Direct Laser Microperforation of Bioresponsive Surface-Patterned

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Direct Laser Microperforation of Bioresponsive Surface-Patterned Films with Through-Hole Arrays for Vascular Tissue-Engineering Application Zuyong Wang,†,‡ Zheren Du,‡ Jerry Kok Yen Chan,§,∥,⊥ Swee Hin Teoh,# Eng San Thian,*,† and Minghui Hong*,‡ †

Department of Mechanical Engineering, National University of Singapore, 9 Engineering Drive 1, Singapore 117576, Singapore Department of Electrical and Computer Engineering, National University of Singapore, 2 Engineering Drive 3, Singapore 117576, Singapore § Department of Reproductive Medicine, KK Women’s and Children’s Hospital, 100 Buikit Timah Road, Singapore 229899, Singapore ∥ Department of Obstetrics and Gynaecology, Yong Loo Lin School of Medicine, National University of Singapore, 14 Medical Drive, Singapore 117599, Singapore ⊥ Cancer and Stem Cell Biology, Duke-NUS Graduate Medical School, 8 College Road, Singapore 169857, Singapore # School of Chemical and Biomedical Engineering, Nanyang Technological University, Singapore 637459, Singapore ‡

ABSTRACT: Tissue architecture plays critical roles in the physiological functions of blood vessels. Surface-patterned films are promising to replicate cellular alignment as in the native vessels. However, for vascular tissue engineering (TE) applications, the current surface-patterned films lack structural support for the myoendothelial communications between tunica media and intima. Herein, we report the development of direct microperforation using a femtosecond laser on surface-patterned films for the native-like architecture reconstruction of blood vessels. Poly(ε-caprolactone) (PCL) thin films were surface-patterned with anisotropic microridges/grooves. Direct femtosecond laser ablation further resulted in microscale through-holes for the PCL films, without invasive thermal damage to the ridges/grooves on the nonprocessed surface. Laser fluence and pulse number were observed to significantly influence the microperforation on both hole quality and dimension. The PCL films after direct femtosecond laser microperforation exhibited improved flexible properties, without sacrificing the yield stress. Meanwhile, direct femtosecond laser microperforation resulted in PCL films with hydrophilic permeability to transport nutritional/signaling biomolecules and allowed for heterocellular protrusion ingrowth into the throughholes for physical myoendothelial contacts. Small-diameter vascular TE scaffolds based on the as-fabricated PCL films could enable a hybrid vascular wall construction with aligned stromal multilayers and a confluent endothelium similar to those of the native vascular tissue. These results showed that direct femtosecond laser microperforation could be a reliable approach for producing biomimetic films with through-holes. The developed vascular TE scaffolds with microridges/grooves and throughholes have the potential to offer structural support for vascular architecture reconstruction with the native-like stromal and endothelial components. KEYWORDS: vascular tissue engineering, myoendothelial interactions, laser microperforation, surface patterns, through-holes

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INTRODUCTION Cardiovascular disease that forms clot in the arteries (making it harder for blood to flow through) is a globally leading cause of death (∼17.3 million in 2008, which could reach ∼23.3 million by 2030).1 Surgical intervention by the implantation of vascular grafts is necessary to replace the stenotic vessels.2 While vascular substitutions made from autogenic tissues are generally used, they are not available to many patients due to widespread atherosclerosis. Meanwhile, alternative grafts prepared from either allogenic or xenogenic tissues lack growth potential, have the issue of transplant rejection, and, thus, are not the optimum choice for surgical implantation.3 These setbacks have caused © 2015 American Chemical Society

significant inadequacy of suitable grafts for millions of patients who need vascular reconstruction surgeries. Tissue engineering (TE) can be promising for cardiovascular patients and can overcome the challenges of autogenic, allogenic, and xenogenic grafts through generating the engineered, autologous, and offthe-shelf substitutes using a combination of cells, biomaterials, and engineering techniques.4 Received: June 19, 2015 Accepted: October 30, 2015 Published: November 2, 2015 1239

DOI: 10.1021/acsbiomaterials.5b00455 ACS Biomater. Sci. Eng. 2015, 1, 1239−1249

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ACS Biomaterials Science & Engineering Critical relationships existing between the tissue architecture and physiological functions are essential for vascular TE grafts in order to recapitulate the native vessels. During vasculogenesis, the monolayer of endothelial cells (ECs) forms the tunica intima to offer an antithrombogenic inner surface for the conduit of blood.5 Stromal cells such as pericytes and smooth muscle cells (SMCs) are recruited subsequently to invest in the nascent endothelial basement membrane (BM) secreted by ECs and organize into a circumferentially aligned structure known as the tunica media, which renders anisotropic tensile strength, compliance, and vasoactive responsiveness for the vessels.4 Cellular interactions between the nascent stromal and endothelial components can occur through the diffusion of signaling biomolecules and physical cell−cell contacts across endothelial BM,6−8 importantly for the functionalized cell phenotype (e.g., myogenic differentiation)9 and neovessel maturation.10 These fundamental elements of blood vessels reflect the need of de novo design for vascular TE scaffolds to facilitate engineering of the native-like vessel architecture. Recently, there is a growing recognition that the nano/ micropatterned surface can be applied to replicate the aligned cellular organization as observed in the tunica media of the native vessels.4,11 These patterned biomaterials are found to further influence the improved functional behaviors of vascular cells (e.g., cellular adhesion/proliferation12 and migration)13 and tissue constructs (e.g., mechanical strength14 and anisotropic contraction).15a,b Meanwhile, our recent studies reveal that when cultured on flexible micropatterned films, human mesenchymal stem cells (MSCs) can adjust their organization into alignment leading to up-regulated myogenic differentiation.4,16 However, except for the control of cell alignment, current nano/ micropatterned biomaterials typically from electrospinning,17 soft lithography,11 and uniaxial stretching16 lack suitable structure support for herterocellular communications between the engineered stromal and endothelial components. For example, electrospinning has been of great interest for vascular TE application due to the high surface-to-volume ratio of the generated fibrous meshes. However, during the fibrogenesis process, the fibers are subjected to uniaxial pulling in order to obtain oriented organization.18 This has been found to induce significant fiber packing with declined porosity to limit cellular ingrowth.19 In addition, although electrospun fibrous meshes from the selective removal of sacrificial fibers can achieve improved porosity, these meshes have been reported with deteriorative mechanical properties for load-bearing application.19 These surface-structured biomaterials remain incapable of engineering vascular TE grafts with the native-like complex architecture. Herein, we report the development of direct microperforation using a femtosecond laser for bioresorbable and surfacepatterned poly(ε-caprolactone) (PCL) films to obtain throughholes for vascular TE application. This method allows for noncontact and programmable microperforation without the usage of solvents or photomasks, being applicable to various films with different patterned-surfaces. The direct femtosecond-lasermicroperforated PCL films could provide a sufficient theoretical burst of strength for vascular TE scaffolds and show coexistence of through-holes and micropatterned surfaces for reconstructing the vascular wall with the native-like aligned stromal cell multilayers, a confluent endothelium, and heterocellular communications between the stromal/endothelial components (Figure 1).

