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. Author manuscript; available in PMC: 2023 Oct 7.
Published in final edited form as: Biomed Mater. 2022 Oct 7;17(6):10.1088/1748-605X/ac94ae. doi: 10.1088/1748-605X/ac94ae

Anisotropicity and flexibility in trilayered microfibrous substrates promote heart valve leaflet tissue engineering

Yuriy Snyder 1, Soumen Jana 1,*
PMCID: PMC9629372  NIHMSID: NIHMS1841312  PMID: 36150373

Abstract

Heart valve leaflet substrates with native trilayer and anisotropic structures are crucial for successful heart valve tissue engineering. In this study, we used the electrospinning technique to produce trilayer microfibrous leaflet substrates using two biocompatible and biodegradable polymers - poly (L-lactic acid) (PLLA) and polycaprolactone (PCL), separately. Different polymer concentrations for each layer were applied to bring a high degree of mechanical and structural anisotropy to the substrates. PCL leaflet substrates exhibited lower unidirectional tensile properties than PLLA leaflet substrates. However, the PLLA substrates exhibited a lower flexural modulus than the PCL substrates. These substrates were seeded with porcine valvular interstitial cells (PVICs) and cultured for one month in static conditions. Both substrates exhibited cellular adhesion and proliferation, resulting in the production of tissue-engineered constructs. The PLLA tissue-engineered constructs had more cellular growth than the PCL tissue-engineered constructs. The PLLA substrates showed higher hydrophilicity, lower crystallinity, and more significant anisotropy than PCL substrates, which may have enhanced their interactions with PVICs. Analysis of gene expression showed higher α-SMA and collagen type 1 expression in PLLA tissue-engineered constructs than in PCL tissue-engineered constructs. The differences in anisotropic and flexural properties may have accounted for the different cellular behaviors in these two individual polymer substrates.

Keywords: Heart valve leaflet substrate, tissue engineering, polycaprolactone, poly-L-lactic acid, trilayer structure, anisotropy, flexural

1. Introduction

Heart valve tissue engineering offers a potential alternative solution to problematic mechanical and bioprosthetic valves by producing three-dimensional substrates using decellularized tissues or biocompatible polymers to support cell growth and neotissue formation [1]. Synthetic polymeric substrates used for tissue engineering are made from biodegradable polymers and possess tunable mechanical and structural properties without the immunological risks associated with decellularized xenogeneic tissues or the scarcity of allogeneic tissues [2].

Various biodegradable biomaterials, such as polyglycolide (PGA), poly(lactic-co-glycolic acid) (PLGA), polyglycerol sebacate (PGS), polyurethane and polycarbonate (PU/PC), are used to fabricate fibrous substrates for heart valve tissue engineering [3, 4]. However, as heart valve substrates, these materials have limitations including rapid biodegradation [5, 6], the requirement for additional carrier polymer or crosslinking support [7], and relatively low mechanical properties [8–10]. Due to the rapid biodegradation of PGA or PLGA scaffolds, investigators first developed in vitro tissue-engineered heart valves with those scaffolds before implanting them in a sheep model to find their efficacies [11]. PU/PC copolymer scaffolds were developed for heart valve tissue engineering for their elasticity and biocompatibility provided by the PU and PC, respectively [12]. However, these copolymers needed hydrolysable groups added to their polymer chains to be biodegradable. These modifications can make electrospun PU/PC substrates unstable at physiological temperatures and requires additional processing, such as crosslinking or carrier polymers [13]. Alternatively, polycaprolactone (PCL) and poly-l-lactic acid (PLLA) were added to PGS and PU, respectively, to improve the stability and mechanical properties of the latter intended for heart valve tissue engineering without requiring additional processing of the substrates [14, 15].

The heart valve leaflets are comprised of three interconnected layers – the circumferentially oriented zona fibrosa, randomly oriented zona spongiosa, and radially oriented ventricularis - that contain valvular interstitial cells (VICs) [1, 16]. These layers dictate the mechanics and anisotropic properties of the leaflet tissue. VICs are the primary cell type residing within leaflets and are responsible for valvular homeostasis, ECM material production, and leaflet repair [17].

Most of the synthetic leaflet substrates investigated so far for heart valve tissue engineering had non-native isotropic microstructure and mechanical properties (e.g., solid-porous, randomly oriented fibrous, and hydrogel substrates) [8, 14, 18]. These non-native properties may trigger pathological phenotype in VICs and cause leaflet retraction and calcification, leading to valve failure [11, 19]. Due to these non-native mechanical cues in the leaflet substrates made of the above-mentioned biomaterials, leaflet retraction was observed in their in vitro and in vivo studies [8, 11, 20, 21]. Multiple cell and computational studies have suggested that VICs exhibit less pathogenic cell activation and more expression of quiescent phenotypes in substrates with anisotropic properties [22, 23]. Therefore, developing substrates with anisotropic properties could be essential to successful heart valve tissue engineering.

