Figure 2

SEM images of electrospun fiber of pure PCL (A), PCL/PTMC-MA (3:1) (B), PCL/PTMC-MA (1:1) (C), PCL/PTMC-MA (1:3) (D) and PTMC-MA (E). Fiber diameter distributions and fiber orientations of PCL fibers (F and I), PCL/PTMC-MA (3:1) (G and J) and PCL/PTMC-MA (1:1) (H and K). Scale bar: 10 μm.

Mechanical Properties of PCL/PTMC-MA Photo-Crosslinked Scaffolds

An ideal scaffold should meet a series of mechanical properties for maintaining the basic function after implanting to the body.42 Mechanical properties of PCL/PTMC-MA composite scaffolds before and after crosslinking were determined by monotonic test (Figure 3). Young’s modulus, yield strain, and max stress at break for pure PCL scaffolds were 11.16 ± 0.55 MPa, 77 ± 10%, and 4.96 ± 0.94 MPa, respectively, and these values incrementally increased as the three-armed PTMC-MA content increased. 15.26 ± 2.37 MPa, 126 ± 22%, and 8.35 ± 1.24 MPa for PCL/PTMC-MA(3:1) before crosslinking; 17.00 ± 1.22 MPa, 149 ± 12%, and 10.31 ± 1.82 MPa for PCL/PTMC-MA(1:1) before crosslinking; 19 ± 3.42 MPa, 159 ± 9%, and 13.21 ± 0.47 MPa for PCL/PTMC-MA(1:3) before crosslinking. Previous study has demonstrated that PCL/PTMC composite scaffolds had superior elasticity and stiffness compared with pure PCL scaffolds due to combination of the advantages of PCL and PTMC.21 The mechanical properties of PCL/PTMC-MA composite scaffolds further enhanced after UV crosslinking, which was reasoned that methyl allyls of functional polyester in UV crosslinked with each other and formed a polymer network.19 As expected, PCL/PTMC-MA(1:3) composite scaffolds that had the highest PTMC-MA content showed the highest Young’s modulus (31.13 ± 1.30 MPa) and max stress at break (23.80 ± 3.44 MPa) after photo-crosslinking. However, the yield strain of photo-crosslinked PCL/PTMC-MA(1:3) scaffolds (170 ± 22%) was lower than that of photo-crosslinked PCL/PTMC-MA(1:1) (190 ± 22%), which may be due to fibers’ coagulating.34

Figure 3

Comparison of tensile mechanical properties of electrospun scaffolds with different PCL/PTMC-MA mass ratios before and after crosslinking. (A) Representative stress–strain curves, CK stands for crosslinked. (B) Max stress at break. (C) Young’s modulus. (D) Yield strain at break. *p< 0.05, **p< 0.01, ***p< 0.001, and ****p< 0.0001, *p < 0.05 stand for statistical significance.

To investigate the influence of photo-crosslinking on mechanical properties of composite scaffolds, pull-relaxation cycles monotonic test was conducted to evaluate the creep resistance of scaffolds (Figure 4). Creep resistance is closely related to scaffold stability and plays a key role during tissue regeneration.6 Each gauge length of the sample was recorded before and after one pull-relaxation monotonic test (Figure 4A). A continuous increase of sample gauge length for non-crosslinked scaffolds was observed during pull-relaxation cycles monotonic tests, and the scaffolds ruptured at the eleventh cycle during the pull-relaxation of non-crosslinked scaffolds. However, the gauge length for crosslinked scaffolds remained almost constant. Note that the Young’s modulus for non-crosslinked scaffolds dramatically increased as the pull sequence cycles increase, which may be caused by the increasing of PCL crystallinity during pulling.43 In contrast, no obvious difference in Young’s modulus for crosslinked scaffolds was observed during pull-relaxation cycles monotonic tests (Figure 4B), indicating more stable scaffolds after photo-crosslinking. Previous studies demonstrated that photo-crosslinked PTMC-MA showed enhanced stability and creep resistance compared with PTMC.19,28

Figure 4

Fatigue test on PCL/PTMC-MA (1:1) before and after crosslinking under 20 times of pull-relaxation cycles. (A) Change of sample gauge length, and (B) Change of Young’s modulus. Scaffolds of PCL/PTMC-MA (1:1) before crosslinking ruptured at the 11th stretch.

Several studies on enhancing mechanical properties of PCL scaffolds have been reported.44–47 For example, Cheol et al introduced human serum albumin (HSA) to PCL to prepare PCL/HSA composite scaffolds.48 The mechanical properties of PCL/HSA (Young’s modulus 14.39±1.30 MPa; Max stress at break 22.47 ± 3.32 MPa) were higher than those of pure PCL scaffolds (Young’s modulus 8.71 ± 0.51 MPa; Max stress at break 5.02 ± 0.41 MPa). Robert et al designed aligned PCL nanofibers to enhance the mechanical properties of PCL scaffold by electrospinning (Young’s modulus 9.90 ± 0.87 MPa in PCL aligned compared to 4.02 ± 0.88 MPa in PCL non-aligned).49 Both mixing with other materials and designing special structures could improve the mechanical properties of PCL scaffolds; nevertheless, the creep resistance of these scaffolds remains no change that limits their applications in tendon tissue engineering. The mechanical properties of our photo-crosslinked PCL/PTMC-MA(1:3) scaffolds show superior mechanical properties to previous reported studies and possess good creep resistance (Table 1). Taken together, photo-crosslinked PCL/PTMC-MA composite scaffolds have comparable mechanical properties to those of native tendon tissue (Young’s modulus 20–1200 MPa, max stress at break 5–100 MPa).50