Figure 1. Schematic diagram illustrating the direct microperforation using a femtosecond laser to engineer through-hole arrays on surfacepatterned PCL films for vascular tissue construction (Mic-ρ-PCL, direct femtosecond-laser-microperforated PCL films; MSCs, mesenchymal stem cells; and ECs, endothelial cells).

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EXPERIMENTAL SECTION

Materials. PCL (MW = 80,000), phosphate buffer saline (PBS), penicillin−streptomycin (PS), fluorescein isothiocyanate-dextran (FITC-dextran, MW = 40,000), paraformaldehyde (PFA), triton-X 100, bovine serum albumin (BSA), 4′,6-diamidino-2-phenylindole (DAPI), TRITC-conjugated phalloidin, and PHK cell linker dye kits were from Sigma-Aldrich (Singapore). Monoclonal mouse antihuman smooth muscle α-actin (SM α-actin) and IgG 2a k were from Dako (Singapore). Trypsin-EDTA, fetal bovine serum (FBS), Dulbecco’s modified Eagle’s medium (DMEM) and goat antimouse IgG (H + L)Alexa Fluor 594 were from Life Technologies (Singapore) and endothelial growth medium (EGM) from Lonza (Singapore). Tissue culture flasks and low-adhesion cell culture plates were from Nunc (Singapore) and Corning Costar (Singapore), respectively. Preparation of PCL Films. PCL pellets after two-roll milling were heat pressed at 80 °C and 300 MPa into a thick film (named as HPPCL). The film was then subjected to uniaxial thermal stretching at 54 °C with a draw ratio of 4.16 The uniaxial-stretched PCL thin film was named as UX-PCL. PCL films were then cut into strips and mounted on a stage. A Ti:sapphire femtosecond laser was used at a wavelength of 800 nm, pulse duration of 110 fs, and repetition rate of 1 kHz (Spectra-Physics, USA). Femtosecond laser beam scanning was controlled by U500 MMI software (Aerotech, UK), and the films were peeled off from the stage. Different laser fluences (F) at a fixed pulse number (Npulse) and different pulse numbers at a fixed laser fluence were investigated. UX-PCL after direct femtosecond laser microperforation was named as Mic-ρ-PCL. For mechanical and biological tests, PCL films were engineered at a consistent density of 400 holes/cm2 with an interhole distance of 500 μm. Morphological Characterization. The morphologies of PCL films were examined using both a light microscope (Olympus; MX50T-F) and field emission scanning electron microscope (SEM; S-4300, Hitachi, Japan). For the measurement of hole diameters, optical microscope images were taken on Mic-ρ-PCL’s top and bottom surfaces. Five samples were used for Mic-ρ-PCL at each fabrication condition (F and Npulse). For the topographical details, PCL films were imaged using SEM. The samples after sputter-gold coating were analyzed at both low and high magnifications with an accelerating voltage of 15 kV. Measurement of Mechanical Properties. PCL film samples were cut into strips of 5 × 30 mm2 (UX-PCL and Mic-ρ-PCL: the strip’s major edge following the stretching direction). Film thickness was measured at five random positions using a digimatic micrometer (APB1D, Mitutoyo corporation, Japan). Mechanical testing was carried out using a tensile testing machine (Model 3345, Instron, USA) at a load cell of 100 N and a pulling rate of 10 mm/min.20 An offset-strain of 0.005 was used for UX-PCL and Mic-ρ-PCL to determine their yield points. For HP-PCL, a low yield point was used to determine the yield stress and strain at linear-elastic regions. To facilitate comparison, the measured mechanical properties were finally normalized to those of HP-PCL. Three samples were used for each kind of PCL film. In Vitro Mass Diffusion Assays. FITC-dextran was used as a marker molecule as previously described to evaluate the mass diffusion 1240