PCL and PLLA – two biodegradable and FDA complaint biomaterials - possess mechanical properties that can bear the hemodynamic load and have degradation times of 6–12 months, which makes them suitable for heart valve substrates [24]. Further, PCL and PLLA can be used to produce highly anisotropic electrospun substrates because they have a large range of mechanical properties [25]. Our previous work used trilayer leaflet PCL substrates for subcutaneous tissue engineering in the rat model, and the developed tissue constructs had a distinct trilayer structure similar to the native trilayer morphology [26, 27]. Substrates made from PLLA have shown good cell adhesion, growth, and tissue development in annulus fibrosus, retinal, and heart valve tissue engineering applications [28, 29]. To date, only one study has directly compared the performance of PCL and PLLA leaflet substrates with isotropic structural and mechanical properties [28]. Further, its 15-day cell culture does not give insightful information for heart valve tissue engineering. So, the effects of anisotropic microfibrous trilayer substrates prepared from PCL or PLLA, on functional leaflet tissue formation have not been investigated in vitro.

In this study, we aim to determine whether electrospun PCL or PLLA fiber is a superior biomaterial for developing trilayer substrates with native anisotropic structure and mechanical properties for heart valve tissue engineering. Varying polymer concentrations and applying native fiber orientations in each layer, highly anisotropic trilayer substrates were produced from PCL or PLLA polymer separately. PVICs were seeded on these substrates and cultured for one month to produce tissue-engineered constructs. We investigated the effects of the trilayer structure, mechanical anisotropy, and flexural properties of the PCL and PLLA substrates on the orientation and growth of the cells and their produced ECM components, protein and gene expression, and the mechanical properties of the tissue-engineered constructs.

2. Materials and methods

2.1. Leaflet-shaped collector fabrication

A leaflet-shaped collector was designed using CAD/CAM software. Using this design, metal collectors were crafted. The collectors had a semilunar shape, a curved top surface, and three cylindrical holes for mounting to the spindle for their rotations.

2.2. Substrate Fabrication

Polycaprolactone (MW: 80 KD, Sigma Aldrich, USA) and poly (L-lactic acid) (MW: 90 KD, Polyscience Inc, USA) solutions were prepared in chloroform (Sigma Aldrich, USA) at 10% (wt/v) and 14% (wt/v) polymer concentration. The solutions were put into syringes and mounted on a syringe pump (Harvard Apparatus, USA) separately. The syringe pump was mounted to a stage for horizontal motion. The collector used for substrate fabrication was mounted to a spindle. By altering the rotational axes of the collector, different fiber alignments were obtained in a single leaflet substrate (Fig. 1). The circumferential layer was electrospun over the collector spinning along its sagittal axis (Fig. 2a) with the following electrospinning parameters: voltage - 10 kV, solution flow rate – 1.2 ml/hr, spinneret-collector distance- 10 cm, and a spin rate of 3,000 RPM to produce aligned microfibers. The random layer was produced by spinning the collector along its sagittal axis (Fig. 2b) and electrospinning directly over the circumferential layer with a slow spin (125 RPM). Finally, the aligned radial layer was produced by mounting the collector along its coronal axis (Fig. 2c) at a spin speed of 4,000 RPM. The random and radial layers were electrospun with electrospinning parameters: voltage - 14 kV, solution flow rate – 0.7 ml/hr, and spinneret-collector distance- 12 cm. The electrospinning time for each layer was determined so that the trilayer construct thickness was approximately 0.3 mm with a layer thickness ratio of 45% circumferential, 35% random, and 20% radial [14, 30]. Both PCL and PLLA leaflet substrates were produced using these specifications.

Figure 1:

Figure 1:

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Schematic of a syringe pump, high-speed spindle, and leaflet-shaped spinning collector applied to produce polymeric leaflet substrates for heart valve tissue engineering. Each layer is made in sequence at different spin speeds, starting with the circumferential, then random and radial layers.

Figure 2:

Figure 2:

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Fabrication of the electrospun trilayer PCL and PLLA substrates and structure of each layer imaged with an SEM. (a-b) A collector was spun along its sagittal axis at two different speeds to produce the circumferentially (a) and randomly (b) oriented fiber layers, respectively. (c) The collector was then rotated along its coronal axis to produce the radial layer. (d-e) The PCL fibers were aligned (shown by a double-headed arrow) along the circumferential direction, as shown at low (d) and high (e) magnifications. (f-g) The circumferentially oriented PLLA fibers appeared aligned and compact at low (f) and high (g) magnifications. (h-i) The randomly oriented PCL microfibers at low (h) and high (i) magnifications. (j-k) PLLA fibers with randomly oriented morphology imaged at low (j) and high (k) magnifications. (l-m) The radial PCL fibers were aligned at low (l) and high (m) magnifications. (n-o) The PLLA microfibrous structure was highly aligned in the radial direction, as shown at low (n) and high (o) magnifications.

2.3. Contact Angle

The hydrophilicity of the substrates was evaluated by applying the sessile drop method using a Model 210 Goniometer (Ramé-Hart Instrument Company, USA). A 1.5 μl distilled water droplet was pipetted onto the surface of the substrate, and the contact angle was measured using DROPimage Pro software (Ramé-Hart Instrument Company, USA).

2.4. Crystallinity

The crystallinity of the substrates was determined using X-ray diffraction (XRD) (Rigaku, Japan) with a monochromatic Cu anode at 20 kV and 30 mA in a 2θ range of 15–80°.

2.5. VIC Isolation and Expansion

Aortic valve leaflets were excised from pig hearts obtained from a local butcher shop. The leaflets were washed with PBS and placed in trypsin solution for 5 min at 37°C. A sterile cotton swab was gently rotated on the leaflet surface to remove the endothelial layer. The leaflets were then digested in 0.5% (wt/v) type I collagenase (Worthington Biochemical, Lakewood, NJ) solution in PBS and incubated at 37 °C for five hours. The digested samples were then placed in a centrifuge for 10 mins at 1000 rpm at 4°C. The isolated PVICs were expanded in tissue culture flasks with normal tissue culture media containing Dulbecco’s modified Eagle’s medium (DMEM) supplemented with 10% fetal bovine serum and 1% penicillin-streptomycin (Sigma Aldrich, USA).