Table 1

Comparison of Mechanical Properties of PCL/PTMC-MA Scaffolds with Reported PCL Scaffoldsa

	Young’s Modulus/MPa	Max Stress at Break/MPa	Reference
Aligned PCL	9.90 ± 0.87	5.20 ± 0.80	[49]
PCL/NaHCO3	32.0 ± 0.59	5.1 ± 2.28	[45]
PCL/HSAc	14.39 ± 1.30	22.47 ± 3.32	[48]
PLGA/PCL	7.32 ± 0.21	3.58 ± 0.09	[46]
PCL/NCb	32.9 ± 4.4	6.03 ± 0.65	[44]
PCL/Gel-MA	10 ± 1.20	3.80 ± 0.20	[47]
PCL/PTMC-MA	31.13 ± 1.30	23.80 ± 3.44	This paper

Notes: Molecular weight of all PCL is 80 kg/mol. bNC represents nanofibrillated chitosan. cHSA represents human serum albumin.

Cell Adhesion and Cell Viability

Besides with adequate mechanical properties, an ideal tendon scaffold should be biocompatible after implantation.4,42 Figure 5 shows fluorescent micrographs of GFP+ MSCs on PCL/PTMC-MA composite scaffolds with different PTMC-MA content at days 1, 4, and 7 after cell seeding. Generally, cell numbers increased steadily on all scaffolds studied during the time course investigated. Thus, PCL/PTMC-MA composite scaffolds support MSCs adhesion. Previous studies have demonstrated that cells would grow along the aligned nanofibers.39,40 In our work, however, due to the properties of PTMC which is soft and elastic, it is difficult to obtain highly aligned PCL/PTMC-MA fibers via a rotating mandrel.

Figure 5

Representative fluorescent micrographs of MSCs density on electrospun scaffolds and tissue culture plates (control) at days 1, 4, and 7. Scale bar: 50 μm.

Cell viability is vital for cell growth and proliferation.51 Cell viability of MSCs on scaffolds was investigated by PrestoBlueTM HS reagent reducing array (Figure 6). Cell viability dramatically increased on all scaffolds from day 1 to day 4 due to the proliferation of MSCs, and it became stable afterward as a result of confluent cell. At day 4, cell viability of MSCs on all electrospun scaffolds was higher than that of tissue culture plate (control), which indicated that electrospun scaffolds were biocompatible and supported MSCs proliferation. PCL/PTMC-MA(3:1) had significantly higher cell viability than PCL/PTMC-MA(1:3) at both day 4 and day 7, which may ascribe to an ultrafine fibrous morphology of PCL/PTMC-MA(3:1) scaffolds leading to a better cell proliferation. For PCL/PTMC-MA(1:3), it is significantly higher than the control plate at day 4. While at day 7, PCL/PTMC-MA(1:3) is lower than the control plate due to cell confluence.

Figure 6

Viability of MSCs on PCL scaffolds and PCL/PTMC-MA composite scaffolds. Tissue culture plates were used as a control. *p< 0.05, ***p< 0.001, and ****p< 0.0001, *p < 0.05 stand for statistical significance.

Quantification of Collagen

Type I collagen is the major component of tendon’s extracellular matrix (ECM) which is closely related to the osteogenic differentiation of MSCs and important for healing tendon injury.2,52 To investigate cell differentiation of MSCs on scaffolds, the amount of collagen I was determined by Collagen I ELISA kit (Figure 7). The amount of collagen I on all substrates dramatically increased during the first week and became relatively constant afterward. No significant difference in collagen secretion was observed between composite scaffolds regardless of PTMC-MA content. At days 14 and 21, the amount of collagen I was significantly higher for composite scaffolds than tissue culture plates (control).

Figure 7

Quantification of collagen I deposited by MSCs on PCL scaffolds and PCL/PTMC-MA composite scaffolds. Tissue culture plates were used as a control. *p< 0.05, ***p< 0.001, and ****p< 0.0001, *p < 0.05 stand for statistical significance.

Conclusions

In the present study, photo-crosslinked PCL/PTMC-MA composite scaffolds were successfully fabricated by electrospinning. The fiber morphology and mechanical properties of PCL/PTMC-MA composite scaffolds were affected by PTMC-MA content. Due to integration of PTMC and formation of crosslinked network, photo-crosslinked PCL/PTMC-MA composite scaffolds showed superior mechanical properties compared to pure PCL scaffolds, including creep resistance, Young’s modulus, and max stress at break. Photo-crosslinked PCL/PTMC-MA (1:3) composite scaffolds which had a Young’s modulus of 31.13 ± 1.30 MPa and max stress at break of 23.80 ± 3.44 MPa showed comparable mechanical properties with native tendon (Young’s modulus 20–1200 MPa, max stress at break 5–100 MPa). Cellular evaluation results demonstrated that PCL/PTMC-MA composite scaffolds were biocompatible for cell adhesion, proliferation, and differentiation. Thus, PCL/PTMC-MA composite scaffold is a promising material for tendon tissue regeneration.