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ACS Biomaterials Science & Engineering across Mic-ρ-PCL.21 Briefly, Eppendorf tubes after being sealed with Mic-ρ-PCL were filled with 1 mL of FITC-dextran solution (2 mg/mL). The tubes sealed with UX-PCL were used as a control group. Each tube was then put in a black bottle prefilling with 21 mL of PBS and incubated in a 37 °C water bath. Samples were taken out from the bottles at predetermined time points and measured via a fluorescence microplate reader (Infinite M1000 PRO, Tecan Group Ltd., Switzerland). FITCdextran concentration was calculated from a standard curve ranging from 0.05 to 0.001563 mg/mL. The diffusion of FITC-dextran across the PCL film was described as a normalized concentration (Ccumulative = Creal/Cfull‑diffusion·100%), where Creal represented the detected concentration of FITC-dextran in the bottle, and Cfull‑diffusion was a specific Creal when FITC-dextran’s concentrations at the inner and outer sides of the Eppendorf tube were equal. Six samples were used for each group. Cellular Isolation and Culture. Human tissue collection for research purposes was approved by the Domain Specific Review Board of National Healthcare Group, in compliance with international guidelines regarding to the use of fetal tissues for research.22 In all cases, patients gave separate written consents for the use of the collected tissues of their fetuses. MSCs were isolated from human fetal bone marrow. Single-cell suspension from the marrow components of femurs was seeded into a tissue culture flask at a density of 106 cells/cm2. Nonadherent cells were removed after 3 days of culturing in D10 (DMEM + 10% FBS + 1% PS). The adherent cells were cultured for an additional week, recovered, and characterized with stem cell phenotype.4 Cell Proliferation Assays. Cell proliferation was evaluated by enumerating nucleus number as previously described.23 Briefly, the wells for cellular culture were separated into top and bottom potions by different PCL films, with the medium inside containing 0% (D0) or 10% (D10) FBS. The group with MSCs cultured on Mic-ρ-PCL in D0/D0 (top/bottom portions) was set to determine the basal levels of cell adhesion and proliferation under FBS free conditions and used as negative control. The group with MSCs cultured on Mic-ρ-PCL in D10/ D10 was set to evaluate the cell adhesion and proliferation under normal nutritional/signaling conditions and used as positive control. For testing groups, MSCs were cultured on UX-PCL and Mic-ρ-PCL in D0/D10. MSCs (5000 cells/cm2) were seeded on the PCL films and cultured for 1 and 3 days. Cells were then fixed with PFA (3.7% in PBS) for 15 min, permeabilized by Triton-X 100 (0.1% in PBS) for 5 min, blocked by BSA (2% in PBS) for 30 min, and finally incubated with DAPI (1:1000 in PBS) for 5 min for nucleus DNA labeling. Cellular images were taken by a confocal laser scanning microscope (CLSM, FV1000, Olympus, Japan), with nucleus numbers enumerated using the built-in function of NIH ImageJ software (USA). To investigate whether the presence of cell layers on the film’s bottom side would affect FBS diffusion, Mic-ρPCL has also been preseeded with MSCs (5000 cell/cm2) on one side for 1 day of culturing and then flipped over for the second-cell seeding (5000 cell/cm2) with further culture in D0/D10 for 1 and 3 days. Three samples were used for each group. Characterization of Cell Morphology. To understand the morphological response of cells to through-holes, MSCs were seeded on Mic-ρ-PCL and cultured in D10 for 3 days. Cells after fixation were dehydrated using a graded series of ethanol−water mixtures to pure ethanol and dried for 1 week. For SEM characterization, the samples were sputter-gold coated and imaged at a high magnification of 2,000× with an accelerating voltage of 15 kV. Cell Labeling and Co-Culturing. Fluorescence labeling was performed for the coculturing of different vascular cells on the two sides of Mic-ρ-PCL, in order to visualize the transmural cell−cell contacts. Briefly, MSCs were labeled with red fluorescence using a PHK26 kit. The cells after culture expansion were collected and resuspended in diluent (2× final concentration, 0.5 mL). Meanwhile, PHK26 dye was diluted as a 2× solution (0.5 mL). The cell suspension and PHK26 dilution were immediately mixed and incubated at room temperature for 5 min. One milliliter of FBS was then added to stop the fluorescencelabeling reaction. Green fluorescent protein (GFP) labeled human umbilical vein endothelial cells (HUVECs) obtained as previously reported were provided by Mr Sandikin Dedy (National University of Singapore, Singapore).24

For coculturing, PHK26-labeled MSCs (10,000 cells/cm2) were seeded on Mic-ρ-PCL and cultured in D10 for 1 day to allow full cellular adhesion. Mic-ρ-PCL was then flipped over and transferred to an inhouse designed ring. GFP-labeled HUVECs (20,000 cells/cm2) were seeded on the opposite side of Mic-ρ-PCL. MSCs/HUVECs coculturing was performed in EGM10 (EGM + 10% FBS). The cells after 3 days of coculturing were imaged using a z-scanning model of CLSM at a magnification of 60×. CLSM images were reconstructed using the built-in function of Imaris (Bitplane, Switzerland). PHK26labeled MSCs and GFP-labeled HUVECs could be identified with red and green fluorescences, respectively. Physical MSC-HUVEC contacts were characterized as the colocalization of red and green, represented as yellow. Construction of Vascular Wall Architecture. To investigate the potential of Mic-ρ-PCL for vascular wall construction, PHK26-labeled MSCs (10,000 cells/cm2) and GFP-labeled HUVECs (20,000 cells/ cm2) seeding on the outer and inner sides of the opened Mic-ρ-PCLbased tubular scaffold were cocultured in EGM10 for 5 days to obtain confluent cell-layers. After fixation, permeabilization, and blocking, the cells were incubated with DAPI for nucleus labeling. The construct of MSCs/PCL/HUVECs was characterized using CLSM via a z-scanning model. CLSM images were reconstructed using the built-in function of Imaris. To characterize the cells of the engineered vessel-wall construct, MSCs (10,000 cells/cm2) and GFP-labeled HUVECs (20,000 cells/ cm2) after coculturing in EGM10 medium for 5 days were fixed, permeabilized, blocked, and further incubated with TRITC-conjugated phalloidin (1:200 in PBS) for 60 min for F-actin labeling as previously reported.20 MSCs and HUVECs were imaged with CLSM using the TRITC and GFP channels, respectively. Immunocytochemistry Assays. The expression of vascular SMC differentiation marker (α-actin) in MSCs was analyzed via immunocytochemistry staining. MSCs (10,000 cells/cm2 on one side of Mic-ρPCL) were monocultured or cocultured with HUVECs (20,000 cells/ cm2 on the other side of Mic-ρ-PCL) in EGM10 for 3 days. Cells were fixed, permeabilized, blocked, and then incubated with the primary monoclonal antibody specifically targeted to SM α-actin (1:100 in 0.5% BSA) or the isotype IgG 2a k (1:100 in 0.5% BSA) at room temperature for 60 min. After washing thrice with PBS, cells were further incubated with the fluorescence-labeled second antibody (goat antimouse IgGAlexa Fluor 594; 1:500 in PBS) at room temperature for another 60 min. After washing thrice with PBS, the cells were finally incubated with DAPI for 5 min for nucleus visualization. Cells images were captured by CLSM using identical parameters for all groups. Statistical Analysis. Data analysis was performed on Prism 5 Software. Results were reported as the mean ± SD. A value of p < 0.05 was considered to be statistically significant.