2.6. Cell Seeding

The leaflet substrates (of both types - PCL and PLLA) were disinfected using 70% ethanol for 24 hours at room temperature. Then, they were rinsed, first with a copious amount of sterile water and then with sterile PBS. The substrates were then soaked in culture media before cell seeding. They were transferred to sterile cell culture plates and seeded with PVICs at passage three and a concentration of 500k cells/cm2. The substrates were cultured for 30 days, and the media was changed every three days.

2.7. Proliferation Test

Approximately 300k PVICs were seeded onto the substrates (n=5 of each type). The substrates were cultured, and an Alamar Blue Assay was conducted at 1-, 3-, 8-, and 14-day time points. The absorbance of each sample was measured using a BioTek Epoch 96-well plate spectrophotometer in triplicate. A calibration curve was generated using known cell numbers and applied to the sample data to determine the cell numbers on the substrates.

2.8. Mechanical Testing

The leaflet substrates and tissue-engineered constructs (n=5 of each type) were cut radially and circumferentially into 2.5 mm × 10 mm strips. The sample thickness was measured with a caliper gauge (Mitutoyo, Japan). Then, the strips were sandwiched between paper frames with a 5.5 mm × 5.5 mm window to develop tensile test samples. Superglue was used to hold the ends of the strips to the paper frames, leaving the middles of the strips open. The samples were loaded into a tensile tester (TestResources, Canada), and the edges of the paper frame were cut. Then, a pre-load of 0.01 N was applied to the sample to remove any slack as described in the previously published literature [31]. The tensile test of the samples was conducted in room-air conditions. However, tissue-engineered constructs remained hydrated in PBS till the completion of tensile testing. The samples were pulled in the axial direction with a head rate of 6 mm/s. The displacement, stress, and strain were recorded to determine the elastic modulus, yield stress, and ultimate tensile strength of the samples. The circumferential elastic modulus or ultimate tensile strength was divided by the radial elastic modulus or ultimate tensile strength to calculate the anisotropic ratios of elastic modulus and ultimate tensile strength, respectively [14].

The flexural modulus of leaflet substrates and tissue-engineered constructs (n=5 of each type) were measured using the method developed in previous literature [5, 32]. According to this method, flexural testing of the leaflet substrates or tissue-engineered constructs was conducted in three steps: I) the substrate/tissue-engineered constructs were cut into rectangle-shaped strip samples of known length, width, and thickness, II) the substrate/tissue-engineered construct strips were glued onto cylindrical plastic holders attached to a compression tester and III) the substrate/tissue-engineered construct strips were compressed at a rate of 1 mm/s. The force and deflation of the substrate were measured, and the flexural modulus was calculated based on linear flexural-bending theory.

2.9. Collagen & GAG quantification

The tissue-engineered constructs (n=4 of each type) were cut into 2.5 mm × 2.5 mm sections and weighted before protein extraction. For collagen extraction, the samples were placed in 1 mL of collagen extraction solution (0.5 M Acetic Acid + 1 U/mL pepsin in water) overnight in an orbital rocker at 20°C. Then, a Sircol insoluble collagen assay kit (Biocolor Ltd., United Kingdom) was used to quantify the amount of collagen in the tissue-engineered constructs following the protocols provided by the company. For GAG extraction, the samples were placed in 1 mL of GAG extraction solution (4 M Guanidine-HCL and 0.5 M Sodium Acetate in water) overnight at 8°C. A Sircol GAG assay kit (Biocolor Ltd., United Kingdom) was used to quantify the amount of GAG in the tissue-engineered constructs following the protocol provided by the company. The collagen and GAG concentrations were determined using an Epoch microplate spectrophotometer (BioTek, USA) set at 550 nm and 656 nm wavelengths, respectively.

2.10. Histology & Immunostaining

Tissue-engineered constructs were fixed in 10% formalin overnight, and then rinsed in PBS before dehydration. The samples were processed and embedded in paraffin. Then, the samples were cut either radially or circumferentially and mounted onto slides. The sections were stained with hematoxylin and eosin (Thermo Fisher Scientific, USA), Masson’s trichrome (Thermo Fisher Scientific, USA), Alcian blue (Abcam, USA), and Von Kossa (Abcam, USA) following the company protocol. Images of the stained samples were obtained using a Zeiss Axiovert 200M with an ORCA-ER camera at 20x magnification.

Unstained slides were further prepared for immunohistochemical analysis using antibodies against vimentin (ab92547, Abcam, USA) and α-smooth muscle actin (α-SMA) (ab124964, Abcam, USA). First, the samples were deparaffinized, then permeabilized and blocked using goat serum in PBS with Triton X-100. Then, the sections were incubated in the primary antibody at 4°C overnight. Finally, the slides were incubated in the secondary antibody for 2 hours. Between the steps, the slides were rinsed with PBS. DAPI was used to mount the glass coverslip. The stained slides were then imaged using a GSD microscope (Lecia, Germany).