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RESULTS AND DISCUSSION Direct Femtosecond Laser Microperforation of PCL Films. Thick HP-PCL was subjected to uniaxial thermal stretching to prepare a flexible and freestanding thin film of UX-PCL. As shown in Figure 2a,b, SEM images of the asfabricated UX-PCL exhibit a patterned topography comprising anisotropic microridges/grooves along the stretching direction. Such topographical structures could function as cues to elicit concurrent control over the alignment and myogenic differentiation of MSCs for vascular TE application.4,16 UX-PCL possessed surface micropatterns on a thin bulk of ∼38 μm (Figure 2b, insert), whereas it remained flexible and freestanding for engineering tubular vascular TE scaffold.4 Figure 2c,d shows SEM images of Mic-ρ-PCL with direct femtosecond laser microperforation at a laser fluence of 51.5 J/ cm2 and pulse number of 60. Mic-ρ-PCL exhibited hole arrays at adjustable hole placements. The holes showed precise positioning on the film’s surface where laser−material interaction occurred and have not caused structure disintegration of Mic1241

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Figure 3. SEM image of HP-PCL after direct femtosecond laser microperforation (red arrow, fibrous redeposit of sputtering PCL melt emanating from the ablated hole edge).

ridges/grooves of holes were confined within a range of ∼30 μm in radius from the center of the hole, without the formation of cracks around hole edges. However, it should be noted that the film integrity of Mic-ρ-PCL could be interfered with if the density and diameter of through-holes were high enough to cause adjacent hole−hole coalescence. In this study, the through-hole arrays were fabricated at a constant density of 400 holes/cm2 with an interhole distance of 500 μm for providing sufficient distance to avoid cellular bridging between adjacent twothrough-holes (cell length of elongated MSCs: ∼100−400 μm), thereby obtaining intact cell alignment along the microridges/grooves of Mic-ρ-PCL.29 Meanwhile, the diameter of through-holes was fabricated at ∼10 μm because at the same through-hole depth, smaller tunnel diameters could interfere with cellular ingrowth for inhibited myoendothelial interactions, whereas bigger tunnels typically larger than 10 μm would cause cell loss during the seeding and culture processes (the diameters of suspended MSCs and HUVECs in culturing medium: ∼15− 20 μm).29 Previously, a femtosecond laser has been applied to microperforate through-holes on metal (e.g., stainless steel),30 ceramic (e.g., silicon),31 and hard polymeric sheets (e.g., polydimethylsiloxane32 and poly(methyl methacrylate)).33 However, these materials were not flexible or biodegradable, thereby having little value for clinical translation in regenerating soft tissues including blood vessels. With the help of masks using stainless steel, we have investigated femtosecond laser ablation of bioresorbable PCL films with replication of through-hole arrays from the mask.28,34 This method, however, relied on the usage of photomasks, which limited the alterations of through-hole parameters (e.g., interhole distances) and could cause damage to and contaminate the film’s surface structures due to direct physical contact.35 Comparatively, this study first reported the direct femtosecond laser microperforation for PCL films, to obtain through-holes toward vascular TE applications. This method has not involved the use of photomasks and could be applicable to surface-patterned films. Toward vascular TE application, different leaching methods using salt,36 sugar,37 and/or water-soluble polymers such as poly(ethylene oxide)38 as the porogen have been extensively investigated for generating porous thin films. Such porogen leaching-induced holes, however, lacked through interconnectivity and were incapable of controlling spatial distribution leading to coalescence of the adjacent holes, thereby interfering with the film’s integrity properties and cellular organization response.39 Other methods such as needle-operated robot and

Figure 2. Morphologies of PCL films. (a) Large-area and (b) highly magnified SEM images of UX-PCL (double-headed arrow, the direction of stretching; insert, the cross-section of UX-PCL). (c,d) Large-area SEM images of Mic-ρ-PCL with precise hole placement (double-headed arrow, ridge direction). (e,f) Highly magnified SEM images of holes on the (e) top and (f) bottom sides of Mic-ρ-PCL (double-headed arrow, ridge direction; blue arrow, femtosecond laser-induced PCL labial melting structure; dashed line, evaginated features of PCL melt; red arrow, femtosecond laser-induced PCL fibrous melting structure).

ρ-PCL on nonprocessed surface areas, with an indication of retained film integrity. Such well-controlled direct microperforation of flexible and freestanding PCL thin films was accomplished principally because of two factors: (1) the low heat conductivity of polymeric biomaterials (∼0.1−0.3 W/(m·K)) that limited the diffusion of heat during laser−material interaction,25 and (2) the confined localization of the excitation energy with a short laser−material interaction time of femtoseconds.26,27 Furthermore, as shown in Figure 2e,f, SEM images of the top and bottom surfaces of Mic-ρ-PCL demonstrate that the hole was composed of a through-tunnel surrounded by a pair of labial melting structures along the ridges. Such geometric shape of the through-holes was reproducible in a large surface area of Mic-ρ-PCL as demonstrated from the low-magnified SEM images (Figure 2c,d). The labial features were consistent with what one expected due to the combined effects of thermal compressive stress imparted to the femtosecond-laser-ablated region28 and residual tensile stress from the prefabrication of uniaxial stretching.16 Both stresses tore the laser-induced PCL melt apart along the ridge’s direction and resulted in the formation of the labial features for Mic-ρ-PCL before the next laser pulse arrived. Consistent with this, due to the lack of uniaxial stretching, HP-PCL after experiencing the direct femtosecond laser ablation resulted in a fibrous redeposit emanating from hole edges, without the formation of labial melting features (Figure 3). Although fibrous redeposit of sputtering PCL melt was also found for Mic-ρ-PCL, the thermal damage of femtosecond-laserablation-induced sputtering melt to the surrounding micro1242

DOI: 10.1021/acsbiomaterials.5b00455 ACS Biomater. Sci. Eng. 2015, 1, 1239−1249

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Figure 4. Effects of laser fluence (F) and pulse number (Npulse) on the microperforation of PCL films. Direct femtosecond laser ablation was performed with different F at a consistent Npulse of 20 or with different Npulse at a consistent F of 51.5 J/cm2. (a) Light microscope images of the engineered throughholes (scale bar = 20 μm). (b) Quantitative analyses of the through-hole diameters on both the top and bottom surfaces of Mic-ρ-PCL (n = 5).

Figure 5. (a) Schematic diagram illustrating femtosecond laser intensity following the Gaussian distribution. Both (b) photochemical and (c) photothermal effects accounted for the formation of through-holes on PCL films (i, ii, and iii representing the temporal stages of a femtosecond laser pulse).

PCL thin films. It was found that with increasing F from 27.2 to 198.4 J/cm2, the laser generated greater labial features with more sputtering PCL melt. Accordingly, the diameters of throughholes on both the film’s top and bottom surfaces were found to increase. Theses results indicated that during direct femtosecond laser microperforation, F influenced the melting and the removal of PCL molecules from the material. The increase in throughhole size could be attributed to Gaussian distribution of the beam intensity of the femtosecond laser, which at a higher F exhibited a wider area at the center of the laser beam (exceeding the ablation threshold; Figure 5a). The influences of Npulse on mask-based femtosecond laser microperforation have been previously studied for biaxially stretched PCL films.28,34 However, with comparison to these reports, this study applied direct femtosecond laser microperforation without the use of masks, for investigating the effects of Npulse on the through-hole fabrication on a newly developed film of UX-PCL. As shown in Figure 4a,b, direct microperforation using a femtosecond laser at a few pulses lower than 40 tended to result in more fibrous sputtering PCL melt, whereas at a Npulse approaching 60 there was almost no fibrous sputtering melt around the through-holes. This observation indicated that multipulses of a femtosecond laser

lithography were advanced over porogen leaching in the control of hole distribution.28,34,40 However, after being placed in cellular culture medium, the needle-punched films exhibited incompetent permeability for hydrophilic mass diffusion because the induced flaps at the sites of punching were found to close.28 Although lithography could allow for the production of throughholes, this method was inflexible due to the use of photomasks and the involvement of a multifarious set of processes.16 In contrast, in this study, the developed direct microperforation using a femtosecond laser could overcome the limitations with current perforation techniques, by offering noncontact and programmable engineering of through-holes with controlled spatial distribution and consistent hole shapes over the film’s surface. Furthermore, this method, which has not involved chemical treatments (often required for porogen leaching and lithography with potential risk of residuals)16 and could be applicable to various functionalized surfaces with nano/micropatterns, represented a reliable approach to produce porously biomimetic films for vascular TE application. Influences of F and Npulse on Direct Femtosecond Laser Microperforation. As shown in Figure 4a,b, direct femtosecond laser ablation at different F and Npulse was carried out for 1243

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Figure 6. Mechanical properties of PCL films. (a) Tensile stress−strain curves of HP-PCL, UX-PCL, and Mic-ρ-PCL. Normalized (b) yield stress, (c) Young’s modulus, and (d) yield strain to those of HP-PCL (n = 3; ***, p < 0.001; **, p < 0.01; *, p < 0.05; and NS, p > 0.05).

breaking of PCL molecules (chemical bond energy of PCL: 3.62−8.28 eV).27 Furthermore, the previous studies have shown that photothermal ablation occurred when polymeric molecules absorbed laser pulses to raise temperature, melt material, and result in material evaporation (Figure 5c).42 The gaseous molecules further formed a plasma plume above the polymer bulk to continuously absorb the laser energy and finally expand to generate shockwaves. These shockwaves rapidly decayed into acoustic waves and propagated within the polymer bulk to eject the melted molecules out from the ablated area edges. In this study, the observed labial-like PCL melting structures and redeposited melting fibers indicated that the direct femtosecond laser microperforation of PCL films has occurred with photothermal ablation. Meanwhile, the evaginated features of PCL melting found over the labial-like structures at both film’s top and bottom sides (Figure 2e,f) were inconsistent with the shockwaveinduced melting ejecta during photothermal ablation. Therefore, it could be argued that during the direct microperforation using a femtosecond laser, both photochemical and photothermal ablation contributed to the formation of through-holes for PCL films. Mechanical Properties of PCL Films with Direct Femtosecond Laser Microperforation. As shown in Figure 6a, the tensile stress−strain curve of HP-PCL possessed a typical elastic deformation up to the proportional limit, followed with reduced loading stress and subsequently plastic deformation. However, uniaxial stretching resulted in a changed tensile stress− strain curve for UX-PCL without the presence of loading stress reduction after achieving the proportional limit. Quantitative analysis demonstrated that compared to HP-PCL, UX-PCL obtained mechanical enhancement in terms of yield stress (1.9x