Tissue-engineered constructs were cut into 3 mm × 3 mm sections. As described above, the samples were dehydrated, permeabilized, blocked, and stained with a vimentin marker. Then, the samples were counterstained using Hoechst 33342 (Thermofisher, USA) according to the manufacturer’s protocol. The samples were washed with PBS and placed into a glass-bottom dish with either the radial or circumferential surfaces facing up and imaged with a spectral confocal microscope (Lecia, Germany).

2.11. Scanning Electron Microscopy imaging

The substrates (n=2 of each type) were cut into 4 mm × 4 mm squares and mounted onto metal stubs. The surfaces were sputter-coated with 25 nm of platinum, and a scanning electron microscope (SEM) (FEI Company, USA) was used to image the surface at 10 kV. Tissue-engineered constructs (n=2 of each type) were cut to the previously specified dimensions and fixed in 10% formaldehyde (Sigma Aldrich, St. Louis, MO) overnight at 4°C. The samples were freeze-dried and mounted onto stubs with either the circumferential or radial direction pointed upward.

2.12. Polymer degradation test

After in vitro tissue engineering for one month, several tissue-engineered constructs (n=2 of each type) were cut into strips and soaked in a 5.25% sodium hypochlorite to remove the tissue materials from the constructs, then rinsed with deionized water. The samples were freeze-dried and imaged with an SEM to examine any degradation of fiber polymers in the substrate.

Degradation was also assessed based on the weight loss of the substrates after 30 days of culture. Substrates (n=5 of each type) were weighted before in vitro tissue engineering. After cell culture, the tissue material was removed from the tissue-engineered constructs (n=5 of each type) as previously described. The constructs were washed with PBS, air dried for 4 hr, and weighed.

2.13. Gene Expression

The tissue-engineered constructs (n = 4 of each type) were rinsed with PBS, disrupted through stirring, and the cells were lysed using RLT buffer (supplied with the kit), supplemented with β-mercaptoethanol (Sigma Aldrich, St. Louis, MO). The samples were then homogenized with QIAshredder spin columns (Qiagen, Germany), and the lysate was transferred to RNAeasy spin columns (Qiagen, Germany). Using the manufacturer’s protocol, the RNA was extracted from the homogenized solutions—DNase I (Life Technologies, USA) was used to treat the extracted RNA. The RNA was reverse-transcribed into cDNA using a High-Capacity cDNA Reverse Transcription Kit and its supplied protocol. The cDNA transcripts were then probed with TaqMan assays for α-smooth muscle actin (ACTA2, Ss04245588_m1), vimentin (VIM, Ss04330801_gH), and type I collagen (COL1A1, Ss03373341_g1) using a TaqMan Universal PCR Master Mix (Applied Biosystems, USA). Amplification was performed using an AriaMx Real-time PCR System (Agilent, USA) using our previously described conditions [33]. Using the comparative cycle threshold (Ct) method, target gene data were normalized using β-actin (ACTβ, Ss03376563_uH) and then analyzed.

2.14. Statistical Analysis

All numerical data are expressed as mean ± standard deviation (S.D.). Unpaired t-tests and one-way ANOVA with Tukey’s post-hoc tests were applied to two-group and 3-group comparisons, respectively. Statistical significance was established at a two-tailed P <0.05.

3. Results and discussion

3.1. Substrate structure and mechanical properties

3.1.1. Physical characteristics of leaflet substrates

As our goals are to determine the superiority between PCL and PLLA polymers for heart valve tissue engineering and determine the effects of anisotropic properties of the substrate on cells, in this study, we designed and developed microfibrous trilayer leaflet substrates from these polymers mimicking the morphologies observed in native tissue [28, 33]. SEM imaging was performed to verify the physical characteristics of the substrates, and the images confirmed the orientations of the microfibers in the trilayer substrates at high and low magnifications (Fig. 2d–o). In both circumferential and radial layers of the PCL and PLLA leaflet substrates, the fibers were highly aligned, and in their random layers, the fibers were randomly oriented. The PCL and PLLA leaflet substrates had fiber diameters of 4–6 μm and 2–4 μm, and pore sizes of 15–18 μm and 10–15 μm, respectively. PCL fibers have a higher crystallinity than PLLA fibers, resulting in the production of larger diameter fibers during electrospinning [28]. The pore sizes of both substrates are above 10 μm, which is adequate for enabling cell infiltration [34].

3.1.2. Contact Angle

The surface energy of a substrate determines its hydrophilicity, i.e., cell adhesion to the substrate. Substrates with higher hydrophilicity have higher surface energy, and thus show greater cell adhesion [35]. The water contact angle test was performed on the PCL and PLLA substrates to find their hydrophilicity and the angles on the PCL and PLLA substrates were 124.28 ± 0.31° and 98.61 ± 0.69°, respectively (Fig. 3a–b). Thus, the fibrous PLLA substrates had higher hydrophilicity than fibrous PCL substrates, i.e., the surface of the PLLA fibers had more hydrophilic characteristics than PCL fibers; so, substrates with PLLA fibers can encourage better cell adhesion and infiltration.

Figure 3:

Figure 3:

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Contact angle and XRD data for the PCL and PLLA fibrous substrates. (a-b) Contact angle images of distilled water droplets on flat PCL (a) and PLLA (b) substrates. (c) XRD data showed the semicrystalline polymer structure of PCL substrates and the amorphous polymer structure of PLLA substrates.