were preferable for microperforating through-holes with better quality, which agreed with the effect of Npulse reported for the mask-based femtosecond laser microperforation of biaxially stretched PCL films and was probably attributed to accelerated concurrent disintegration of materials at a larger Npulse.28,34 Furthermore, the increase in Npulse from 20 to 60 showed little change to the dimensions of the labial-like melting structures or through-holes at both the film’s top and bottom sides. In the published studies, different mechanisms including thermal, mechanical, photophysical, photochemical, and defect have been established for describing the laser ablation of polymers.41,42 Comparing the characteristics of these models to the experimental conditions and observations on Mic-ρ-PCL, it was suggested that photochemical and photothermal ablation could contribute to the direct femtosecond laser microperforation of PCL films. On the basis of the published studies, photochemical ablation occurred when the incident photon energy was sufficiently high to elicit direct bond breaking.41,42 In this process, when a laser pulse duration was in the order of femtoseconds, the high laser peak power could result in significant optical nonlinearity and multiphoton absorption to induce conductive band electrons at the first part of a femtosecond laser pulse (Figure 5b).41,42 The ionized electrons could further absorb energy from the laser pulse’s later part and break free from the polymer bulk to form a strong electric field. The polymeric clusters disintegrated via Coulomb explosion when the magnitude of the electron energy was higher than the molecule’s binding energy.28 In this study, the femtosecond laser used was at a pulse duration of 110 fs with a wavelength of 800 nm and a repetition rate of 1000 Hz. The high laser peak power was able to induce multiphoton absorption and result in the bond 1244

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Figure 7. Hydrophilic permeability of Mic-ρ-PCL. (a) Diffusion of FITC-dextran across Mic-ρ-PCL as a function of incubation time in PBS at 37 °C (n = 5; UX-PCL, nonperforated film). (b) Diffusion of FBS across Mic-ρ-PCL as a nutritional/signaling biomolecule supplement for MSC adhesion and proliferation. MSCs (5,000 cells/cm2) were seeded on the PCL films of different groups and cultured for 1 and 3 days (D0, DMEM with 0% FBS; D10, DMEM with 10% FBS; Neg. Ctrl., negative control; Pos. Ctrl., positive control; n = 3; ***, p < 0.001; and NS, p > 0.05).

+, p < 0.001; Figure 6b) and Young’s modulus (0.7x + , p < 0.001; Figure 6c) without sacrificing the film’s yield strain (p > 0.05; Figure 6d). Comparing Mic-ρ-PCL and UX-PCL, both exhibited similar tensile stress−strain curves (Figure 6a) with comparable yield stress (p > 0.05; Figure 6b) and strain (p > 0.05; Figure 6d). This indicated that the direct femtosecond laser microperforation of this study has not changed the film’s elastic property. Furthermore, as compared to UX-PCL, Mic-ρ-PCL exhibited a declined Young’s modulus, suggesting that the incorporation of through-holes could make PCL films more flexible. On the basis of the Laplace’s law (burst strength = material’s yield stress· thickness/radius),43 vascular TE scaffolds fabricated from a singular Mic-ρ-PCL layer (∼38 μm in thickness) could be similar to those engineered from UX-PCL, to provide a sufficient theoretical burst strength (e.g., > 5000 mmHg for a scaffold at a diameter of 2.3−4.6 mm) to withstand the normal inside pulsatile pressure of the human mammary artery (∼3200 mmHg).4 These indicated that a femtosecond laser could be a suitable tool for microperforating biomaterials at a proper through-hole size and density in regenerating mechano-active tissues such as blood vessels, which require not only the reconstruction of the native-like complex architecture but also the mechanical retention for ensuring safety considerations. It should be noted that this study produced Mic-ρ-PCL at a constant through-hole density (400 hole/cm2) and diameter (∼10 μm) and thus has not investigated the influences of through-hole density and diameter on a film’s mechanical properties. However, previous studies have shown that film stiffness tended to decrease with the incorporation of holes,44,45 which agreed with the findings on Mic-ρ-PCL as compared to that of UX-PCL. In addition, the previous study showed that with increased hole diameter and density, the percentage surface area of holes increased and the film’s stiffness tended to decline.44 It should also be noted that after implantation, the mechanical properties of Mic-ρ-PCL might be altered during the degradation process. In this study, Mic-ρ-PCL could be considered as two portions: nonprocessed area and through-holes. The nonprocessed area comprising UX-PCL, based on the published study, exhibited preferential hydrolysis in superficial amorphous structure and could retain little change to the film’s yield stress up to experiencing ∼30% weight loss of the whole film.20