3.1.3. Crystallinity

Amorphous polymers are typically more flexible because they contain randomly oriented polymer chains with lower chain densities and less intermolecular bonding than crystalline polymers [36–38]. The crystallinity of polymeric materials that determines the physical and mechanical properties of the fibers in the substrates was measured by XRD. The XRD pattern for PCL substrates shows one prominent peak at a 2θ of 21.28° and a smaller peak at 23.71° (Fig. 3c). These peaks indicated that PCL substrates had a semicrystalline polymer structure. The PLLA substrate showed no major peaks, indicating that PLLA substrates had a more amorphous polymer structure than PCL substrates [28].

3.1.4. Mechanical properties of the leaflet substrates

Heart valve leaflets have a trilayer fiber structure that exhibits high tensile properties along the circumferential direction and relatively low tensile properties in the radial direction. These anisotropic properties influence cell proliferation, their gene expression, and the production of ECM components [39]. Using a tensile tester, the tensile properties - in their circumferential and radial directions – and flexural properties of PCL and PLLA substrates were measured (Fig 4). The tensile moduli and strengths in the circumferential direction were 10.45 ± 0.85 MPa and 2.23 ± 0.42 MPa for the PCL substrates and 40.63 ± 2.76 MPa and 2.01 ± 0.33 MPa for the PLLA substrates, respectively (Fig. 5a). The tensile moduli and strengths in the radial direction were 2.89 ± 0.17 MPa and 1.06 ± 0.21 MPa for the PCL substrates and 9.21 ± 0.99 MPa and 0.41 ± 0.26 MPa for the PLLA substrates (Fig. 5b). The PLLA substrates had a significantly higher tensile modulus than the PCL substrates, but their tensile strengths were quite similar. The difference in tensile moduli could be due to the higher glass transition temperature (Tg) of PLLA (60–65 °C) than PCL (−60 °C) [40]. The higher unidirectional tensile properties of the PLLA substrate may enable it to bear the hemodynamic load more effectively than PCL substrates.

Figure 4:

Figure 4:

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Schematic of flexural testing of the leaflet substrates: I) loading a testing sample into the tester in an extended but relaxed position and II) compressing the sample to obtain its deflection for flexural modulus measurement.

Figure 5:

Figure 5:

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Mechanical testing of the PCL and PLLA heart valve leaflet substrates. (a-b) Tensile moduli and strengths of the PCL and PLLA substrates in the circumferential (a) and radial (b) directions, respectively. (c) The anisotropic ratios (circumferential/radial) of the tensile modulus and strength of the PCL and PLLA substrates. (d) The flexural moduli of the PCL and PLLA substrates.

The tensile modulus and tensile strength anisotropic ratios were 3.61 ± 0.87 and 2.11 ± 0.47 for the PCL and 4.41 ± 0.99 and 5.03 ± 0.42 for the PLLA substrates (Fig. 5c). The flexural moduli of the PCL and PLLA leaflet substrates were 14.45 ± 0.36 MPa and 10.67 ± 0.59 MPa, respectively (Fig. 5d), i.e., PLLA leaflet substrates were more flexible than PCL substrates. PLLA substrates had lower crystallinity than PCL substrates, which is associated with their lower flexural modulus [28, 41, 42]. The higher anisotropic tensile properties and flexibility of PLLA substrates could bring a more native-like environment to the cells in PLLA substrates than PCL substrates during heart valve tissue engineering [14].

3.2. Development of Tissue-engineered Constructs

3.2.1. Cell proliferation on the leaflet substrates

A proliferation assay was conducted to evaluate the increase in cell numbers and compare the cytocompatibility of the PCL and PLLA substrates. Initially, there were significantly more cells on the PLLA than on PCL substrates, suggesting that PLLA had more cell adhesion. The cell proliferated continuously on both substrates. After 14 days of cell culture, the PLLA substrates showed 16% greater cell proliferation than the PCL substrates (Fig. 6a). The superior cell adhesion and proliferation on PLLA substrates compared to PCL could be due to the PLLA polymer chains containing more ester, ether, and carboxyl groups leading to higher surface energy [29, 40, 43]. Also, this result could have indicated that cells responded positively to the more anisotropic and flexible PLLA substrates [28, 44].

Figure 6:

Figure 6:

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Assessment of cellular proliferation on PCL and PLLA substrates. (a) Cellular proliferation in PCL and PLLA substrates over a 14-day cell culture. (b-e) Images of ECM material on the circumferential and radial surfaces of the PCL (b-c) and PLLA (d-e) tissue-engineered constructs after a one-month static culture. Circumferential and radial surfaces of the PCL (f-g) and PLLA (h-i) tissue-engineered constructs stained for vimentin and Hoechst. Cross-sections of PCL (j) and PLLA (k) tissue-engineered constructs stained for vimentin and Hoechst.

SEM imaging of the PCL and PLLA tissue-engineered constructs confirms that ECM material deposited on the radial and circumferential surfaces of the substrates and formed a confluent layer (Fig. 6b–e). The confocal immunostaining images of those surfaces stained with a vimentin marker showed that cells adhered to the surface of both substrates and produced vimentin proteins (Fig. 6f–i). Further, their cross-sectional images at high magnification indicated that cells infiltrated the trilayer structures of both PCL and PLLA substrates (Fig.6j–k). The sufficient pore size (>10 μm) in the substrates was favorable for cell infiltration into the substrates.

3.2.2. Biodegradation of the leaflet substrate

Substrates for tissue engineering should biodegrade and support tissue formation through cell proliferation and ECM production. PCL and PLLA substrates generally biodegrade completely in vivo in 8 months and 1–2 years, respectively. However, in previous studies, both materials have shown mild biodegradation in vitro due to cell-mediated enzymatic and oxidative processes, resulting in fiber damage [24, 45]. The ECM and cellular materials on the substrates of the tissue-engineered constructs were removed to identify any fiber degradation. Although there was no visible surface erosion on the PCL fibers (Fig. 7a), mild surface erosion was observed on the PLLA fibers (shown by arrows, Fig. 7b). Further, the weights of the original PCL and PLLA substrates and their cleaned counterparts obtained from tissue-engineered constructs after tissue removal were measured to determine any loss of materials during tissue engineering. There were no significant differences in the weights of the original PCL or PLLA substrates compared to the PCL or PLLA substrates in the tissue-engineered constructs (Fig. 7c). Moreover, both substrates conserved their fibrous structures; thus, the substrate’s ability to guide cell growth within the tissue-engineered constructs was not affected [46, 47].

Figure 7:

Figure 7:

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Biodegradation of the substrate fibers and tensile properties of the PCL and PLLA tissue-engineered constructs after a one-month culture. (a-b) SEM images of the PCL (a) and PLLA (b) fibers in the tissue-engineered constructs (arrows show surface erosion on the fibers). (c) Weight (mg) of the acellular PCL or PLLA substrates and tissue-engineered constructs. (d-e) Tensile moduli and strengths of PCL and PLLA tissue-engineered constructs in the circumferential (d) and radial (e) directions. (f) Tensile modulus and tensile strength anisotropic ratios of the PCL and PLLA tissue-engineered constructs. (g) Flexural moduli of both PCL and PLLA tissue-engineered constructs.

3.2.3. Mechanical properties of the tissue-engineered constructs

The substrate’s materials and physical and mechanical properties influence cell adhesion, proliferation, and ECM production in tissue engineering, which ultimately determines the mechanical properties of the tissue-engineered constructs. So, the tensile properties of tissue-engineered constructs were measured and compared. Further, they were compared with those of PCL and PLLA substrates to find the influence of substrate properties on tissue engineering. After tissue engineering for four weeks, the tensile modulus and strengths in the circumferential direction were 10.47 ± 1.55 MPa and 1.75 ± 0.51 MPa for the PCL tissue-engineered construct and 34.01 ± 3.85 MPa and 1.58 ± 0.44 MPa for the PLLA tissue-engineered construct, respectively (Fig. 7c). The tensile modulus and strengths in the radial direction were 3.14 ± 0.46 MPa and 0.81 ± 0.18 MPa for the PCL tissue-engineered construct and 9.71 ± 0.96 MPa and 0.85 ± 0.28 MPa for the PLLA tissue-engineered construct (Fig. 7d). Tensile properties of the PCL tissue-engineered constructs compared to PLLA tissue-engineered constructs showed a similar trend observed in the tensile properties of PCL and PLLA substrates.

Tensile properties (modulus and strength) of the PCL tissue-engineered constructs in the circumferential (10.47 ± 1.55, 1.75 ± 0.51 MPa) or radial direction (3.14 ± 0.46, 0.81 ± 0.18 MPa) were not significantly different from that of PCL substrates in the circumferential (10.45 ± 0.85, 2.23 ± 0.42 MPa) and radial directions (2.89 ± 0.17, 1.06 ± 0.21 MPa). The PLLA tissue-engineered constructs had a lower tensile modulus (34.01 ± 3.85 MPa) (statistically significant, p <0.05) and similar strength (1.58 ± 0.44 MPa) in the circumferential direction compared to the PLLA substrates (40.63 ± 2.76, 2.01 ± 0.33 MPa). Conversely, the PLLA tissue-engineered constructs had a similar tensile modulus (9.71 ± 0.96 MPa) and higher strength (0.85 ± 0.28 MPa) (statistically significant, p <0.05) in the radial direction compared to the original PLLA substrates (9.21 ± 0.99, 0.41 ± 0.26 MPa). Substrates made from electrospinning solutions with a low PLLA polymer concentration have been shown to have greater cell growth [48, 49]. Greater cell proliferation in the radial layer, made from a low PLLA polymer concentration, of the PLLA tissue-engineered constructs could account for the higher tensile modulus and strength in the radial direction than the PLLA substrate.

The tensile modulus anisotropy ratio (3.33 ± 0.84) and strength anisotropy ratio (2.18 ± 0.54) of the PCL tissue-engineered constructs (Fig. 7e) were not significantly different from (3.61 ± 0.87, 2.11 ± 0.47) that of the PCL substrates. The tensile modulus anisotropy and tensile strength anisotropy ratios for the PLLA tissue-engineered constructs were 3.52 ± 0.97 and 1.86 ± 0.51, respectively (Fig. 7e). Although the tensile modulus anisotropy of the PLLA tissue-engineered constructs was not significantly different from that of the PLLA substrates (4.41 ± 0.99), the tensile strength anisotropy of the PLLA tissue-engineered constructs was significantly lower (p <0.005) than that of the PLLA substrates (5.03 ± 0.42). Cell proliferation in the radial layer of the PLLA tissue-engineered constructs and surface erosion mentioned earlier could account for the disparity in tensile properties of PLLA tissue-engineered constructs.