Meanwhile, the through-holes were covered with superficial laser-ablated layers, which have been reported with enhanced resistance against hydrolysis and could delay the degradation of the underlying structure.20 These indicated that Mic-ρ-PCL could have mechanical retention before undergoing a large weight loss (e.g., ∼30%) Transmural Biomolecule Diffusion and Physical MSCsHUVECs Contacts. The in-depth influences of direct femtosecond-laser-microperforated through-holes have been linked to the biological functions of Mic-ρ-PCL for myoendothelial communications such as those occurring in native blood vessels. As shown in Figure 7a, Mic-ρ-PCL showed a hydrophilic permeability with continuous transmural diffusion of FITCdextran (known as a molecular marker)21 over an investigated period of ∼12 days. This hydrophilic permeability could be attributed to the altered physicochemical properties of both increased roughness28,46 and oxygen gain27 occurring on the tunnel surfaces of through-holes. Consistently with this, there was no diffusion of FITC-dextran observed for the group of UXPCL. To further interrogate whether Mic-ρ-PCL could enable the transmural diffusion of nutrition and signaling biomolecules, a simulated milieu for the stromal cells in tunica media was built, which in the actual vessels was physically separated from blood. MSCs instead of SMCs were selected as the cell source of tunica media due to their more expandable capability in vitro, immune compatibility, and unique antithrombogenicity for vascular TE.4 Using FBS as the nutritional/signaling supplement, Figure 7b shows a proliferative capability for the cells cultured in Pos. Ctrl., Mic-ρ-PCL, and Mic-ρ-PCL/Cell-Layer groups. Compared to both the Neg. Ctrl. and UX-PCL groups, MSCs cultured on Micρ-PCL showed an improved cellular adhesion (1.7−1.9x + , p < 0.001) and proliferation (2.2−2.4x + , p < 0.001) with larger cell nucleus numbers found for 1 and 3 days of culturing, respectively. Such observations were in keeping with the anticipated FBS diffusion across Mic-ρ-PCL over to the cells. Furthermore, comparing the groups of Mic-ρ-PCL and Mic-ρ-PCL/CellLayer, MSCs exhibited comparable cellular adhesion and proliferation in both groups, with approximate cell nucleus numbers for 1 (p > 0.05) and 3 (p > 0.05) days of culturing, respectively. This result indicated that the presence of cell layers on the bottom side of Mic-ρ-PCL has not inhibited transmural 1245

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Figure 8. Cellular ingrowth into the through-holes of Mic-ρ-PCL. MSCs (5,000 cells/cm2) were cultured on Mic-ρ-PCL in D10 for 3 days. SEM images of cells exhibited various morphological responses such as (a) bridging, (b) sidewall adhesion, and (c) covering onto the tunnel surfaces of the throughholes (blue, anchoring of cell protrusions to the tunnel surface; red arrow, cell protrusions approaching the hole depth).

Figure 9. Transmural heterotypic cell−cell contacts via the through-holes of Mic-ρ-PCL. PHK-26-labeled MSCs (10,000 cells/cm2) and GFP-labeled HUVECs (20,000 cells/cm2) were seeded on the two sides of Mic-ρ-PCL and cocultured in EGM10 for 3 days. (a) 3D z-stacked and (b) cross-sectional views of CLSM images representing the structural support of Mic-ρ-PCL for physical contacts between MSCs and HUVECs (red, PHK-26-labeled MSCs; green, GFP-labeled HUVECs).

Figure 10. (a) Small-diameter vascular scaffold fabricated from Mic-ρ-PCL and a confocal microscope image of a hybrid vessel wall construction comprising of MSCs/PCL/HUVECs multilayers. (b) Confluent HUVECs showing a typical cobblestone-like morphology on the intima side of Mic-ρPCL. (c) Confluent MSCs exhibiting an aligned organization of cytoskeletal stress filaments on the media side of Mic-ρ-PCL. MSCs (10,000 cells/cm2) and HUVECs (20,000 cells/cm2) seeded on the two surfaces of the opened scaffold were cocultured in EGM10 for 5 days (blue, DAPI-labeled cell nuclei; green, GFP-labeled HUVECs; red, MSCs labeled with PHK-26 in panel a and F-actin in panel c; double-headed arrow, ridge direction).

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ACS Biomaterials Science & Engineering FBS diffusion, probably attributed to the semipermeability of the cell membrane and the not fully blocked through-tunnels by cells. These results on transmural FITC-dextran and FBS diffusion suggested that femtosecond laser microperforation offered Micρ-PCL with improved hydrophilic permeability, which could be for the diffusion of nutritional/signaling biomolecules between the TE tunica media and intima. Further studies demonstrated that the direct femtosecondlaser-ablated through-holes of Mic-ρ-PCL could enable cellular ingrowth. As shown in Figure 8a−c, SEM images reveal the different manners by which cellular protrusions approach the hole depth: bridging (cell bridging across the hole at two anchors; Figure 8a), sidewall adhesion (cell adhering to the steep hole-sidewall; Figure 8b), and covering (cells spread and cover the open hole via multiple anchors; Figure 8c). When coculturing MSCs and HUVECs (as a cell source of tunica intima)47 on the opposing sides of Mic-ρ-PCL for 3 days, CLSM images exhibit physically direct cellular contacts (colocalization represented in yellow; Figure 9a,b) between the PHK26-labeled MSCs (red) and GFP-labeled HUVECs (green). The mutual embedment of red and green suggested an interdigitation interaction between MSCs and HUVECs as observed from both 3D z-stacked and cross-sectional views. Such heterotypic cell−cell contacts across Mic-ρ-PCL were attributed to the through-holes, which facilitated the cell protrusion’s ingrowth. This heterocellular interaction between MSCs and HUVECs was akin to the myoendothelial contacts found between the stromal (e.g., SMCs8 and pericytes)7 and endothelial components in both developmental and mature vessels, and has been known to account for important vascular functions (e.g., signaling pathways, cellular myogenic differentiation, and nascent vessel maturation).9,10 Construction of Functionalized 3D Vascular TE Scaffolds. Finally, we applied Mic-ρ-PCL to engineer a 3D tubular vascular TE scaffold, with the spatial distribution of through-holes and circumferentially aligned ridges/grooves through rolling of the film around a rod and then attaching it via heat welding (∼3.5 mm in diameter; Figure 10a). With seeding of PHK26-labeled MSCs and GFP-labeled HUVECs on the outer and inner sides of the opened Mic-ρ-PCL-based scaffold, respectively, an engineered vascular wall construction of MSCs/PCL/HUVECs was achieved after 5 days of coculturing as shown from the z-stacked CLSM image. It was found that Micρ-PCL could act as structural support for the cells to adhere on its surfaces and simultaneously to separate them into distinct stromal and endothelial layers. On one side of Mic-ρ-PCL, HUVECs exhibited a typical cobblestone-like morphology with spreading and elongated cellular shapes and formed the TE tunica intima comprising a confluent singular cell-layer (Figure 10b). Meanwhile, MSCs on the opposite side of Mic-ρ-PCL appeared to organize into aligned cell multilayers such as that of the native tunica media, with cytoskeleton F-actin fibers oriented along the ridges (Figure 10c). To understand whether actual heterocellular interaction has occurred between the engineered stromal/endothelial components, the expression of a vascular SMCs differentiation marker (SM α-actin) was investigated in MSCs with 3 days of monoculturing (MSCs/Mic-ρ-PCL) or coculturing with HUVECs (MSCs/Mic-ρ-PCL/HUVECs). CLSM images show that there was no cellular red fluorescence found for the negative control (Figure 11a), suggesting that no nonspecific immune conjunction occurred. In contrast, with the labeling of SM αactin, both mono- and cocultured MSCs exhibited red