The flexural modulus (15.97 ± 0.66 MPa) of the PCL tissue-engineered constructs (Fig. 7f) was significantly higher (statistically significant, p <0.05) than that (14.45 ± 0.36 MPa) of PCL substrates. The produced ECM within the PCL tissue-engineered constructs could be a possible reason for this increase in PCL tissue-engineered constructs. In comparison, the flexural modulus (10.27 ± 0.81 MPa) of the PLLA tissue-engineered construct (Fig. 7f) was not significantly different from that (10.67 ± 0.59 MPa) of the original substrate.

3.2.4. Histology

The fabricated substrates applied for leaflet tissue engineering had a trilayered structure with circumferential, random, and radial orientations in three respective layers mimicking the native leaflet tissue structure. Electrospun fibers within the substrates guide the orientation and morphology of cells and their produced ECM along the fibers in three layers of the substrates [50, 51]. Thus, the tissue-engineered constructs were histologically assayed to verify the orientation of cells and their produced ECM components in different layers of the substrates. Radial and circumferential sections stained with H&E showed that the PCL tissue-engineered constructs had an evident trilayer structure, with distinct regions of aligned and randomly oriented cells and collagen fibrils (Fig. 8a–b). Due to the temperature sensitivity of PCL polymer, the sample embedding (in paraffin) process partially deformed the PCL substrate in the PCL tissue constructs [52]. Similarly, H&E stained PLLA tissue-engineered constructs showed trilayer, oriented structure, with distinct regions of aligned and randomly oriented cells and collagen fibrils (Fig. 8c–d). Further, it is apparent that PLLA tissue-engineered constructs had more cellularization and collagen in the circumferential and radial layers compared to PCL tissue-engineered constructs, possibly due to aligned fibrous structure combined with high flexural capacity in those layers of PLLA substrates than of PCL substrates.

Figure 8:

Figure 8:

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Histological images of PCL and PLLA tissue-engineered constructs developed after one month of static culture. (a-d) H&E-stained circumferential cross-sections and radial cross-sections of PCL (a-b) or PLLA (c-d) tissue-engineered constructs show aligned circumferential (C) and radial (R) layers and randomly oriented middle (M) layer. (e-h) Masson’s trichrome-stained circumferential cross-sections and radial cross-sections of PCL (e-f) or PLLA (g-h) tissue-engineered constructs show aligned circumferential (C) and radial (R) layers and randomly oriented middle (M) layer. (i-l) Alcian blue-stained circumferential cross-sections and radial cross-sections of PCL (i-j) or PLLA (k-l) tissue-engineered constructs show aligned circumferential (C) and radial (R) layers and randomly oriented middle (M) layer. (m-p) Von Kossa-stained circumferential cross-sections and radial cross-sections of PCL (m-n) or PLLA (o-p) tissue-engineered constructs show aligned circumferential (C) and radial (R) layers and randomly oriented middle (M) layer. Alignment in the circumferential and radial directions can be observed in the C and R layers (shown with a double-headed arrow). Dots in the R layer of the circumferential cross-sections and the C layer of the radial cross-sections show the alignment perpendicular to the cross-section. The random-middle (M) layer showed the random orientation.

Images of Masson’s trichrome-stained PCL tissue-engineered constructs confirmed that collagen fibrils (blue color) were aligned in their circumferential and radial layers (Fig 8e–f). The middle layer had randomly oriented collagen fibrils. The developed collagen fibrils were quite uniformly distributed throughout the constructs. Similarly, PLLA tissue-engineered constructs had significant collagen presence and alignment in the outer two layers, possibly due to aligned fibrous structure combined with high flexural capacity in those layers of PLLA substrates (Fig. 8g–h). Collagen fibrils also were randomly oriented in the middle random layer.

Alcian blue was used to stain the tissue-engineered constructs for GAG - an essential ECM component making up most of the random layer in leaflet tissue [53]. Circumferential and radial sections of the PCL tissue-engineered constructs exhibited GAG protein production throughout the constructs (Fig. 8i–j). The PLLA tissue-engineered constructs showed more uniform GAG distribution throughout the structure compared to the PCL tissue-engineered constructs (Fig. 8k–l). One possible reason for this observation is that there was more extensive cellular growth in the PLLA tissue-engineered constructs due to the high flexibility of the substrate. GAG production is elevated by cells in flexible and soft substrates [54, 55]. Further, PLLA tissue-engineered constructs had higher cell densities. High cell densities are associated with increased GAG production because the negatively charged polysaccharides enable increased binding to essential nutrients and proteins for cell and tissue growth [53, 56].

Calcification is considered one of the most frequent causes of heart valve failure. It can occur when activated VICs are exposed to non-native stresses, bone morphogenetic proteins, and transcription factors that cause their differentiation into osteoblasts and osteoclasts phenotypic cells [57]. The tissue-engineered constructs were stained using Von Kossa to determine if any calcium nodules (black stain) were formed. Neither PCL nor PLLA tissue-engineered constructs showed evidence of any calcification (Fig. 8m–p).

3.2.5. Quantification of ECM Components within the leaflet substrates

Native leaflet structure and functions are dictated by its ECM components, including collagen, GAG, and elastin [58, 59]. Thus, the ECM (collagen and GAG) contents in PCL and PLLA tissue-engineered constructs were quantified using biochemical assays. Although the quantities of collagen and GAG in PLLA tissue-engineered constructs were higher than that in the PCL tissue-engineered constructs, they were not significantly different (Fig. 9).

Figure 9:

Figure 9:

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Quantification of ECM components - collagen and glycosaminoglycans (GAG) in PCL and PLLA tissue-engineered constructs.

3.2.6. Gene expression and Immunofluorescence staining of the leaflet substrates

During tissue-engineering, VICs become activated and undergo transdifferentiation into myofibroblast, expressing high levels of vimentin, alpha-smooth muscle actin (α-SMA), and collagen type 1 [14, 60]. The structural and mechanical properties of the substrates also influence the phenotypes of VICs [1]. Therefore, the gene expression of the PVICs in the tissue-engineered constructs was characterized by reverse transcription PCR analysis (Fig. 10a). Although the vimentin expression of the cells in the PCL and PLLA tissue-engineered constructs was not significantly different, their α-SMA gene expression in the PLLA tissue-engineered constructs was significantly higher than that in the PCL tissue-engineered constructs. This observation indicates that more cells were in a growth state in the PLLA tissue-engineered constructs. Also, collagen type 1 expression of cells was significantly higher in the PLLA tissue-engineered constructs than in the PCL tissue-engineered constructs. PLLA substrates were unidirectionally stiffer but more flexural and anisotropic than PCL substrates, which may have caused greater expression of α-SMA and collagen type-1 [61].

Figure 10:

Figure 10:

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α-SMA, vimentin, and collagen type 1 expression in PCL and PLLA tissue-engineered constructs after one month of in vitro tissue engineering. (a) Vimentin, α-SMA, and collagen type 1 gene expression of the cells in the tissue-engineered constructs. (b-e) Images of α-SMA marker stained circumferentially and radially sectioned PCL (b-c) or PLLA (d-e) tissue-engineered constructs showed distinct orientations (C, M, and R) of their α-SMA protein. (f-i) Images of vimentin marker stained circumferentially and radially sectioned PCL (f-g) or PLLA (h-i) tissue-engineered constructs showed distinct orientations (C, M, and R) of their vimentin filament protein. Alignment in the circumferential and radial directions can be observed in the C and R layers (shown with a double-headed arrow). Dots in the R layer of the circumferential cross-sections and the C layer of the radial cross-sections show the alignment perpendicular to the cross-section. The random-middle (M) layer showed the random orientation.

Immunofluorescence staining was performed with α-SMA (red) and vimentin (green) protein markers to identify the growing status of the tissue-engineered constructs and confirm the above gene expression results. The PCL tissue-engineered constructs showed α-SMA positive protein expression throughout the structure, indicating that remodeling occurred throughout the trilayer constructs (Fig. 10b–c). Comparatively, PLLA tissue-engineered constructs had lower α-SMA protein expression than the PCL tissue-engineered constructs (Fig. 10d–e). Vimentin protein expression was detected throughout the PCL tissue-engineered constructs (Fig. 10f–g), whereas, in the PLLA tissue-engineered constructs, this protein expression was slightly less (Fig. 10h–i). It appeared that more PVICs were in a proliferation state in the PLLA tissue-engineered constructs than in the PCL tissue-engineered constructs, which could be caused by the higher flexibility of the PLLA fibers than PCL fibers [62, 63]. Considering the above results, we believe that during tissue engineering, the cultured PVICs were transdifferentiated into myofibroblast phenotypic cells for their growth and production of ECM components [1, 11]. In these constructs, three layers were distinctly observed; protein materials were aligned in circumferential and radial layers, and they were randomly oriented in the random layers of the tissue constructs.

This is the first report of a comparative study on the development of native-mimicked trilayer, oriented microfibrous leaflet substrates from two biocompatible and biodegradable polymers - PCL and PLLA, and their effectiveness in heart valve leaflet tissue engineering. Among them, tissue-engineered constructs developed with PLLA substrates had better cell proliferation, gene expression, and growth compared to that with PCL substrates due to higher flexibility and mechanical anisotropicity in PLLA substrates. This study brings a potential design and development of native-mimicked, microfibrous PLLA substrates for successful heart valve tissue engineering.

4. Conclusion

PCL and PLLA trilayer microfibrous leaflet substrates mimicking the native leaflet structure were produced by electrospinning. PLLA substrates had higher tensile properties than PCL substrates but were more flexible and anisotropic. The PLLA substrates had higher hydrophilicity, lower crystallinity, and higher flexibility and anisotropic properties than the PCL substrates, leading to greater cell adhesion and proliferation. The flexural properties of the PLLA substrate did not change after tissue engineering; however, the PCL tissue-engineered construct became less flexible as ECM material formed. The GAG and collagen quantities in the PCL and PLLA tissue-engineered constructs were not significantly different. Both tissue-engineered constructs had similar vimentin gene expression, whereas the PLLA tissue-engineered constructs showed higher α-SMA and collagen type 1 expression than PCL tissue-engineered constructs. This observation demonstrated that the higher flexibility and anisotropic properties of the PLLA substrates caused more cells to be in a growth state in the PLLA tissue-engineered constructs than in the PCL tissue-engineered constructs. This study directly compared trilayer PCL and PLLA leaflet substrates and provided significant insight into the positive influence of substrates with the appropriate hydrophilicity, flexural properties, and anisotropicity on PVICs to promote healthy heart valve tissue engineering.

Acknowledgments

This work is supported by the National Institutes of Health (NIH R00HL134823).

Footnotes

Declaration of Competing Interest

The authors have declared that there is no competing financial interest or personal relationships with other people or organizations that could inappropriately influence this work.

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