Figure 11. CLSM images of MSCs with the positive expression of vascular SM α-actin using immunocytochemistry staining. MSCs (10,000 cells/cm2; seeded on one side of Mic-ρ-PCL) were monocultured or cocultured with HUVECs (20,000 cells/cm2; seeded on the other side of Mic-ρ-PCL) in EGM10 for 3 days. (a) Negative control of IgG 2a. (b) Expression of SM α-actin in MSCs. SM α-actin representing the protein marker related to smooth muscle early term differentiation, and IgG 2a representing the isotype of SM α-actin (blue, nucleus DNA; red, investigated marker).

fluorescence, with an indication of positive expression (Figure 11b). Comparatively, CLSM images of the cocultured group exhibit red fluorescence around most cell nuclei of the MSCs, while in the monocultured group the red fluorescence of SM αactin was observed only around a few MSC nuclei, with lots of cells lacking the positive labeling. These observations suggested that the coculture of MSCs with HUVECs using Mic-ρ-PCL could promote myogenic differentiation in MSCs with elevated α-actin expression for vascular TE application. Such results agreed with those of the previous reports on direct MSCs/ HUVECs coculturing using a collagen matrix48a,b and demonstrated the occurrence of actual heterocellular interactions between the engineered stromal/endothelial layers. Previously, different strategies (e.g., electrospinning, soft lithography, and uniaxial stretching) have been developed to produce bioresponsive films with nano/micropatterned topographies for the architectural reconstruction of cell alignment in blood vessels.4,49 However, these films lacked structural support like that of the native vessels for the transportation of nutritional/ signaling biomolecules and physical cell−cell contacts between stromal and endothelial components,4,11,39,50 while the incorporation of hole structures on surface micropatterns (e.g., via porogen leaching)39 has been reported with interfered cellular alignment. Although various routes including microcarriers, collagen gels, and direct cellular coculture have been developed for achieving myoendothelial interactions, they were incapable of guiding stromal cell alignment or the formation of the correct vessel structure with distinct stromal and endothelial layers.29 In contrast, the developed Mic-ρ-PCL of this study exhibited micropatterned topographies and through-holes, and not only 1247

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obtained concurrent functionality for the vascular wall reconstruction comprising aligned stromal cell multilayers and a confluent endothelium but also enabled both transmural biomolecular diffusion and physical cell−cell contacts for myoendothelial interactions between the engineered stromal/ endothelial components. Such interactions based on Mic-ρ-PCL could influence MSCs in the engineered stromal layers with upregulated expression of SM α-actin. In addition to biological benefits, the developed PCL films with precise microperforation showed retained mechanical properties, which could ensure the engineered vascular TE scaffolds with sufficient theoretical burst strength for the safety considerations. In short, the developed film with direct femtosecond laser microperforation exhibited unique bio- and mechano-properties. The as-fabricated vascular TE scaffolds, in the incorporation of MSCs and HUVECs, could have potential for generating “off-the-shelf” grafts with the native-like complex architecture of blood vessels.

REFERENCES

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CONCLUSIONS Herein, we presented the direct femtosecond laser microperforation of surface-patterned PCL films for a functionalized vascular TE scaffold’s construction. Direct femtosecond laser ablation resulted in through-holes with precise placement for PCL films via photochemical and photothermal effects. Laser fluence and pulse number were found to influence both throughhole quality and dimensions. With direct femtosecond laser microperforation, the PCL films of this study displayed improved flexibility without sacrificing the film’s yield stress and showed hydrophilic permeability allowing for transmural biomolecule diffusion and physical cell−cell contacts, leading to MSCs with up-regulated expression of α-actin. Such a PCL film could be rolled into a small-diameter 3D tubular TE scaffold, and in the incorporation of MSCs and HUVECs, it facilitated a hybrid vascular wall construction comprising aligned stromal and confluent endothelium components similar to the architecture of the native vessels. This work offered a direct microperforation method using a femtosecond laser for bioresorbable and surfacepatterned films to obtain through-holes toward vascular TE application. The as-fabricated 3D tubular scaffolds could have the potential for vascular reconstruction with the native-like stromal and endothelial components.

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AUTHOR INFORMATION

Corresponding Authors

*(E.S.T.) Tel: 65 6516 5233. Fax: 65 6779 1459. E-mail: [email protected]. *(M.H.) Tel: 65 6516 1636. Fax: 65 6779 1103. E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We thank Dr. Jing Lim and Mr. Sandikin Dedy from National University of Singapore for their assistance in the FITC-dextran diffusion test and cell isolation for this project, respectively. This research is supported by the National Research Foundation, Prime Minister’s Office, Singapore (Competitive Research Program: NRF-CRP10-2012-04) and grant from the Ministry of Education, Singapore (R-265-000-300-112). J.K.Y.C. is supported by a salary from the Ministry of Health, Singapore (NMRC Clinician-Scientist Award: NMRC/CSA/043/2012). 1248

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